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PVP composite scaffold for cell culture
Journal Pre-proof Electrohydrodynamic jet 3D printing of PCL/PVP composite scaffold for cell culture Kai Li, Dazhi Wang, Kuipeng Zhao, Kedong Song, Junsheng Liang PII:
S0039-9140(20)30041-2
DOI:
https://doi.org/10.1016/j.talanta.2020.120750
Reference:
TAL 120750
To appear in:
Talanta
Received Date: 10 September 2019 Revised Date:
11 January 2020
Accepted Date: 14 January 2020
Please cite this article as: K. Li, D. Wang, K. Zhao, K. Song, J. Liang, Electrohydrodynamic jet 3D printing of PCL/PVP composite scaffold for cell culture, Talanta (2020), doi: https://doi.org/10.1016/ j.talanta.2020.120750. This is a PDF file of an article that has undergone enhancements after acceptance, such as the addition of a cover page and metadata, and formatting for readability, but it is not yet the definitive version of record. This version will undergo additional copyediting, typesetting and review before it is published in its final form, but we are providing this version to give early visibility of the article. Please note that, during the production process, errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain. © 2020 Published by Elsevier B.V.
Electrohydrodynamic jet 3D printing of PCL/PVP composite scaffold for cell culture Kai Li1, Dazhi Wang1, 2, *, Kuipeng Zhao1, Kedong Song3 and Junsheng Liang1, 2 1
Key Laboratory for Micro/Nano Technology and System of Liaoning Province, Dalian University of Technology, Dalian, 116024, China. 2 Key Laboratory for Precision and Non-traditional Machining Technology of Ministry of Education, Dalian University of Technology, Dalian, 116024, China. 3 State Key Laboratory of Fine Chemicals, Dalian R&D Center for Stem Cell and Tissue Engineering, Dalian University of Technology, Dalian 116024, China.
Abstract Controlled printing of biodegradable and bioresorbable polymers at desired 3D scaffold is of great importance for cell growth and tissue regeneration. In this work, a novel electrohydrodynamic jet 3D printing technology with the resultant effect of electrohydrodynamic force and thermal convection was developed, and its feasibility to fabricate controllable filament composite scaffolds was verified. This method introduces an effective thermal field under the needle to simultaneously enhance the ink viscosity, jetting morphology controllability and printing structure solidify. The fabrication mechanisms of thermal convection on jetting morphology and printed structures feature were investigated through theoretical analysis and experimental characterization. Under optimized conditions, a stable and finer jet was formed; then with the use of this jet, various 3D structures were directly printed at a high aspect ratio ~30. Furthermore, the PCL/PVP composite scaffolds with the controllable filament diameter (~10 µm) which is closed to living cells were printed. Cell culture experiments showed that the printed scaffolds had excellent cell biocompatibility and facilitated cellular proliferation in vitro. It is a great potential that the developed electrohydrodynamic jet 3D printing technology might provide a novel approach to directly print composite synthetic biopolymers into flexibly scale structures for tissue engineering applications.
Keywords: 3D printing; electrohydrodynamic jet; thermal field; composite scaffold; synthetic biopolymers; cell culture Corresponding author: Dazhi Wang. Email: +86(411)84707170-2171; Fax: +86(411)84707940 Supplementary Information is available online. 1
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1. Introduction Biopolymers can be derived from natural or synthetic materials and can take an important role in drug screening, organ repair and tissue engineering, which provide mass transport and temporary mechanical support to facilitate regenerated cells in lots of successful applications [1-4]. Compared with naturally derived polymers, synthetic biopolymers have the advantages of high biological compatibility, flexible controlling of the molecule and high function mimic of organs. And it can be biodegradable through hydrolysis or enzymatic degradation. In addition, the synthetic biopolymers with high tensile strength, high materials degradation rate and other unique properties can be created using chemistries such as photo-mediated degradation or crosslinking [5-8]. Hence, the synthetic biopolymer is now receiving increased attention because of their enormous need from regenerate damaged tissue to tissue culture [9-12]. Recently various synthetic biopolymers, such as polylactic acid (PLA) [13, 14], polycaprolactone (PCL) [15, 16], polyvinyl alcohol (PVA) [17, 18], polyvinyl pyrrolidone (PVP) [19, 20] and poly (lactic-co-glycolicacid) (PLGA) [21, 22], have been employed for tissue engineering and regenerative medicine aim to develop replacement tissues for our body. A number of biofabrication technologies enable the fabrication of biological constructs with precise control over the positioning of synthetic biopolymers. Unlike conventional biofabrication technologies that require moulds, masks and complex process, printing makes it possible to rapidly turn computer-aided designs into complex 3D objects on-demand [23-25]. The extrusion-based printing is one of the most commonly used biofabrication techniques which relies on “pushing” of high viscous synthetic biopolymers ink and its molten status. In this kind of technique, inkjet printing relies on the generation of droplets at the needle outlet, followed by 2
noncontact deposition on a substrate to fabricate synthetic biopolymers structures. The droplets jet from the needle due to the action of pressure pulses generated, which induced by thermal or piezoelectric means. It is obvious that a challenge for inkjet printing is the achievable levels of resolution, establishing minimum sizes of patterned features about 10 µm [26]. In addition, fused deposition modeling (FDM) and direct ink writing (DIW) extrudes the viscous state ink in a steady flow until deposition and must rapidly stabilize upon delivery [27-30]. These “pushing” printing techniques are simple and low cost. However, the resolution of extrusion-based printing technologies has a dependence on the needle inner size. The needle inner size used usually hundreds of micrometers to maintain the steady flow due to the high viscosity of printing ink. Thus, the feature size of the printed 3D structure is also larger than the needle size, which restricts application scope and device performance of synthetic biopolymer. Electrohydrodynamic jet printing (E-Jet) was thus developed to fabricate synthetic biopolymer patterning by controlling the movement of the collecting substrate because of its high-resolution characteristic [31, 32]. In the E-Jet printing process, the mobile ions in the synthetic biopolymers ink are induced by applying a high voltage potential and accumulate near the surface at the needle tip, and this leads to an intense electric field force around the pendent drop. E-Jet printing utilizes this electric field force to induce synthetic biopolymer ink forming a fine jetting from needle tip, which is based on “pulling” force rather than “pushing” force, exhibiting unique advantages of high resolution (sub-100 nm) [26]. Previous works have been demonstrated the compatibility of E-Jet printing with a wide variety of synthetic biopolymers inks, ranging from P(S-ran-MMA), PCL, PVA and many others. Various structures of dot, line and film with a feature size from hundreds of nanometer to micrometer have been printed [33-36]. Lee et al developed a metal
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electrode assisted E-Jet printing process and the printed polyethylene oxide (PEO) filaments could be precisely stacked on a platinum line layer-by-layer deposition [37]. Wang et al integrated an air infusion modulus into the E-Jet printing system to increase the rapid solidification of printed PCL filaments [38]. Choe et al proposed a modified E-Jet printing technology to fabricate the micro/nanoscale hybrid scaffold with a controllable pore using the cellulose acetate biomaterial [39]. However, the previous E-Jet printing technologies are often used to print single synthetic biopolymers. And it is difficult to build high aspect ratio composite 3D structures because the low viscous synthetic biopolymers ink is likely to spread on the substrate and the printed structure is easy to collapse. In this work, a new E-Jet 3D printing technique under the resultant effects of electrohydrodynamic and thermal convection was developed to print PCL/PVP composite scaffold for cell culture. Firstly, a novel E-Jet 3D printing equipment with a thermal platform under the needle was designed and prepared for printing in which the electric field was applied on the needle and the thermal field was supplied by the thermal platform. A low viscous composite synthetic biopolymer ink (PCL/PVP) was prepared after their particles were dissolved in low boiling point solvent. Then, the electrohydrodynamic effect provided an electrical shearing force to induce a fine jet and the thermal convection was employed to increase the viscosity of the ink by evaporating the low boiling point solvent during printing. Eventually, the stable and controllable fine jet was pulled out and formed at the needle tip. It should be noted that atomization could be likely to happen if the electrohydrodynamic effect was only applied to this low viscous biopolymer ink. Even though the jet can be formed it is impossible to pattern 3D structures because of the low viscosity of the ink. And electrospinning normally occurs when a high viscous ink was directly used for 4
printing under high electrical field, causing the jet whipping, which resulting in difficult to control of printing. The resultant effects of thermal field and electric field on jetting behaviors and size of printed 3D structures were deeply investigated through experimental characterization. By optimizing the parameters of printing process, various 3D structures with a high aspect ratio (~ 30) were efficiently created. Furthermore, 3D PCL/PVP composite scaffolds cultured with Murine MC3T3-E1 Subclone14 cartilage cells were printed flexibly as a platform to study cell regeneration in vitro. Excellent cell compatibility of printed 3D scaffolds promoted cell growth and proliferation. Higher cell density and filling the internal pores of the scaffolds under 5 days of culture were observed compared to 3 days. Overall, this new E-Jet printing technique develops a promising way to fabricate high aspect ratio 3D structure and provides a great potential for tissue engineering applications.
