High-speed on-demand 3D printed bioresorbable vascular scaffolds

Materials Today Chemistry 7 (2018) 25e34

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High-speed on-demand 3D printed bioresorbable vascular scaffolds Henry Oliver T. Ware a, Adam C. Farsheed a, Banu Akar b, Chongwen Duan b, Xiangfan Chen a, Guillermo Ameer b, c, **, Cheng Sun a, * a

Mechanical Engineering Department, Northwestern University, 2145 Sheridan Road, Evanston, IL 60208, USA Biomedical Engineering Department, Northwestern University, 2145 Sheridan Road, Evanston, IL 60208, USA c Department of Surgery, Feinberg School of Medicine, Chicago, IL, 60611, USA b

a r t i c l e i n f o

a b s t r a c t

Article history: Received 18 July 2017 Received in revised form 11 October 2017 Accepted 14 October 2017

Recent development of the high-resolution Micro-Continuous Liquid Interface Production (mCLIP) process has enabled 3D printing of biomedical devices with micron-scale precision. Despite our recent success in demonstrating fabrication of bioresorbable vascular scaffolds (BVS) via mCLIP, key technical obstacles remain. Specifically, achieving comparable radial stiffness to nitinol stents required strut thickness of 400 mm. Such large struts would negatively affect blood flow through smaller coronary vessels. Low printing speed also made the process impractical for potential on-demand fabrication of patient-specific BVSs. Lack of a systematic optimization strategy capturing the sophisticated processmaterials-performance dependencies impedes development of on-demand fabrication of BVSs and other biomedical devices. Herein, we developed a systematic method to optimize the entangled process parameters, such as materials strength/stiffness, exposure dosage, and fabrication speed. A dedicated speed working curve method was developed to calibrate the mCLIP process, which allowed experimental determination of dimensionally-accurate fabrication parameters. Composition of the citric acid-based bioresorbable ink (B-Ink™) was optimized to maximize BVS radial stiffness, allowing scaffold struts at clinically-relevant sizes. Through the described dual optimization, we have successfully fabricated BVSs with radial stiffness comparable to nitinol stents and strut thickness of 150 mm, which is comparable to the ABSORB GT1BVS. Fabrication of 2-cm long BVS with 5 mm, 10 mm, and 15 mm layer slicing can now be accomplished within 26.5, 15.3, and 11.3 min, respectively. The reported process optimization methods and high-speed, high-resolution 3D printing capability demonstrate a promising solution for on-demand fabrication of patient-specific biomedical devices. © 2017 Elsevier Ltd. All rights reserved.

Keywords: 3D printing Micro-continuous liquid interface production (mCLIP) Bioresorbable vascular scaffold

1. Introduction Stents are a class of biomedical implants used within the body to reopen narrowed or obstructed passageways in the vasculature, esophagus, and gastrointestinal tract [1e5]. Coronary stents help to restore the lumen and normalize the optimal distal blood flow of blood vessels. While stenting for obstruction relief is safe and effective, issues such as in-stent restenosis, stent thrombosis, and inflammation still pose a challenge [6,7]. The drug eluting stent

* Corresponding author. Mechanical Engineering Department, Northwestern University, 2145 Sheridan Road, Evanston, IL 60208, USA. ** Corresponding author. Biomedical Engineering Department, Northwestern University, 2145 Sheridan Road, Evanston, IL 60208, USA. E-mail addresses: [email protected] (G. Ameer), c-sun@northwestern. edu (C. Sun). https://doi.org/10.1016/j.mtchem.2017.10.002 2468-5194/© 2017 Elsevier Ltd. All rights reserved.