2. Materials and methods 2.1 Experimental set-up The schematic illustration of printing set-up is shown in figure 1a, including an electrohydrodynamic needle, a syringe pump, a high voltage power supply, a thermal platform, a computer-controlled X-Y-Z movement stage and a camera. The electrohydrodynamic needle is a quartz needle with an inner/outer diameter of 100 µm/380 µm (New Objective, Inc, USA). The quartz needle is installed in a channel of metallic fixture. And the metallic fixture is connected to the output of high voltage power supply (Physical Instruments, Germany). In this condition, the high voltage could be applied to the ink through this metallic fixture. The inlet of metallic fixture is connected to the glass syringe (100 µL, Hamilton, USA) by a hose to feed the low viscous biopolymer composite ink. The glass syringe was placed in a syringe pump
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(Harvard Apparatus, Holliston, USA). The thermal platform with a resistance heater is fixed on X-Y-Z movement stage (uKSA, Zolix Instruments Co., Ltd, China) to provide thermal field during printing. A temperature-control device is used for ensuring the heating uniformity of thermal platform. A control program is developed based on LabVIEW, which is employed to coordinate the movement of X-Y-Z motion stage. A camera (Point Grey, Canada) is used for observing the jetting behaviors. 2.2 Preparation of printable PCL/PVP composite ink PCL ink and PVP ink were prepared separately and two inks were mixed to form the PCL/PVP composite ink. PCL was not only a good biologically compatible synthetic biopolymer but also insoluble in water and alcohol. Therefore, the PCL and PVP biopolymer were chosen as the scaffold material while PVP was added to modulate the viscosity of the printable ink. 9.1 wt.% of PCL particle (Mw=800,000, yuanye Bio-Technology Co., Ltd, China) was dissolved into acetic acid (Adamas Reagent Co., Ltd, China) and mixed for 50 minutes at 80 °C. The preparation of PCL ink was completed. 11.7 wt.% of PVP powder (Mw=1,300,000, Aladdin Industrial Corporation, China) was dissolved into ethanol. And 0.43 wt.% of PEG (Mw=800, Kemiou Chemical Reagent Co., Ltd, China) was added to the solution and mixed for 5 minutes at 50 °C. After the mixed solution was cooled to 25 °C, 1.7 wt.% of Triton X-100 (Kemiou Chemical Reagent Co., Ltd, China) was added to the solution and mixed for 10 minutes, the PVP ink can be obtained. Finally, the PCL ink and PVP ink were mixed for 20 minutes at 25 °C to form printable PCL/PVP composite ink (figure S1a). As shown in figure S1a, the surface tension of PCL/PVP ink was measured by an automatic surface tensiometer (K100C, Kruess Scientific Instrument, Germany). The measurement resolution is 0.01 mN m-1 and the measurement range is 1-500 mN m-1. The surface tension of the PCL/PVP composite ink was tested ten times and its 6
mean value was calculated. 2.3 3D printing scaffold process The low viscous PCL/PVP composite biopolymer ink with low boiling point was placed in the glass syringe and infused to the quartz needle through the syringe pump which provides the hydrodynamic force to push PCL/PVP ink at a constant flow rate. Subsequently, an applicable temperature of thermal platform was selected to increase the viscosity of composite ink. Afterwards, the viscosity increased with the solvent evaporation under the thermal field (figure 1b). And the jetting formed under the action of high electric shearing force provided by a high voltage power supply. The stable and controllable finer jet was pulled out and formed at the needle tip. Finally, the PCL/PVP composite scaffolds were overlaid layer by layer through the X-Y-Z movement stage controlled by a LabVIEW program. In addition, the camera was used to observe the printing process in real-time. 2.4 Cell culture on printed scaffolds The Murine MC3T3-E1 Subclone14 cartilage cells were cultured in α-Dulbecco's Modified Eagle's Medium (α-DMEM, Hyclone, USA) containing 10 % v/v fetal bovine serum (FBS, MINHAI, China), 100 units mL-1 penicillin and 100 µg mL-1 streptomycin (Hyclone, USA). Before cell seeding, the printed PCL/PVP scaffolds were sterilized in 75 v/v% alcohol for 5 h. Subsequently, 20 µl cell suspensions with a density of 5×105 µg mL-1 were drop-seeded on both sides of each scaffold and incubated for 3 h at 37 °C and 5 % v/v CO2 atmosphere, followed by supplement of fresh medium. 2.5 Evaluation of cell growth on composite scaffolds On 3 days and 5 days after cell seeding, Dead/Live staining was performed to evaluate cell viability and attachment. Calcein-AM, propidium iodide (PI) and 7
Hoechst 33258 were purchased from Calbiochem (San Diego, CA, USA). Briefly, cell-scaffold constructs were incubated in PBS containing 2 µmol L-1 calcein-AM, 4 µmol L-1 propidiumiodide (PI) and 5 µg L-1 Hoechst 33258 staining solution at 37 °C for 30 min. Cell viability on the PCL/PVP composite scaffolds was observed through four fluorescent images which were randomly taken using a fluorescence microscope (IX71 + DP71, Olympus, Japan). The scale in each confocal micrograph was saved in Image-Pro Plus 6 (IPP, Media Cybernetics, USA) software and set as the benchmark. The normalized cell densities (the initial cell numbers per mm2 are normalized to the level of those in the composite scaffolds) are measured within 5 days. The cell viabilities are estimated after culturing for 3 days and 5 days. Furthermore, the cells on the PCL/PVP scaffolds were visualized using field emission scanning electron microscopy (FE-SEM, SU8220, HITACHI, Japan). For FE-SEM observing, the cell-scaffold constructs were fixed with 2.5% glutaraldehyde at 4 °C for 3 h and gradually dehydrated in an ascending series of ethanol (50%, 70%, 90% and 100%) for 30 min for each solution. And these FE-SEM samples were dried in a vacuum condition. After sputter-coated with gold, the samples were observed by FE-SEM with an accelerating voltage of 5.0 kV.
3. Results and discussion 3.1 Effect of thermal field and electric field on jetting behaviors and feature of printed 3D structures The temperature of thermal platform has a great effect on printing behaviors and printed 3D structure features. In order to obtain a stable jet and desired composite scaffold, the effect of thermal platform on the printing mechanisms were examined. In this paper, the thermal platform (as a heat source) under the needle is made by a square aluminum plate. Therefore, the thermal convection during printing is the heat 8
dissipation from a thermal platform surface to its upper PCL/PVP composite ink. And thermal convection takes place in the composite ink through collisions or oscillations of each molecule. The equations used for simulation analysis are shown in the Supplementary Information. Figure 2 shows the simulated results of temperature distribution between the thermal platform and needle tip using COMSOL Multiphysics software. When the temperatures of thermal platform were 30 °C (figure 2a), 35 °C (figure 2b) and 45 °C (figure 2c), the temperatures of needle tip were 27 °C, 32 °C and 38 °C, respectively. And the temperature of needle rose to 47 °C when the temperature of thermal platform increased at 55 °C (figure 2d-e). The thermal convection near the needle tip is more intense when the temperature of thermal platform increased. And this thermal convection was employed in evaporating the solvent in PCL/PVP composite ink to increase its viscosity during printing. Figure 2f shows the simulated temperature contrasting of thermal platform and needle tip. In the PCL/PVP composite ink, the boiling points of two most solvents (ethanol and acetic acid) were 78 °C and 117.9 °C. Furthermore, these two solvents were volatile. It is important for a suitable temperature of thermal platform to evaporate volatile solvents and obtaining viscous ink. And a stable and controllable fine jet was pulled out and formed at the needle tip. Figure 3a-d show the jetting behavior affected by the temperature of thermal platform, ranging from 30 °C to 55 °C. The applied voltage, flow rate, printing height and speed of movement stage were at 650 V, 0.025 µL min-1, 150 µm and 2 mm s-1, respectively. Figure 3a-b shows that the Taylor cone was formed and the PCL/PVP biopolymer droplet was ejected from the cone tip when the temperature of thermal platform was lower than 35 °C. This is because the temperature of needle tip (32 °C showing in figure 2b) is far below the boiling point of two solvents and most solvents 9
have still stayed in the jetting. When the temperature of thermal platform rose to 45 °C, the temperature of the needle tip was 38 °C. In this printing condition, the acceleration of solvent evaporation (mainly ethanol) and increase of the viscosity of PCL/PVP ink occurred. The jetting behaviors of droplets along with jet at the tip of the cone were formed (figure 3c). However, the temperature was not high enough to form a semisolid jet and maintain the continuous jet. Figure 3d shows the stable semisolid jet was formed when the temperature of thermal platform rose to 55 °C. At this temperature the two volatile solvents (ethanol and acetic acid) were evaporated rapidly, then the viscosity of the PCL/PVP ink at the outlet of needle was further increased, resulting in the semisolid state. And it was viscous enough to support the formation of a stable continuous semisolid jet. It was suggested that the thermal platform temperature of 55 °C could significantly promote the evaporation of solvents and rising its viscosity. Figure 3e-h show the 3D profiles of printed PCL/PVP structures with 20 layers when the temperatures of thermal platform at 30 °C, 35 °C, 45 °C and 55 °C, respectively. Figure 3e shows the width and height of the printed PCL/PVP structure were 229±10.1 µm and 11±4.2 µm. And the aspect ratio of PCL/PVP structure was small. This was because the viscosity of PCL/PVP ink at the outlet of needle was low which would lead to ink dispersed on the substrate after printed. In contrast, as shown in figure 3e and 3h, the width of printed 3D PCL/PVP structures gradually decreased from 229±10.1 µm to 13±4.1 µm while the height increased from 11±4.2 µm to 110±3.1 µm as the temperature of thermal platform increased from 30 °C to 55 °C. Five positions were randomly selected on printed PCL/PVP structure and measured its sizes. The statistical analysis in figures 3i shows the increase of temperature could help to print high aspect ratio 3D structures. Therefore, the optimized temperature of 10
thermal platform was 55 °C. Figure 4 shows the effect of different applied voltages on jetting behaviors and size of printed 3D structures. In E-Jet printing process, the applied voltage plays a key role in controlling jet behaviors. And it affects the jet diameter and the aspect ratio of printed 3D structures simultaneously. During printing, the surface charge of ink increased continually and a meniscus was induced when the voltage was applied. Then the ink meniscus was deformed and the jetting was formed at the end of needle. In general, in order to overcome the surface tension of ink meniscus and forming a jetting, the charge on the meniscus is satisfied with the Rayleigh limiting charge [40],
Qmax = π(8ε0Fst dd3 )1/2 where Qmax is the maximum amount of carrying charge on the drop surface, ε0 is the ink vacuum permittivity, Fst is the ink’s surface tension, dd is the diameter of the droplet. In this work, the mean surface tension of PCL/PVP composite ink is 23.48 mN m-1 at 20 °C (ten measured data showing in figure S2). And equation 11 suggests that the ink with low surface tension just requires a lower voltage to induce the jetting. As shown in figure 4a, a jet formed at the end of needle when a voltage of 400 V was applied. When the applied voltage was increased, the jet diameter was decreased (figure 4a-c). This thin jet can help to print uniform 3D structures with a high respect ratio. It can be seen from the experiment process that multi-jet was generated when 1200 V was applied (figure 4d), which is an undesirable jetting behavior and cannot be used in precise printing 3D structures. From figure 4a to 4d, the angle (θ) between the jet and the direction of printed structures (horizontal direction) increased from 34.31° to 84.47° when the applied voltage increased from 400 V to 1000 V. It can be seen the applied voltage has a 11
marked impact on the orientation of jetting filament. When the applied voltage was lower (400 V, figure 4a), the drawing force (Fd) is the leading force to pull the jetting filament. And the drawing force promotes the stacking of filaments on 3D structures layer-by-layer. The orientation of jetting filament is more curved and the θ is small (34.31°, figure 4a) because of the effect of the drawing force. When the applied voltage increased the jetting rate (U) and the electric field force (Fe) increased. Thus the effect of Fd decreases while the effect of Fe increases and the angle (θ) increases at the same time. Figure 4e-h shows 3D profiles of printed PCL/PVP structures with 20 layers when the applied voltage was 400 V, 600 V, 800 V and 1200 V, respectively. As shown in figure 4e-g, the 3D structure width gradually decreased from 15±6 µm to 8±1.5 µm while the height increased from 51±4 µm to 86±3 µm. However, the multi-jet occurred when the applied voltage rose to 1200 V (figure 4d) because the normal electrical field force was much larger than the surface tension. In this condition, the cone-jet mode was destroyed and the flat meniscus with small cones formed around the end of needle. And tiny jets were ejected from these small cones and arrived at the substrate (figure 4h). Five positions were randomly selected on the printed PCL/PVP structure and measured its sizes. The statistical analysis of the effect of applied voltage on 3D structures size is showed in figure 4i. It can be seen that higher aspect ratio 3D structures can be obtained when the applied voltages increased up to 1200 V. 3.2 Printing 3D structures with high aspect ratio Figure 5a shows SEM image of 3D lattice structure programmed printed layer-by-layer. The size of every square lattice was 200 µm × 200 µm. During the printing process the temperature of thermal platform, applied voltage, flow rate,
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printing height and speed of movement stage were at 55 °C, 600 V, 0.03 µL min−1, 150 µm and 2 mm s-1, respectively. After 15 layers printing, the width of the printed 3D lattice was about 11 µm and the height was about 80 µm, the aspect ratio was about 7. The 3D profile in figure 5b shows the printed 3D lattice had the uniform and smooth features. At the condition of low applied voltage (600 V), small distance (150 µm) and low printing speed (2 mm s-1), the misplacing has not appeared in layer upon layer and the jet was kept vertical throughout the printing process. With the movement stage went back and forth at a speed of 2 mm s-1 in X-axial direction, a 3D structure consisting of 40-layers polyline were printed and the filaments were stacked orderly together on the substrate. As showed in figure 5c, the aspect ratio of printed 3D perpendicular walls was about 25 and each layer was completely overlapped. During the printing process the temperature of thermal platform, flow rate and speed of movement stage were at 55 °C, 0.03 µL min-1 and 2 mm s-1, respectively. Furthermore, in order to maintain an effective electric field force, the applied voltage increased (600 V-800 V) as the height of printed 3D perpendicular walls increased. These various 3D structures proved that this printing method is capable of constructing 3D microstructures with high aspect ratio. 3.3 Printing and evaluation of PCL/PVP composite scaffold The meaning of bioactivity is a biomaterial that has the ability to promote the formation of hydroxyapatite on its surface, while the biomaterial immersed in simulated body fluid. Previous studies have been proved that PCL scaffold has poor bioactivity probably due to the bio-inertness and slow degradation nature [14]. Thus 13
the incorporation of PVP into printed PCL microscale filaments would effectively promote the bioactivity of PCL/PVP composite scaffold. Furthermore, the PVP has higher mechanical stiffness and modulus. The incorporation of PVP into PCL microscale filaments would significantly increase the mechanical properties of the composite scaffold. In order to determine the existence of PCL and PVP in the printed composite scaffold, the PCL scaffold, PVP scaffold and the PCL/PVP composite scaffold were printed respectively. Then the FTIR spectra of these three scaffolds were performed. As shown in the red line of figure 6, the PCL scaffold spectrum exhibited the peak of C=O and stretching vibrations at 1702 cm-1, CH2 bending modes at 1359 cm-1, 1408 cm-1 and 1438 cm-1 and COO– vibrations at 1286 cm-1. The C–O– C stretching vibration peaks at 1221 cm-1, 1161 cm-1 and 1106 cm-1 [41]. The PVP scaffold spectrum (blue line) exhibited the intense sharp peak of C=O stretching vibrations of PVP at 1663 cm-1. The CH2 bending modes were observed at 1378 cm-1, 1424 cm-1, 1458 cm-1 and 1487 cm-1 and that of the C–N vibrations occurred at 1045 cm-1 [42]. The spectrum of PCL/PVP composite scaffold exhibited the dominant peak of both PCL (C=O stretch 1702 cm-1) and PVP (C=O stretch at 1663 cm-1). Moreover, it was found the COO– stretch of PCL at 1286 cm-1 and C–N stretch of PVP at 1045 cm-1. These results indicated the PCL/PVP composite scaffold caused almost no change in the FTIR spectra of each characteristic peaks, indicating no chemical bonding or physical interaction occurs between PCL and PVP components. The internal microstructure of tissue engineered 3D scaffolds plays an important role to provide a proper growth microenvironment for improving cell growth and proliferation. Previous studies have proved that the appropriate filament size and its spacing in the scaffold not only promotes cell seeding but also provides an effective nutrient exchange [7, 33]. As shown in figure 7a and figure 8a, the 3D porous 14
PCL/PVP composite scaffolds were fabricated on a glass substrate (Supplementary Video S1). During the printing process the temperature of thermal platform, flow rate, printing height and speed of movement stage were at 55 °C, 0.05 µL min-1, 200 µm and 8 mm s-1, respectively. It was printed for 8 layers, one layer was printed as X direction, then the adjacent layer was printed as Y direction (Supplementary Video S2). With the increase of the number of printing layers, the distance between the needle and the substrate increases gradually. Meanwhile, the applied voltage was slightly adjusted (600-700 V) in order to keep electric field intensity between the needle and ground electrode to obtain the stable printing process. The average distance between two adjacent filaments was ~80 µm. It also can be seen that the distance between the two adjacent filaments was a small difference, which was because that the previously printed filament can attract the printing jet when the printing height is small, thus resulting in the position error between the two adjacent filaments. The average width of the scaffold filaments was about 10 µm. And this filament size is closed to living cells and is suitable to provide sufficient surface area for cell survival. Furthermore, the scaffold size can be flexibly fabricated to mimic cell growth microenvironment, which can provide desired structures for improving cell growth and proliferation. Dead/Live staining of the 3D microtissue of Murine MC3T3-E1 Subclone14 cartilage cells grown in printed PCL/PVP composite scaffold was conducted 3 days and 5 days of culture. As shown in figure 7 and figure 8, the LSCM fluorescent images of live/dead-stained on the two scaffolds exhibit a high cell seeding percentage and a high cell survival rate. An overwhelming proportion of living cells were positively stained by calcein-AM with strong green fluorescence signals. The cell density in the scaffolds under 5 days of culture (figure 8) was apparently higher 15