(DES) was developed to release antiproliferative therapeutics that help prevent restenosis, but other problems still persist, including late stent restenosis, stent fracture, and myocardial infarction [8]. Further, after many years within the body and as the antiproliferative drug coating dissipates, target lesion revascularization can be compromised and very late stent thrombosis can also occur [9]. In some pediatric cases, a permanent stent is not ideal as it will not accommodate the growth of the body as the patient matures [10]. Further, it has been suggested that stents are not needed for longer than 6 months to perform their desired function [11]. In response, the bioresorbable vascular scaffold (BVS) has been developed. The BVS is meant to remain within the body while arterial remodeling and lesion healing occur, at which point it will naturally be resorbed by the body [12,13]. BVSs provide the potential to accommodate the growth of pediatric patients and to eliminate some of the long-term problems associated with traditional stents [8].

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Metal and polymer-based stents (standard and BVS) have commonly been manufactured via laser machining of a hollow tube of primary material [1]. Although laser machining has high accuracy, it can cause thermal and/or chemical defects at the machining sites [14]. With the relatively recent advent of additive manufacturing (3D printing) processes, the opportunity to create patient specific medical implants, at relatively low cost, has arisen [15e17]. Several groups have reported coronary stent/scaffold fabrication via additive manufacturing methods. Park et al. fabricated a bioresorbable drug-coated scaffold via extrusion techniques on the surface of a cylindrical template followed by the spraycoating of the immunosuppressive drug sirolimus. The sustained release of sirolimus from the stent was observed in the porcine animal studies [18]. However, the fabricated stent geometry has to be strictly conformal to the cylindrical template being used, which severely limits the flexibility in design and fabrication of the customized stent/scaffold geometries. Flege et al. fabricated coronary scaffolds consisting of Poly-L-lactic acid (PLLA) and PLLA-copolycaprolactone via selective laser melting (SLM) process [19]. SLM tends to have poor surface finish, which requires additional dip and spray coating processes to smoothen stent/scaffold surfaces [20,21]. Fused deposition modeling and traditional stereolithography have been used in stent fabrication as well, but have only been reported in tracheobronchial stents and other larger diameter devices, such as stents for heart valve replacement [22,23]. All additive methods listed above require point-by-point scanning of the material to produce/polymerize each fabrication layer, which leads to long production times and inhomogeneous structural properties [22,24e26]. Projection microstereolithography (PmSL) manufacturing speeds up the print time by utilizing projected UV light patterned via dynamic mask to photopolymerize a full cross-sectional layer in a single exposure [16,17,24]. The continuous liquid interface production (CLIP) method further developed projection-based stereolithography with the addition of an oxygen permeable window and a continuously moving build substrate. This method has been shown to create 3D geometries at very high speeds, with good surface finish, and uniform mechanical properties [27]. Utilizing a similar concept, we have developed a high-resolution micro-continuous liquid interface production (mCLIP) method. Through experimental study, we evaluated the manufacturability of a BVS composed of a citric acid-based bioresorbable biomaterial ink (B-Ink) as the primary building material. We have previously reported through use of a PmSL system, fabrication of a 2-cm long scaffold takes approximately 16 h [28]. In contrast, using a mCLIP system, BVSs of the same size were able to be fabricated in 70 min with much better surface finish. Despite the successful demonstration of a 3D printed BVS with strength comparable to commercially available bare metal nitinol stents, it was imperative that the 3D printing process and B-Ink formulation be further optimized to reduce the profile of the device [28]. In particular, while previously reported 3D printed scaffolds have comparable strut width to standard stents/scaffolds, the 3D manufactured scaffolds required rather large strut thicknesses (400 mm and higher) to achieve radial stiffness comparable to commercial nitinol stents. Currently used coronary BVSs, such as the Absorb GT1, have strut thicknesses at or below 200 mm [29,30]. A strut thickness of 400 mm would have a negative impact blood flow in small diameter arteries, hence this dimension should be reduced to maximize the impact of 3D printed BVS. In addition, Moore et al. stated that if the supply chain of stent fabrication could be reduced to a singular machine and process time to approximately 20 min, the genuine possibility of patient-specific implants for emergency and urgent care patients may be achieved [1]. This represents a target goal for 3D printed coronary scaffold fabrication. Thus, the primary