than that of the scaffold under 3 days of culture (figure 7). And the cells under 5 days of culture were found to significantly proliferate and fill the internal pores of the scaffolds (figure 8e). As shown in figure 8c (5 days of culture), the increased number of dead cells (a very small proportion) stained by PI was inevitable and acceptable over time, possibly due to the lack of available nutrients and undischarged waste in the interior of printed scaffolds. After the PCL/PVP composite scaffolds were sterilized in 75 % ethanol aqueous solution for 5 h, the surface roughness of the printed filaments increased due to the partial dissolution of PVP components. And the overall PCL/PVP composite scaffold and its filaments could be stably maintained. Nanoscale pores exposed on the surface of the printed filaments could be promoted cellular attachment. [43]. Furthermore, 3D view images as shown in figure 7f and figure 8f displayed that the cells attached from the surface to the inside of the scaffold and the depth of the ingrowth of the cells reached about 25 µm. As shown in figure 9a, the normalized Murine MC3T3-E1 Subclone14 cartilage cells densities in PCL/PVP composite scaffold are measured within 5 days. The results indicate that the cell densities increased as the days of cell culture increased. And the number of cells has been more than four times within 5 days in composite scaffolds. The cell viability of cells has been measured according to the results of live/dead stain (figure 9b). The cell viability has been evaluated by calculating the ratio of number of live cells to number of total cells. The cell viability after cell culture of 3 days and 5 days are 91 %±2.1 % and 95 %±3.5%. In addition, after cell culture on the printed composite scaffolds for 3 days and 5 days, SEM observation was performed to examine the proliferation and morphology of cells. As shown in figure 9(c-f), the cells with the long pseudopodia attached well on the channel surface of composite scaffolds. At cell culture of 5 days, cell density on the composite 16
scaffolds was obviously higher than that at 3 days and the most area of the surface of scaffolds was covered with cells. These results were consistent with the results of LSCM fluorescent shown in figure 7 and figure 8. In a word, these results indicated favorable viability of cells, suggesting good cell compatibility of the printed PCL/PVP composite scaffolds. This novel E-Jet 3D printing technique proposed in this paper demonstrates a new and promising way to pattering high molecular weight biopolymer for tissue engineering. 4. Conclusion A novel E-Jet 3D printing technique, with the unique advantages of high aspect ratio feature and large controllability, is introduced to direct print microscale synthetic biopolymers composite scaffold for cell culture. The translation of synthetic biopolymers composite ink from low viscous to high viscous was realized with the thermal field effect and the local solution evaporated in printing process. The external electric field force created by a constant voltage induced this high viscous ink to form a focused and stable jet. Under the resultant effect of electrical shearing force and thermal field effect, the controllable fine filament was ejected from the needle tip. The jetting behavior of jetting filament and printed 3D features have been studied at different experimental parameters and the optimized printing condition was selected to fabricate high aspect ratio 3D structures layer-by-layer. Microscale PVL/PVP composite 3D scaffolds with excellent cell compatibility were printed for tissue growth and proliferation. And the cell culture experiments proved that with the increase of culture time, higher cell density was also observed on composite 3D scaffolds. The new E-Jet 3D printing technique presented in this paper demonstrates a promising way to print composite synthetic biopolymers for tissue engineering.
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Acknowledgment This research was supported by National Natural Science Foundation of China (51975104), National Key R&D Program of China (2018YFA0703200), the aerospace science foundation (2018ZD63004), the Fundamental Research Funds for the Central Universities (DUT18LAB17), State Key Laboratory of Precision Measuring Technology and Instruments (Tianjin University) (pilab1804) Science Fund for Creative Research Groups of NSFC (51321004) and the Collaborative Innovation Center of Major Machine Manufacturing in Liaoning. Reference [1] D. F. Williams, On the nature of biomaterials, Biomaterials 30 (2009) 5897-5909. [2] R. Parenteau-Bareil, R. Gauvin, F. Berthod, Collagen-Based Biomaterials for Tissue Engineering Applications, Materials 3 (2010) 1863-1887. [3] M. J. Webber, E. A. Appel, E. W. Meijer, R. Langer, Supramolecular biomaterials, Nature Materials 15 (2016) 13-26. [4] M. Hospodiuk, M. Dey, D. Sosnoski, I. T. Ozbolat, The bioink: A comprehensive review on bioprintable materials, Biotechnology Advances 35 (2017) 217-239. [5] E. S. Place, J. H. George, C. K. Williams, M. M. Stevens, Synthetic polymer scaffolds for tissue engineering, Chemical Society Reviews 38 (2009) 1139-1151. [6] D. F. Williams, On the nature of biomaterials, Biomaterials 30 (2009) 5897-5909. [7] T. G. Kim, H. Shin, D. W. Lim, Biomimetic scaffolds for tissue engineering, Advanced Functional Materials 22 (2012) 2446-2468. [8] Y. L. Kong, M. K. Gupta, B. N. Johnson, M. C. McAlpine, 3D Printed bionic nanodevices, Nano Today 11 (2016) 330-350. [9] S. Bose, D. Ke, H. Sahasrabudhe, A. Bandyopadhyay, Additive manufacturing of biomaterials, Progress in Materials Science 93 (2018) 45-111. [10] A. Sorushanova, L. M. Delgado, Z. Wu, N. Shologu, A. Kshirsagar, R. Raghunath, A. M. Mullen, Y. Bayon, A. Pandit, M. Raghunath, D. I. Zeugolis, The collagen suprafamily: from biosynthesis to advanced biomaterial development, Advanced Materials 31 (2019) 1801651. [11] S. V. Murphy, A. Atala, 3D bioprinting of tissues and organs, Nature
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Figure captions Figure 1. (a) Schematic diagram showing the E-Jet printing process; (b) The solvent evaporation during printing. Figure 2. Simulations of the temperature distribution at different temperatures of thermal platform; (a-d) Simulation images of thermal platform at 30 °C, 35 °C, 45 °C and 55 °C, respectively; (e) Showing the red frame in (d); (f) Simulated temperature relationship of thermal platform and needle tip. Figure 3. Experimental 3D printing at different temperatures of thermal platform; (a-d) The jet behaviors at 30 °C, 35 °C, 45 °C and 55 °C, respectively (scale bar=50 µm); (e-h) 3D profiles of printed structure at 30 °C, 35 °C, 45 °C and 55 °C, respectively; (i) Statistical graph showing the sizes of the printed 3D structures at different temperatures of thermal platform. Figure 4. Experimental 3D printing at different applied voltages; (a-d) The jet behaviors at 400 V, 800 V, 1000 V and 1200 V, respectively (scale bar=50 µm, θ meaning the included angle between jetting filament and the direction of printed structures); (e-h) 3D profiles of printed structure at 400 V, 600 V, 800 V and 1200 V, respectively; (i) The statistical graph showing the sizes of the printed 3D structures at different applied voltages. Figure 5. (a-b) SEM image and 3D profile of 3D lattice structure; (c) 3D profile of high aspect ratio walls. Figure 6. (a) Fourier transform infrared (FTIR) spectra for PCL scaffold, PVP scaffold, PCL/PVP composite scaffold, respectively; (b) The image of the boxed region in (a) and its characteristic peaks. Figure 7. Printing 3D porous PCL/PVP scaffold and evaluation of Murine MC3T3-E1 22
Subclone14 cartilage cells activity and distribution after 3 days culture. (a) The printed PCL/PVP composite scaffold; (b-f) Live/dead staining of cells on the scaffold; Live cells were stained with calcein-AM, dead cells were stained with PI and cell nuclei were stained Hoechst 33258, showing green, red and blue, respectively. Scale bars in a-e=200 µm, f=100 µm. Figure 8. Printing 3D porous PCL/PVP scaffold and evaluation of Murine MC3T3-E1 Subclone14 cartilage cells activity and distribution after 5 days culture. (a) The printed PCL/PVP composite scaffold; (b-f) Live/dead staining of cells on the scaffold; Scale bars in a-e=200 µm, f=100 µm. Figure 9. (a) The normalized densities of Murine MC3T3-E1 Subclone14 cartilage cells inside PCL/PVP composite scaffolds; (b) Viability of Hela cells inside composite scaffolds. ** indicate statistical significance of P < 0.01 (n=4). (c-f) The morphology of cells on the printed PCL/PVP scaffolds. SEM observations in (c-d) and (e-f) were performed to examine the morphology of cartilage cells after cells seeded on the scaffolds for 3 days and 5 days. Scale bars in c-f=100 µm.
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Highlights A novel E-Jet 3D printing method was developed for patterning synthetic biopolymers. Various 3D biopolymers structures were directly printed at a high aspect ratio ~30. 3D printed PVP/PCL composite scaffolds with controllable filament were developed. The scaffolds could act as favorable synthetic ECM for cartilage regeneration.
Declaration of interests ☒ The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. ☐The authors declare the following financial interests/personal relationships which may be considered as potential competing interests:
S0039-9140(20)30041-2
DOI:
https://doi.org/10.1016/j.talanta.2020.120750
Reference:
TAL 120750
To appear in:
Talanta
Received Date: 10 September 2019 Revised Date:
11 January 2020
Accepted Date: 14 January 2020
Please cite this article as: K. Li, D. Wang, K. Zhao, K. Song, J. Liang, Electrohydrodynamic jet 3D printing of PCL/PVP composite scaffold for cell culture, Talanta (2020), doi: https://doi.org/10.1016/ j.talanta.2020.120750. This is a PDF file of an article that has undergone enhancements after acceptance, such as the addition of a cover page and metadata, and formatting for readability, but it is not yet the definitive version of record. This version will undergo additional copyediting, typesetting and review before it is published in its final form, but we are providing this version to give early visibility of the article. Please note that, during the production process, errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain. © 2020 Published by Elsevier B.V.
Electrohydrodynamic jet 3D printing of PCL/PVP composite scaffold for cell culture Kai Li1, Dazhi Wang1, 2, *, Kuipeng Zhao1, Kedong Song3 and Junsheng Liang1, 2 1
Key Laboratory for Micro/Nano Technology and System of Liaoning Province, Dalian University of Technology, Dalian, 116024, China. 2 Key Laboratory for Precision and Non-traditional Machining Technology of Ministry of Education, Dalian University of Technology, Dalian, 116024, China. 3 State Key Laboratory of Fine Chemicals, Dalian R&D Center for Stem Cell and Tissue Engineering, Dalian University of Technology, Dalian 116024, China.
Abstract Controlled printing of biodegradable and bioresorbable polymers at desired 3D scaffold is of great importance for cell growth and tissue regeneration. In this work, a novel electrohydrodynamic jet 3D printing technology with the resultant effect of electrohydrodynamic force and thermal convection was developed, and its feasibility to fabricate controllable filament composite scaffolds was verified. This method introduces an effective thermal field under the needle to simultaneously enhance the ink viscosity, jetting morphology controllability and printing structure solidify. The fabrication mechanisms of thermal convection on jetting morphology and printed structures feature were investigated through theoretical analysis and experimental characterization. Under optimized conditions, a stable and finer jet was formed; then with the use of this jet, various 3D structures were directly printed at a high aspect ratio ~30. Furthermore, the PCL/PVP composite scaffolds with the controllable filament diameter (~10 µm) which is closed to living cells were printed. Cell culture experiments showed that the printed scaffolds had excellent cell biocompatibility and facilitated cellular proliferation in vitro. It is a great potential that the developed electrohydrodynamic jet 3D printing technology might provide a novel approach to directly print composite synthetic biopolymers into flexibly scale structures for tissue engineering applications.
Keywords: 3D printing; electrohydrodynamic jet; thermal field; composite scaffold; synthetic biopolymers; cell culture Corresponding author: Dazhi Wang. Email: +86(411)84707170-2171; Fax: +86(411)84707940 Supplementary Information is available online. 1
[email protected];
Tel.:
1. Introduction Biopolymers can be derived from natural or synthetic materials and can take an important role in drug screening, organ repair and tissue engineering, which provide mass transport and temporary mechanical support to facilitate regenerated cells in lots of successful applications [1-4]. Compared with naturally derived polymers, synthetic biopolymers have the advantages of high biological compatibility, flexible controlling of the molecule and high function mimic of organs. And it can be biodegradable through hydrolysis or enzymatic degradation. In addition, the synthetic biopolymers with high tensile strength, high materials degradation rate and other unique properties can be created using chemistries such as photo-mediated degradation or crosslinking [5-8]. Hence, the synthetic biopolymer is now receiving increased attention because of their enormous need from regenerate damaged tissue to tissue culture [9-12]. Recently various synthetic biopolymers, such as polylactic acid (PLA) [13, 14], polycaprolactone (PCL) [15, 16], polyvinyl alcohol (PVA) [17, 18], polyvinyl pyrrolidone (PVP) [19, 20] and poly (lactic-co-glycolicacid) (PLGA) [21, 22], have been employed for tissue engineering and regenerative medicine aim to develop replacement tissues for our body. A number of biofabrication technologies enable the fabrication of biological constructs with precise control over the positioning of synthetic biopolymers. Unlike conventional biofabrication technologies that require moulds, masks and complex process, printing makes it possible to rapidly turn computer-aided designs into complex 3D objects on-demand [23-25]. The extrusion-based printing is one of the most commonly used biofabrication techniques which relies on “pushing” of high viscous synthetic biopolymers ink and its molten status. In this kind of technique, inkjet printing relies on the generation of droplets at the needle outlet, followed by 2
noncontact deposition on a substrate to fabricate synthetic biopolymers structures. The droplets jet from the needle due to the action of pressure pulses generated, which induced by thermal or piezoelectric means. It is obvious that a challenge for inkjet printing is the achievable levels of resolution, establishing minimum sizes of patterned features about 10 µm [26]. In addition, fused deposition modeling (FDM) and direct ink writing (DIW) extrudes the viscous state ink in a steady flow until deposition and must rapidly stabilize upon delivery [27-30]. These “pushing” printing techniques are simple and low cost. However, the resolution of extrusion-based printing technologies has a dependence on the needle inner size. The needle inner size used usually hundreds of micrometers to maintain the steady flow due to the high viscosity of printing ink. Thus, the feature size of the printed 3D structure is also larger than the needle size, which restricts application scope and device performance of synthetic biopolymer. Electrohydrodynamic jet printing (E-Jet) was thus developed to fabricate synthetic biopolymer patterning by controlling the movement of the collecting substrate because of its high-resolution characteristic [31, 32]. In the E-Jet printing process, the mobile ions in the synthetic biopolymers ink are induced by applying a high voltage potential and accumulate near the surface at the needle tip, and this leads to an intense electric field force around the pendent drop. E-Jet printing utilizes this electric field force to induce synthetic biopolymer ink forming a fine jetting from needle tip, which is based on “pulling” force rather than “pushing” force, exhibiting unique advantages of high resolution (sub-100 nm) [26]. Previous works have been demonstrated the compatibility of E-Jet printing with a wide variety of synthetic biopolymers inks, ranging from P(S-ran-MMA), PCL, PVA and many others. Various structures of dot, line and film with a feature size from hundreds of nanometer to micrometer have been printed [33-36]. Lee et al developed a metal
3
electrode assisted E-Jet printing process and the printed polyethylene oxide (PEO) filaments could be precisely stacked on a platinum line layer-by-layer deposition [37]. Wang et al integrated an air infusion modulus into the E-Jet printing system to increase the rapid solidification of printed PCL filaments [38]. Choe et al proposed a modified E-Jet printing technology to fabricate the micro/nanoscale hybrid scaffold with a controllable pore using the cellulose acetate biomaterial [39]. However, the previous E-Jet printing technologies are often used to print single synthetic biopolymers. And it is difficult to build high aspect ratio composite 3D structures because the low viscous synthetic biopolymers ink is likely to spread on the substrate and the printed structure is easy to collapse. In this work, a new E-Jet 3D printing technique under the resultant effects of electrohydrodynamic and thermal convection was developed to print PCL/PVP composite scaffold for cell culture. Firstly, a novel E-Jet 3D printing equipment with a thermal platform under the needle was designed and prepared for printing in which the electric field was applied on the needle and the thermal field was supplied by the thermal platform. A low viscous composite synthetic biopolymer ink (PCL/PVP) was prepared after their particles were dissolved in low boiling point solvent. Then, the electrohydrodynamic effect provided an electrical shearing force to induce a fine jet and the thermal convection was employed to increase the viscosity of the ink by evaporating the low boiling point solvent during printing. Eventually, the stable and controllable fine jet was pulled out and formed at the needle tip. It should be noted that atomization could be likely to happen if the electrohydrodynamic effect was only applied to this low viscous biopolymer ink. Even though the jet can be formed it is impossible to pattern 3D structures because of the low viscosity of the ink. And electrospinning normally occurs when a high viscous ink was directly used for 4
printing under high electrical field, causing the jet whipping, which resulting in difficult to control of printing. The resultant effects of thermal field and electric field on jetting behaviors and size of printed 3D structures were deeply investigated through experimental characterization. By optimizing the parameters of printing process, various 3D structures with a high aspect ratio (~ 30) were efficiently created. Furthermore, 3D PCL/PVP composite scaffolds cultured with Murine MC3T3-E1 Subclone14 cartilage cells were printed flexibly as a platform to study cell regeneration in vitro. Excellent cell compatibility of printed 3D scaffolds promoted cell growth and proliferation. Higher cell density and filling the internal pores of the scaffolds under 5 days of culture were observed compared to 3 days. Overall, this new E-Jet printing technique develops a promising way to fabricate high aspect ratio 3D structure and provides a great potential for tissue engineering applications.