objective of this study was to optimize the mCLIP process in order to fabricate, within 20 min, a 3D printed BVS with mechanical properties comparable to those of nitinol stents, with strut thickness less than 200 mm. Lastly, the biocompatibility of diethyl fumarate, the solvent used in Ref. [28], may be a challenge for regulatory approval. This solvent was utilized previously to demonstrate feasibility due to its very slow evaporation rate for PmSL. Replacement of the diethyl fumarate with a solvent that has better biocompatibility is necessary to produce a product suitable for implantation. Herein we developed an in-process calibration method suitable for the mCLIP process to ensure the dimensional accuracy of 3D printed parts and experimentally optimized the B-Ink pre-polymer and photoinitiator system concentrations to maximize the final scaffold radial stiffness. Upon optimization of the composition of the B-Ink as well as the key mCLIP process parameters, strut thickness was successfully reduced below 200 mm while maintaining a radial stiffness comparable to that of a 4-cm long control bare metal nitinol stent. We also successfully reduced the fabrication time from the 70 min described in our previous work [28] to 26.5 min, 15.3 min, and 11.5 min for a 2-cm long BVS at 5 mm, 10 mm, and 15 mm layer slicing thickness, respectively. 2. Material and methods 2.1. Material information Methacrylated poly(1e12 dodecamethylene citrate) (mPDC) was prepared as described in the previous literature [31,32]. mPDC has been shown to be biocompatible, biodegradable, photopolymerizable, and antioxidant [28,32]. Bioresorption and antioxidant properties of mPDC also extend to the 3D printed scaffolds, showing approximately 25% degradation in 37
C phosphate buffered saline and capability to scavenge free radicals [28]. The photopolymerizable ink in this study (hereafter known as B-Ink when referring to full ink) for mCLIP was composed of the prepolymer (mPDC); solvent (ethanol, Sigma Aldrich); primary photoinitiator (Irgacure 819, BASF); co-photoinitiator Ethyl 4-dimethylamino benzoate (EDAB, Sigma Aldrich); and UV absorber (Tinuvin 171, Sigma Aldrich). Throughout this study, various B-Ink ingredient formulations were investigated experimentally. mPDC and EDAB concentrations were varied between 50% (wt./wt.) to 70% (wt./wt.) and 0% (wt./wt.) to 4.4% (wt./wt.), respectively. Irgacure 819 and Tinuvin 171 concentrations were held constant at 2.2% (wt./wt.) and 0.1% (wt./wt.), respectively. 2.2. BVS fabrication The BVS design was created with SolidWorks CAD software. The unit cell used in this study was designed to have increased rigidity rather than flexibility. Fig. 1(b) shows a schematic of the scaffold unit length cell design. Strut width and thickness (radial thickness) were set at 150 mm. The CAD model of scaffold design in the STL format was digitally sliced into layers for additive manufacturing. These layers were then converted into bitmap cross sectional images via a custom Matlab program. 3D fabrication of scaffolds was performed using a custom-made mCLIP system that utilized UV light (l ¼ 365 nm) for photopolymerization and ambient air as the oxygen source. Fig. 1(b) shows a few example cross-sectional images of our fabricated scaffolds as well as an illustration of our scaffold fabrication with mCLIP. The dynamic mask generator for projection of scaffold cross sections was a 1080p digital micro mirror device (DMD, Texas Instruments). Projection optics of the printer were optimized to have a pixel resolution of 7.1 mm  7.1 mm

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Fig. 1. (a) BVS design schematic (unit length) with intended dimensions. (b) Illustration of mCLIP fabrication of B-Ink scaffolds. (c) SEM micrograph of 3D fabricated scaffold.