2. Materials and methods 2.1 Experimental set-up The schematic illustration of printing set-up is shown in figure 1a, including an electrohydrodynamic needle, a syringe pump, a high voltage power supply, a thermal platform, a computer-controlled X-Y-Z movement stage and a camera. The electrohydrodynamic needle is a quartz needle with an inner/outer diameter of 100 µm/380 µm (New Objective, Inc, USA). The quartz needle is installed in a channel of metallic fixture. And the metallic fixture is connected to the output of high voltage power supply (Physical Instruments, Germany). In this condition, the high voltage could be applied to the ink through this metallic fixture. The inlet of metallic fixture is connected to the glass syringe (100 µL, Hamilton, USA) by a hose to feed the low viscous biopolymer composite ink. The glass syringe was placed in a syringe pump
5
(Harvard Apparatus, Holliston, USA). The thermal platform with a resistance heater is fixed on X-Y-Z movement stage (uKSA, Zolix Instruments Co., Ltd, China) to provide thermal field during printing. A temperature-control device is used for ensuring the heating uniformity of thermal platform. A control program is developed based on LabVIEW, which is employed to coordinate the movement of X-Y-Z motion stage. A camera (Point Grey, Canada) is used for observing the jetting behaviors. 2.2 Preparation of printable PCL/PVP composite ink PCL ink and PVP ink were prepared separately and two inks were mixed to form the PCL/PVP composite ink. PCL was not only a good biologically compatible synthetic biopolymer but also insoluble in water and alcohol. Therefore, the PCL and PVP biopolymer were chosen as the scaffold material while PVP was added to modulate the viscosity of the printable ink. 9.1 wt.% of PCL particle (Mw=800,000, yuanye Bio-Technology Co., Ltd, China) was dissolved into acetic acid (Adamas Reagent Co., Ltd, China) and mixed for 50 minutes at 80 °C. The preparation of PCL ink was completed. 11.7 wt.% of PVP powder (Mw=1,300,000, Aladdin Industrial Corporation, China) was dissolved into ethanol. And 0.43 wt.% of PEG (Mw=800, Kemiou Chemical Reagent Co., Ltd, China) was added to the solution and mixed for 5 minutes at 50 °C. After the mixed solution was cooled to 25 °C, 1.7 wt.% of Triton X-100 (Kemiou Chemical Reagent Co., Ltd, China) was added to the solution and mixed for 10 minutes, the PVP ink can be obtained. Finally, the PCL ink and PVP ink were mixed for 20 minutes at 25 °C to form printable PCL/PVP composite ink (figure S1a). As shown in figure S1a, the surface tension of PCL/PVP ink was measured by an automatic surface tensiometer (K100C, Kruess Scientific Instrument, Germany). The measurement resolution is 0.01 mN m-1 and the measurement range is 1-500 mN m-1. The surface tension of the PCL/PVP composite ink was tested ten times and its 6
mean value was calculated. 2.3 3D printing scaffold process The low viscous PCL/PVP composite biopolymer ink with low boiling point was placed in the glass syringe and infused to the quartz needle through the syringe pump which provides the hydrodynamic force to push PCL/PVP ink at a constant flow rate. Subsequently, an applicable temperature of thermal platform was selected to increase the viscosity of composite ink. Afterwards, the viscosity increased with the solvent evaporation under the thermal field (figure 1b). And the jetting formed under the action of high electric shearing force provided by a high voltage power supply. The stable and controllable finer jet was pulled out and formed at the needle tip. Finally, the PCL/PVP composite scaffolds were overlaid layer by layer through the X-Y-Z movement stage controlled by a LabVIEW program. In addition, the camera was used to observe the printing process in real-time. 2.4 Cell culture on printed scaffolds The Murine MC3T3-E1 Subclone14 cartilage cells were cultured in α-Dulbecco's Modified Eagle's Medium (α-DMEM, Hyclone, USA) containing 10 % v/v fetal bovine serum (FBS, MINHAI, China), 100 units mL-1 penicillin and 100 µg mL-1 streptomycin (Hyclone, USA). Before cell seeding, the printed PCL/PVP scaffolds were sterilized in 75 v/v% alcohol for 5 h. Subsequently, 20 µl cell suspensions with a density of 5×105 µg mL-1 were drop-seeded on both sides of each scaffold and incubated for 3 h at 37 °C and 5 % v/v CO2 atmosphere, followed by supplement of fresh medium. 2.5 Evaluation of cell growth on composite scaffolds On 3 days and 5 days after cell seeding, Dead/Live staining was performed to evaluate cell viability and attachment. Calcein-AM, propidium iodide (PI) and 7
Hoechst 33258 were purchased from Calbiochem (San Diego, CA, USA). Briefly, cell-scaffold constructs were incubated in PBS containing 2 µmol L-1 calcein-AM, 4 µmol L-1 propidiumiodide (PI) and 5 µg L-1 Hoechst 33258 staining solution at 37 °C for 30 min. Cell viability on the PCL/PVP composite scaffolds was observed through four fluorescent images which were randomly taken using a fluorescence microscope (IX71 + DP71, Olympus, Japan). The scale in each confocal micrograph was saved in Image-Pro Plus 6 (IPP, Media Cybernetics, USA) software and set as the benchmark. The normalized cell densities (the initial cell numbers per mm2 are normalized to the level of those in the composite scaffolds) are measured within 5 days. The cell viabilities are estimated after culturing for 3 days and 5 days. Furthermore, the cells on the PCL/PVP scaffolds were visualized using field emission scanning electron microscopy (FE-SEM, SU8220, HITACHI, Japan). For FE-SEM observing, the cell-scaffold constructs were fixed with 2.5% glutaraldehyde at 4 °C for 3 h and gradually dehydrated in an ascending series of ethanol (50%, 70%, 90% and 100%) for 30 min for each solution. And these FE-SEM samples were dried in a vacuum condition. After sputter-coated with gold, the samples were observed by FE-SEM with an accelerating voltage of 5.0 kV.
3. Results and discussion 3.1 Effect of thermal field and electric field on jetting behaviors and feature of printed 3D structures The temperature of thermal platform has a great effect on printing behaviors and printed 3D structure features. In order to obtain a stable jet and desired composite scaffold, the effect of thermal platform on the printing mechanisms were examined. In this paper, the thermal platform (as a heat source) under the needle is made by a square aluminum plate. Therefore, the thermal convection during printing is the heat 8
dissipation from a thermal platform surface to its upper PCL/PVP composite ink. And thermal convection takes place in the composite ink through collisions or oscillations of each molecule. The equations used for simulation analysis are shown in the Supplementary Information. Figure 2 shows the simulated results of temperature distribution between the thermal platform and needle tip using COMSOL Multiphysics software. When the temperatures of thermal platform were 30 °C (figure 2a), 35 °C (figure 2b) and 45 °C (figure 2c), the temperatures of needle tip were 27 °C, 32 °C and 38 °C, respectively. And the temperature of needle rose to 47 °C when the temperature of thermal platform increased at 55 °C (figure 2d-e). The thermal convection near the needle tip is more intense when the temperature of thermal platform increased. And this thermal convection was employed in evaporating the solvent in PCL/PVP composite ink to increase its viscosity during printing. Figure 2f shows the simulated temperature contrasting of thermal platform and needle tip. In the PCL/PVP composite ink, the boiling points of two most solvents (ethanol and acetic acid) were 78 °C and 117.9 °C. Furthermore, these two solvents were volatile. It is important for a suitable temperature of thermal platform to evaporate volatile solvents and obtaining viscous ink. And a stable and controllable fine jet was pulled out and formed at the needle tip. Figure 3a-d show the jetting behavior affected by the temperature of thermal platform, ranging from 30 °C to 55 °C. The applied voltage, flow rate, printing height and speed of movement stage were at 650 V, 0.025 µL min-1, 150 µm and 2 mm s-1, respectively. Figure 3a-b shows that the Taylor cone was formed and the PCL/PVP biopolymer droplet was ejected from the cone tip when the temperature of thermal platform was lower than 35 °C. This is because the temperature of needle tip (32 °C showing in figure 2b) is far below the boiling point of two solvents and most solvents 9
have still stayed in the jetting. When the temperature of thermal platform rose to 45 °C, the temperature of the needle tip was 38 °C. In this printing condition, the acceleration of solvent evaporation (mainly ethanol) and increase of the viscosity of PCL/PVP ink occurred. The jetting behaviors of droplets along with jet at the tip of the cone were formed (figure 3c). However, the temperature was not high enough to form a semisolid jet and maintain the continuous jet. Figure 3d shows the stable semisolid jet was formed when the temperature of thermal platform rose to 55 °C. At this temperature the two volatile solvents (ethanol and acetic acid) were evaporated rapidly, then the viscosity of the PCL/PVP ink at the outlet of needle was further increased, resulting in the semisolid state. And it was viscous enough to support the formation of a stable continuous semisolid jet. It was suggested that the thermal platform temperature of 55 °C could significantly promote the evaporation of solvents and rising its viscosity. Figure 3e-h show the 3D profiles of printed PCL/PVP structures with 20 layers when the temperatures of thermal platform at 30 °C, 35 °C, 45 °C and 55 °C, respectively. Figure 3e shows the width and height of the printed PCL/PVP structure were 229±10.1 µm and 11±4.2 µm. And the aspect ratio of PCL/PVP structure was small. This was because the viscosity of PCL/PVP ink at the outlet of needle was low which would lead to ink dispersed on the substrate after printed. In contrast, as shown in figure 3e and 3h, the width of printed 3D PCL/PVP structures gradually decreased from 229±10.1 µm to 13±4.1 µm while the height increased from 11±4.2 µm to 110±3.1 µm as the temperature of thermal platform increased from 30 °C to 55 °C. Five positions were randomly selected on printed PCL/PVP structure and measured its sizes. The statistical analysis in figures 3i shows the increase of temperature could help to print high aspect ratio 3D structures. Therefore, the optimized temperature of 10
thermal platform was 55 °C. Figure 4 shows the effect of different applied voltages on jetting behaviors and size of printed 3D structures. In E-Jet printing process, the applied voltage plays a key role in controlling jet behaviors. And it affects the jet diameter and the aspect ratio of printed 3D structures simultaneously. During printing, the surface charge of ink increased continually and a meniscus was induced when the voltage was applied. Then the ink meniscus was deformed and the jetting was formed at the end of needle. In general, in order to overcome the surface tension of ink meniscus and forming a jetting, the charge on the meniscus is satisfied with the Rayleigh limiting charge [40],
Qmax = π(8ε0Fst dd3 )1/2 where Qmax is the maximum amount of carrying charge on the drop surface, ε0 is the ink vacuum permittivity, Fst is the ink’s surface tension, dd is the diameter of the droplet. In this work, the mean surface tension of PCL/PVP composite ink is 23.48 mN m-1 at 20 °C (ten measured data showing in figure S2). And equation 11 suggests that the ink with low surface tension just requires a lower voltage to induce the jetting. As shown in figure 4a, a jet formed at the end of needle when a voltage of 400 V was applied. When the applied voltage was increased, the jet diameter was decreased (figure 4a-c). This thin jet can help to print uniform 3D structures with a high respect ratio. It can be seen from the experiment process that multi-jet was generated when 1200 V was applied (figure 4d), which is an undesirable jetting behavior and cannot be used in precise printing 3D structures. From figure 4a to 4d, the angle (θ) between the jet and the direction of printed structures (horizontal direction) increased from 34.31° to 84.47° when the applied voltage increased from 400 V to 1000 V. It can be seen the applied voltage has a 11
marked impact on the orientation of jetting filament. When the applied voltage was lower (400 V, figure 4a), the drawing force (Fd) is the leading force to pull the jetting filament. And the drawing force promotes the stacking of filaments on 3D structures layer-by-layer. The orientation of jetting filament is more curved and the θ is small (34.31°, figure 4a) because of the effect of the drawing force. When the applied voltage increased the jetting rate (U) and the electric field force (Fe) increased. Thus the effect of Fd decreases while the effect of Fe increases and the angle (θ) increases at the same time. Figure 4e-h shows 3D profiles of printed PCL/PVP structures with 20 layers when the applied voltage was 400 V, 600 V, 800 V and 1200 V, respectively. As shown in figure 4e-g, the 3D structure width gradually decreased from 15±6 µm to 8±1.5 µm while the height increased from 51±4 µm to 86±3 µm. However, the multi-jet occurred when the applied voltage rose to 1200 V (figure 4d) because the normal electrical field force was much larger than the surface tension. In this condition, the cone-jet mode was destroyed and the flat meniscus with small cones formed around the end of needle. And tiny jets were ejected from these small cones and arrived at the substrate (figure 4h). Five positions were randomly selected on the printed PCL/PVP structure and measured its sizes. The statistical analysis of the effect of applied voltage on 3D structures size is showed in figure 4i. It can be seen that higher aspect ratio 3D structures can be obtained when the applied voltages increased up to 1200 V. 3.2 Printing 3D structures with high aspect ratio Figure 5a shows SEM image of 3D lattice structure programmed printed layer-by-layer. The size of every square lattice was 200 µm × 200 µm. During the printing process the temperature of thermal platform, applied voltage, flow rate,
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printing height and speed of movement stage were at 55 °C, 600 V, 0.03 µL min−1, 150 µm and 2 mm s-1, respectively. After 15 layers printing, the width of the printed 3D lattice was about 11 µm and the height was about 80 µm, the aspect ratio was about 7. The 3D profile in figure 5b shows the printed 3D lattice had the uniform and smooth features. At the condition of low applied voltage (600 V), small distance (150 µm) and low printing speed (2 mm s-1), the misplacing has not appeared in layer upon layer and the jet was kept vertical throughout the printing process. With the movement stage went back and forth at a speed of 2 mm s-1 in X-axial direction, a 3D structure consisting of 40-layers polyline were printed and the filaments were stacked orderly together on the substrate. As showed in figure 5c, the aspect ratio of printed 3D perpendicular walls was about 25 and each layer was completely overlapped. During the printing process the temperature of thermal platform, flow rate and speed of movement stage were at 55 °C, 0.03 µL min-1 and 2 mm s-1, respectively. Furthermore, in order to maintain an effective electric field force, the applied voltage increased (600 V-800 V) as the height of printed 3D perpendicular walls increased. These various 3D structures proved that this printing method is capable of constructing 3D microstructures with high aspect ratio. 3.3 Printing and evaluation of PCL/PVP composite scaffold The meaning of bioactivity is a biomaterial that has the ability to promote the formation of hydroxyapatite on its surface, while the biomaterial immersed in simulated body fluid. Previous studies have been proved that PCL scaffold has poor bioactivity probably due to the bio-inertness and slow degradation nature [14]. Thus 13
the incorporation of PVP into printed PCL microscale filaments would effectively promote the bioactivity of PCL/PVP composite scaffold. Furthermore, the PVP has higher mechanical stiffness and modulus. The incorporation of PVP into PCL microscale filaments would significantly increase the mechanical properties of the composite scaffold. In order to determine the existence of PCL and PVP in the printed composite scaffold, the PCL scaffold, PVP scaffold and the PCL/PVP composite scaffold were printed respectively. Then the FTIR spectra of these three scaffolds were performed. As shown in the red line of figure 6, the PCL scaffold spectrum exhibited the peak of C=O and stretching vibrations at 1702 cm-1, CH2 bending modes at 1359 cm-1, 1408 cm-1 and 1438 cm-1 and COO– vibrations at 1286 cm-1. The C–O– C stretching vibration peaks at 1221 cm-1, 1161 cm-1 and 1106 cm-1 [41]. The PVP scaffold spectrum (blue line) exhibited the intense sharp peak of C=O stretching vibrations of PVP at 1663 cm-1. The CH2 bending modes were observed at 1378 cm-1, 1424 cm-1, 1458 cm-1 and 1487 cm-1 and that of the C–N vibrations occurred at 1045 cm-1 [42]. The spectrum of PCL/PVP composite scaffold exhibited the dominant peak of both PCL (C=O stretch 1702 cm-1) and PVP (C=O stretch at 1663 cm-1). Moreover, it was found the COO– stretch of PCL at 1286 cm-1 and C–N stretch of PVP at 1045 cm-1. These results indicated the PCL/PVP composite scaffold caused almost no change in the FTIR spectra of each characteristic peaks, indicating no chemical bonding or physical interaction occurs between PCL and PVP components. The internal microstructure of tissue engineered 3D scaffolds plays an important role to provide a proper growth microenvironment for improving cell growth and proliferation. Previous studies have proved that the appropriate filament size and its spacing in the scaffold not only promotes cell seeding but also provides an effective nutrient exchange [7, 33]. As shown in figure 7a and figure 8a, the 3D porous 14
PCL/PVP composite scaffolds were fabricated on a glass substrate (Supplementary Video S1). During the printing process the temperature of thermal platform, flow rate, printing height and speed of movement stage were at 55 °C, 0.05 µL min-1, 200 µm and 8 mm s-1, respectively. It was printed for 8 layers, one layer was printed as X direction, then the adjacent layer was printed as Y direction (Supplementary Video S2). With the increase of the number of printing layers, the distance between the needle and the substrate increases gradually. Meanwhile, the applied voltage was slightly adjusted (600-700 V) in order to keep electric field intensity between the needle and ground electrode to obtain the stable printing process. The average distance between two adjacent filaments was ~80 µm. It also can be seen that the distance between the two adjacent filaments was a small difference, which was because that the previously printed filament can attract the printing jet when the printing height is small, thus resulting in the position error between the two adjacent filaments. The average width of the scaffold filaments was about 10 µm. And this filament size is closed to living cells and is suitable to provide sufficient surface area for cell survival. Furthermore, the scaffold size can be flexibly fabricated to mimic cell growth microenvironment, which can provide desired structures for improving cell growth and proliferation. Dead/Live staining of the 3D microtissue of Murine MC3T3-E1 Subclone14 cartilage cells grown in printed PCL/PVP composite scaffold was conducted 3 days and 5 days of culture. As shown in figure 7 and figure 8, the LSCM fluorescent images of live/dead-stained on the two scaffolds exhibit a high cell seeding percentage and a high cell survival rate. An overwhelming proportion of living cells were positively stained by calcein-AM with strong green fluorescence signals. The cell density in the scaffolds under 5 days of culture (figure 8) was apparently higher 15