at the focal plane, which correspond to the maximum build volume for the system of 7.67 mm  13.69 mm  40 mm. All scaffold designs were sliced with a layer slice thickness (LST) between 5 mm and 15 mm. Fig. 1(c) shows a SEM image of 3D fabricated BVS with measured strut width within 3.6% deviation of design thickness. 2.3. Speed calibration and power density optimization The maximum light power value of our system at the focal plane was 6.04 mW with a power density of 17.055 mW/cm2 across a 5 mm diameter circular cross-section. For each B-Ink formulation, speed sweep tests were performed to determine the maximum speed and light intensities needed to accomplish a target dimensional accuracy. Speed sweep tests yielded “speed working curves”, which are discussed in the Theory section. During each speed sweep test, UV light power density (intensity) was held constant while the build platform speed was varied. As the test structure was fabricated, the build platform's vertical speed was incrementally increased in command by 2.5 mm/s after each ridge (Fig. 2(a)). Speed sweep structures were fabricated at various light power densities ranging from 1.543 to 7.50 mW/cm2 (power values of 0.6e2.67 mW, respectively). The initial fabrication speed for each speed sweep structure with UV power density below 4.03 mW/cm2 was 5 mm/s (average speed ¼ 4.22 mm/s). For speed structures fabricated at UV power density equal to or greater than 4.03 mW/ cm2, the initial fabrication speed command was set at 12.5 mm/s (average speed of 8.4 mm/s). Speed values hereafter will refer to the average speed of the full part print, as there was a discrepancy of the input command and the actual average speed of the print. The 3D printed speed sweep structures were measured using scanning electron microscopy (NOVA600, FEI) to quantify the dimensional accuracy, which was defined when the ridge thickness was measured to be within 5% tolerance of the 100 mm design thickness. Maintaining dimensional accuracy ensured the average curing

depth to be within 5% tolerance of the designed value of LST. For LST values of 10 and 15 mm, an UV power sweep was also performed accordingly. For this test, a constant upward speed (for our purposes, the maximum average stage speed) was utilized for the full print and power density was incrementally lowered from full UV power density (17.055 mW/cm2) down to 0.4255 mW/cm2. Ridge thicknesses and working curves were analyzed to determine dimensional accurate UV exposure and stage speeds. 2.4. BVS post-fabrication-processing BVSs after fabrication were further cured via either UV flood exposure or convection heating. All scaffolds used to test effects of B-Ink constituent concentrations were post-fabrication-cured in a UV flood exposure system (Inpro F300S) for 2 min on each side. BVSs were placed 200 distance from the UV source. The Inpro F300S UVB range (280e320 nm) was rated to be as high as 105 2 mW/cm2 and UVA range (320e390 nm) as high as 745 mW/cm2. For heat post-fabrication-processing, scaffolds were placed in a vacuum oven at 110
C for 2, 4, 8 and 10 h. To analyze UV post-fabricationprocess conditions, scaffolds were placed in the UV flood exposure system for 15, 30, 60, 120, 180, and 240 s. A BVS with no postfabrication-processing was also used as a control for both postprocessing methods. 2.5. Mechanical testing All BVSs were soaked in water for at least 12 h to partially simulate their operating environment. Immediately prior to testing, scaffolds were removed from water and excess water was wiped off to prepare them for parallel plate compression test. The BVSs were then loaded into a universal testing machine (Sintech 20G) and compressed to 75% of their original diameter. The experimentally measured force-strain curve was then analyzed to determine the

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Fig. 2. (a) CAD design of speed sweep structure. (b) SEM of mCLIP fabricated speed structure. (c) Speed Working Curves for B-Ink consisting of 60% mPDC and 3% EDAB.

stiffness of the 3D printed BVS. BVSs were fabricated with each of the various B-Ink formulations at the corresponding dimensional accurate UV exposure conditions to evaluate whether faster speed and higher input UV power density affected mechanical properties of the scaffolds, while maintaining the dimensional accuracy. 2.6. Degradation test 3D printed BVSs were placed in a test tube containing 10 ml of 0.1 M NaOH to evaluate accelerated degradation. BVSs were incubated at 37
C, washed with water and dried under vacuum overnight at each measurement time point. Weight loss was calculated by comparing the initial weight (W0) with the weight measured at reported time point (Wt) (Eq. (1)). The results are shown as standard ± standard deviation (n ¼ 3).

Weight loss ð%Þ ¼ 100*ðW0  Wt Þ=W0

(1)

2.7. Biocompatibility evaluation 3D printed BVSs were incubated in series of Ethanol solutions (100% e 20 min, 70% e 20 min and 50% e 30 min) and basal media (24 h) to remove any excess uncured monomers. Next, they were incubated in complete cell culture media at 37
C and media were collected at 24, 48 and 72 h. Human umbilical vein endothelial cells (HUVECs, Lonza) were cultured in endothelial growth medium-2 (EGM-2, Lonza) using standard cell culture practice. HUVECs were seeded in 24 well plates (20,000 cells/well) and cultured with the extracted cell media from BVS incubation for 3 days. Analysis of cell proliferation and cytotoxicity was performed via Alamar Blue assay (Fisher Scientific) by following manufacture's protocol. The cells were incubated with alamar blue dye (440 mmol) and fluorescence was measured using fluorescent spectrophotometer. 3. Theory CLIP can be considered as the 3rd generation of stereolithography process, where scanning and projection stereolithography are 1st and 2nd generation, respectively. It distinguishes from projection stereolithography process in that it incorporates an oxygen permeable window below the fabrication ink. The presence of oxygen in the reaction region creates a small

pocket where the photopolymerization reaction is inhibited near the window, also known as the “dead zone”. During the printing process, the stage travels upward and the “dead zone” is replenished with fresh photopolymer, allowing for a continuous part fabrication [27]. A standardized calibration method is critical to maintain the dimensional accuracy of 3D printed parts. While the lateral resolution is determined by the projection optics, the vertical resolution of the 3D printing process is governed mainly by the absorbing characteristic of the photocurable ink. We have developed a speed working curve method dedicated to the calibration of the mCLIP process that builds the 3D structures via the continuous Z-motion and ties Curing Depth directly to the stage speed. This allows the user to print a single structure for calibration means. The calibration structures were designed and fabricated to characterize the effective curing depth of the B-Ink to ensure the dimensional accuracy. The well-established model for curing depth (CD) in stereolithography processes follows [33].


 CD ¼ Dp *ln Eexp E C

(2)

where Eexp is the total energy density input by the projected UV light onto the surface of the photocurable ink (units of mJ/cm2), Dp is the penetration depth of the UV light (units of mm), and EC is the critical energy density necessary for polymerization to occur. Dp is determined as the slope of the semi-log plot of the CD vs Eexp (also known as the material “Working Curve”), and EC is the X-axis intercept of the working curve [33]. While this equation is generally written in terms of energy flux for conventional SLA processes, based on the static exposure of the building layers, the mCLIP process employs nearly continuous upward motion and thus, this equation needs to be rewritten as the function of stage speed and UV power density. The curing depth equation can be written in terms of speed with a few assumptions: Total energy density (Eexp) is defined as UV light power density at resin surface (Pf) multiplied by exposure time (texp):

Eexp ¼ Pf *texp

(3)

Exposure time (texp) is defined as layer slice thickness (LST) divided by stage speed (vs):

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texp ¼ LST=v

(4)

s

accuracy of the 3D printed part. The OCL sets the stage for quantitative evaluation of the mechanical strength by removing the influence from the variation of the BVS strut dimension.

The critical energy density (Ec) is dependent on both the critical UV light power density (Pc) and the critical stage speed (vc).

4. Results and discussion

Ec ðP; vÞ ¼ ðPc *LSTÞ=v

4.1. Validation of the speed working curve method

c

(5)

The critical value in Eq. (5) is dependent on which user controlled value (UV power density or stage speed) is held constant. The standard working curve equation can then be expressed as the function of the stage speed under constant user controlled UV light power density Pf. Here were refer it as the speed working curve equation:


C D ¼ DP  ln vc=v s

(6)

The critical speed of polymerization vc (analogous to critical energy density in Eq. (2)) is the X-intercept of the new speed working curve, which represents the critical maximum speed at which polymerization can occur under the given Pf value. The speed working curve equation (Eq. (6)) is most useful in performing a speed variation test to determine the dimensionally accurate stage speed. Alternatively, under constant user-controlled stage speed (vs), the standard working curve can be defined as the function of the UV Pf:


 . C D ¼ Dp  ln P f P c

(7)

In this case, the “critical value” of polymerization is the UV power density (Pc). This is used when an exact stage speed is desired, such as our system's maximum stage speed(s) in our study. Because vs is constant, the effective exposure time is also constant, meaning that the energy input is only dependent on the UV light Pf. Eq. (7) is most like the standard curing depth tests performed for traditional stereolithography machines. The CAD design of the calibration structure is illustrated in Fig. 2(a). The calibration structure comprises a series freestanding ridges being attached to a thick post. The ridges are designed to have thickness of 100 mm and vertical spacing of 500 mm to minimize the interference between neighboring ridges. The ridges within the same calibration structure were fabricated under varying stage speeds and thus, we can extract multiple data points from one 3D printed calibration structure to determine the speed working curve. The scanning electron microscope (SEM) image of a representative mCLIP fabricated calibration structure is shown in Fig. 2(b). Fig. 2(c) shows experimentally measured working curves of the B-Ink formulation consisting of 60% mPDC and 3% EDAB. Ridge thicknesses that had been measured during speed sweep tests performed for each B-Ink formulation were averaged by the number of projected layers in each ridge to obtain average curing depth for our speed working curves. Each of the curves plotted in Fig. 2(c) represents a speed sweep at a given Pf value. Line fit performed on the working curves showed good agreement with Eq. (5) (average R2 value across all tested B-Ink formulations and Pf values was 0.972). The dimensional accuracy condition was defined where the curing depth matched the LST of 5 mm, which is represented in Fig. 2(c) as the red horizontal line. The dimensional accurate stage speed at given Pf value is then determined where the speed working curve intercepts the red line. The dimensionally accurate stage speeds under varying Pf value can then be collected to define the optimal curing line (OCL) in the stage speed (vs) and Pf plot. In this study, OCL represent the experimentally determined dependence between vs and the Pf, while maintaining the dimensional

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The speed working curve was developed to allow for fabrication of BVSs at the maximum allowable speed, while maintaining dimensional accuracy. Although the motorized stage is intended to provide a continuous motion, additional computational overhead for loading and processing the bitmap mask patterns cause small time delay in between the projected layers. Thus, the actual stage speed also depends on the LST, which determines the number of mask layers need to be processed per unit time. In our system, the maximum stage speed (vmax) was measured to be 17.77 mm/s for 5 mm LST. As discussed in the theory section, the OCLs are obtained from the speed working curves of the various tested B-Ink formulations. Fig. 3(a and b) show the OCLs of the experimentally measured speed working curves with plotted with UV Pf as the independent variable and log vertical speed (vs) as the dependent variable. As shown in Fig. 3(a and b), the red horizontal lines represent the vmax of the stage at 5 mm LST. The Optimal Pf, which corresponds to the dimensionally accurate Pf at the maximum stage speed (vmax) for a given LST, were thus determined as the intercept of the OCLs and the vmax (red horizontal line). Fig. 3(a) shows OCLs for B-Ink formulations composed of 50%, 60%, and 70% mPDC and constant 3% EDAB by weight, while Fig. 3(b) shows the OCLs for BInk formulations composed of 0%, 1%, 2%, and 3% EDAB and constant 60% mPDC by weight. Fig. 3(c) and (d) show the Optimal Pf vs [mPDC] and Optimal Pf vs [EDAB] plots obtained from the optimal curing speed curves. The Optimal Pf noticeably decreased with increasing mPDC concentration. This was to be expected, as an increase in the volume fraction of prepolymer concentration results in much more effective photopolymerization. Optimal Pf for 50%, 60% and 70% mPDC B-Ink were measured to be 6.621, 4.644, and 2.670 mW/cm2, respectively (Fig. 3(c)). In contrast, EDAB concentration had a smaller influence on the Optimal Pf compared to mPDC concentration, with the Optimal Pf varying from 3.741 mW/ cm2 to 5.293 mW/cm2 (Fig. 3(d)). The presence of EDAB allows a slight lowering of the Optimal Pf, but does not show a large apparent dependence with respect to EDAB concentration. EDAB is typically used as a co-photoinitiator (amine synergist) for Type 2 visible light photoinitiator systems, such as camphorquinone in dental resins. Typical exposure time used in evaluating dental resins containing EDAB is on the order of 30 s to a minute at UV light power densities much higher than that in our printer [34e36]. It was our intention for EDAB to be utilized as an additional cross-linker for post-fabrication-processing of the BVSs. With the effective exposure time of each fabrication layer being shorter than 2 s at maximum and at a shorter light wavelength (365 nm), it is reasonable that the EDAB would largely be inactive during the printing process. The contribution by the EDAB during the 3D printing step is relatively small as expected. By maintaining conditions for dimensional accuracy, the resulting mechanical stiffness of the 3D printed BVSs should mainly be determined by the material properties. The energy variation test was performed on the B-Ink formulation consisting of 60% mPDC and 2% EDAB. mCLIP process parameters (Pf, vs), as well as final postfabrication-cured BVS Stiffness are summarized in Table 1. BVSs were post-fabrication-cured under UV flood exposure for 2 min. BVS Stiffness refers to the slope of the parallel plate compression loading plots. Analyzing the full population of scaffolds printed with the 60% mPDC and 2% EDAB formulation of B-Ink yields the

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Fig. 3. (a) Optimal Curing Lines vs UV Pf, with respect to [mPDC] (b) Optimal Curing Lines vs UV Pf, with respect to [EDAB]. (c) Optimal Pf (intercept points of OCL and vmax) vs. [mPDC]. (d) Optimal Pf vs [EDAB].

Table 1 Effect of ideal curing speed on BVS stiffness for B-Ink formulation of 60% mPDC and 2% EDAB. Pf (mW/cm2)

vs (mm/s)

BVS Stiffness (N/(mm/mm))

2.67 3.32 3.60

13.57 15.40 17.16

2.51 ± 0.21 2.76 ± 0.05 2.62 ± 0.05

standard deviation less than 5%. This relatively small standard deviation underlies the effectiveness of the reported process calibration procedure based on the speed working curve method in determining the optimal process condition with highly repeatable performance of the 3D printed devices. The method for optimizing the printing parameters, such as the stage speed (vs) and UV light power density (Pf), is crucial for both fine part quality and fast printing times. By tying the stage speed directly to the Curing Depth allows the user to essentially disregard the oxygen dead zone thickness. This calibration method works well with ambient air as the oxygen source and relatively low UV power at the focal plane. Using the speed working curve method reported here, one can conduct quick and accurate curing depth tests to optimize printing parameters. BVSs printed with optimized vs and Pf pairs following the OCL all had comparable strength, further validating the effectiveness of the speed working curve method. During the calibration process of various B-Inks, it was observed that the stage motion acted as a constraint for our fabrication. Data shown in the preceding paragraphs was collected from 3D fabricated parts that utilized 5 mm LST. Use of 5 mm LST allowed for full BVS fabrication in 26.5 min with a maximum stage speed, vmax, of 17.77 mm/s. This new speed represents a 62.1% reduction compared to what was reported in our previous work [28]. The vmax was found to be dependent on LST, mostly due to the computational

overhead of converting, loading the bitmap layer masks to the DMD control board, and a pause of 1 ms to ensure projection and motion synchronization for every iterating layers. By utilizing a coarser LST, BVS fabrication speed could be further increased as a higher vmax could be achieved. vmax was found to be 34.35 mm/s and 49.9 mm/s for 10 and 15 mm LST, respectively. These speeds allowed for full BVS fabrication in 15.3 min and 11.3 min, respectively. The significantly shortened fabrication times allowed by the coarser LST make ondemand BVS fabrication possible [1]. While coarser layer slicing reduces the fabrication time, there is a tradeoff with the surface finish. Vertical faces such as the inner diameter and outer diameter of the BVS are largely unaffected by changes in LST, but angled structures, such as the struts, begin to show evidence of “stair stepping” or layering as LST increases. Further work is ongoing to make part fabrication with fully continuous motion, which would negate the low speed limits for small LST. Optimization of ink properties for mechanical strength reported in the following sections is reported for 5 mm LST, as this maintained very high resolution. 4.2. Effects of post-fabrication-processing method on BVS radial stiffness Most stereolithographic parts utilize post-fabrication-processing or post-fabrication-cure to obtain their final mechanical properties [37]. The two most common post-processing methods are UV flood post-fabrication-cure and convection heat post-fabrication-cure. It was observed that both the post-processing method and the duration of each method had large effects on BVS stiffness. Initial, unprocessed, scaffolds were very compliant, with an average stiffness value of 0.48 N/(mm/mm). All scaffolds used in these postfabrication-cure validation tests were fabricated using the B-Ink formulation consisting of 60% mPDC and 3% EDAB. Fig. 4(a) shows

H.O.T. Ware et al. / Materials Today Chemistry 7 (2018) 25e34

the effect of convection heat post-fabrication-cure on the BVSs at 110
C. The heat post-fabrication-cure method led to the tested maximum value of 3.83 N/(mm/mm) although for practical purposes, the heat post-cure can be considered to saturate at approximately 4 h with an average stiffness value of 3.57 N/(mm/mm). Using convection heat post-fabrication-cure, BVS stiffness exceeds the nitinol control stent (red line) at the post-fabrication-curing time beyond 1.5 h. Fig. 4(b) shows BVS stiffness with respect to UV postfabrication-cure exposure. BVS stiffness values increased with time, but the rate of increase reduced significantly at approximately 60 s of exposure. While BVS stiffness continuously increased with time under UV post-fabrication-cure, only 60 s was necessary for the B-Ink scaffolds to have stiffness values that approximately match the stiffness of the nitinol control stent (red line). The maximum stiffness values of the BVSs fully post-fabrication-cured were greater than the stiffness value of the nitinol control (2.574 N/(mm/mm)). Maximum BVS stiffness obtained from the heat post-fabrication-processing was 3.832 N/(mm/mm). Within the tested time scale (0e240 s), the maximum BVS stiffness obtained from UV post-fabrication-processing was 3.395 N/(mm/ mm). From the results in both Fig. 4 plots, we concluded that heat post-fabrication-cure typically yielded stiffer scaffolds, but took much more time to reach saturation. UV post-cured scaffolds effectively saturated in stiffness at 120 s of exposure. The much slower increase shown after 60 s appears to be dependent on the penetration of the UV light within the UV flood device. After 60 s, high crosslink density at the BVS exterior surfaces could inhibit further cross linkage of interior polymer chains. The postfabrication-cure method can be chosen depending on the application. For potential on-demand fabrication of BVSs, where lead time is understandably short, UV post-fabrication-cure can be utilized. Whereas if traditional lead time for the stent/scaffold diagnosis and implantation is needed, the heat post-fabrication-cure method is a valid option. In the literature, heat post-fabrication-cure was shown to have higher degree of elongation at failure than UV postfabrication-cure, which ultimately makes it the better option for traditional lead times [37]. For BVSs fabricated