than that of the scaffold under 3 days of culture (figure 7). And the cells under 5 days of culture were found to significantly proliferate and fill the internal pores of the scaffolds (figure 8e). As shown in figure 8c (5 days of culture), the increased number of dead cells (a very small proportion) stained by PI was inevitable and acceptable over time, possibly due to the lack of available nutrients and undischarged waste in the interior of printed scaffolds. After the PCL/PVP composite scaffolds were sterilized in 75 % ethanol aqueous solution for 5 h, the surface roughness of the printed filaments increased due to the partial dissolution of PVP components. And the overall PCL/PVP composite scaffold and its filaments could be stably maintained. Nanoscale pores exposed on the surface of the printed filaments could be promoted cellular attachment. [43]. Furthermore, 3D view images as shown in figure 7f and figure 8f displayed that the cells attached from the surface to the inside of the scaffold and the depth of the ingrowth of the cells reached about 25 µm. As shown in figure 9a, the normalized Murine MC3T3-E1 Subclone14 cartilage cells densities in PCL/PVP composite scaffold are measured within 5 days. The results indicate that the cell densities increased as the days of cell culture increased. And the number of cells has been more than four times within 5 days in composite scaffolds. The cell viability of cells has been measured according to the results of live/dead stain (figure 9b). The cell viability has been evaluated by calculating the ratio of number of live cells to number of total cells. The cell viability after cell culture of 3 days and 5 days are 91 %±2.1 % and 95 %±3.5%. In addition, after cell culture on the printed composite scaffolds for 3 days and 5 days, SEM observation was performed to examine the proliferation and morphology of cells. As shown in figure 9(c-f), the cells with the long pseudopodia attached well on the channel surface of composite scaffolds. At cell culture of 5 days, cell density on the composite 16
scaffolds was obviously higher than that at 3 days and the most area of the surface of scaffolds was covered with cells. These results were consistent with the results of LSCM fluorescent shown in figure 7 and figure 8. In a word, these results indicated favorable viability of cells, suggesting good cell compatibility of the printed PCL/PVP composite scaffolds. This novel E-Jet 3D printing technique proposed in this paper demonstrates a new and promising way to pattering high molecular weight biopolymer for tissue engineering. 4. Conclusion A novel E-Jet 3D printing technique, with the unique advantages of high aspect ratio feature and large controllability, is introduced to direct print microscale synthetic biopolymers composite scaffold for cell culture. The translation of synthetic biopolymers composite ink from low viscous to high viscous was realized with the thermal field effect and the local solution evaporated in printing process. The external electric field force created by a constant voltage induced this high viscous ink to form a focused and stable jet. Under the resultant effect of electrical shearing force and thermal field effect, the controllable fine filament was ejected from the needle tip. The jetting behavior of jetting filament and printed 3D features have been studied at different experimental parameters and the optimized printing condition was selected to fabricate high aspect ratio 3D structures layer-by-layer. Microscale PVL/PVP composite 3D scaffolds with excellent cell compatibility were printed for tissue growth and proliferation. And the cell culture experiments proved that with the increase of culture time, higher cell density was also observed on composite 3D scaffolds. The new E-Jet 3D printing technique presented in this paper demonstrates a promising way to print composite synthetic biopolymers for tissue engineering.
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Acknowledgment This research was supported by National Natural Science Foundation of China (51975104), National Key R&D Program of China (2018YFA0703200), the aerospace science foundation (2018ZD63004), the Fundamental Research Funds for the Central Universities (DUT18LAB17), State Key Laboratory of Precision Measuring Technology and Instruments (Tianjin University) (pilab1804) Science Fund for Creative Research Groups of NSFC (51321004) and the Collaborative Innovation Center of Major Machine Manufacturing in Liaoning. Reference [1] D. F. Williams, On the nature of biomaterials, Biomaterials 30 (2009) 5897-5909. [2] R. Parenteau-Bareil, R. Gauvin, F. Berthod, Collagen-Based Biomaterials for Tissue Engineering Applications, Materials 3 (2010) 1863-1887. [3] M. J. Webber, E. A. Appel, E. W. Meijer, R. Langer, Supramolecular biomaterials, Nature Materials 15 (2016) 13-26. [4] M. Hospodiuk, M. Dey, D. Sosnoski, I. T. Ozbolat, The bioink: A comprehensive review on bioprintable materials, Biotechnology Advances 35 (2017) 217-239. [5] E. S. Place, J. H. George, C. K. Williams, M. M. Stevens, Synthetic polymer scaffolds for tissue engineering, Chemical Society Reviews 38 (2009) 1139-1151. [6] D. F. Williams, On the nature of biomaterials, Biomaterials 30 (2009) 5897-5909. [7] T. G. Kim, H. Shin, D. W. Lim, Biomimetic scaffolds for tissue engineering, Advanced Functional Materials 22 (2012) 2446-2468. [8] Y. L. Kong, M. K. Gupta, B. N. Johnson, M. C. McAlpine, 3D Printed bionic nanodevices, Nano Today 11 (2016) 330-350. [9] S. Bose, D. Ke, H. Sahasrabudhe, A. Bandyopadhyay, Additive manufacturing of biomaterials, Progress in Materials Science 93 (2018) 45-111. [10] A. Sorushanova, L. M. Delgado, Z. Wu, N. Shologu, A. Kshirsagar, R. Raghunath, A. M. Mullen, Y. Bayon, A. Pandit, M. Raghunath, D. I. Zeugolis, The collagen suprafamily: from biosynthesis to advanced biomaterial development, Advanced Materials 31 (2019) 1801651. [11] S. V. Murphy, A. Atala, 3D bioprinting of tissues and organs, Nature
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Figure captions Figure 1. (a) Schematic diagram showing the E-Jet printing process; (b) The solvent evaporation during printing. Figure 2. Simulations of the temperature distribution at different temperatures of thermal platform; (a-d) Simulation images of thermal platform at 30 °C, 35 °C, 45 °C and 55 °C, respectively; (e) Showing the red frame in (d); (f) Simulated temperature relationship of thermal platform and needle tip. Figure 3. Experimental 3D printing at different temperatures of thermal platform; (a-d) The jet behaviors at 30 °C, 35 °C, 45 °C and 55 °C, respectively (scale bar=50 µm); (e-h) 3D profiles of printed structure at 30 °C, 35 °C, 45 °C and 55 °C, respectively; (i) Statistical graph showing the sizes of the printed 3D structures at different temperatures of thermal platform. Figure 4. Experimental 3D printing at different applied voltages; (a-d) The jet behaviors at 400 V, 800 V, 1000 V and 1200 V, respectively (scale bar=50 µm, θ meaning the included angle between jetting filament and the direction of printed structures); (e-h) 3D profiles of printed structure at 400 V, 600 V, 800 V and 1200 V, respectively; (i) The statistical graph showing the sizes of the printed 3D structures at different applied voltages. Figure 5. (a-b) SEM image and 3D profile of 3D lattice structure; (c) 3D profile of high aspect ratio walls. Figure 6. (a) Fourier transform infrared (FTIR) spectra for PCL scaffold, PVP scaffold, PCL/PVP composite scaffold, respectively; (b) The image of the boxed region in (a) and its characteristic peaks. Figure 7. Printing 3D porous PCL/PVP scaffold and evaluation of Murine MC3T3-E1 22
Subclone14 cartilage cells activity and distribution after 3 days culture. (a) The printed PCL/PVP composite scaffold; (b-f) Live/dead staining of cells on the scaffold; Live cells were stained with calcein-AM, dead cells were stained with PI and cell nuclei were stained Hoechst 33258, showing green, red and blue, respectively. Scale bars in a-e=200 µm, f=100 µm. Figure 8. Printing 3D porous PCL/PVP scaffold and evaluation of Murine MC3T3-E1 Subclone14 cartilage cells activity and distribution after 5 days culture. (a) The printed PCL/PVP composite scaffold; (b-f) Live/dead staining of cells on the scaffold; Scale bars in a-e=200 µm, f=100 µm. Figure 9. (a) The normalized densities of Murine MC3T3-E1 Subclone14 cartilage cells inside PCL/PVP composite scaffolds; (b) Viability of Hela cells inside composite scaffolds. ** indicate statistical significance of P < 0.01 (n=4). (c-f) The morphology of cells on the printed PCL/PVP scaffolds. SEM observations in (c-d) and (e-f) were performed to examine the morphology of cartilage cells after cells seeded on the scaffolds for 3 days and 5 days. Scale bars in c-f=100 µm.
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Highlights A novel E-Jet 3D printing method was developed for patterning synthetic biopolymers. Various 3D biopolymers structures were directly printed at a high aspect ratio ~30. 3D printed PVP/PCL composite scaffolds with controllable filament were developed. The scaffolds could act as favorable synthetic ECM for cartilage regeneration.
Declaration of interests ☒ The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. ☐The authors declare the following financial interests/personal relationships which may be considered as potential competing interests: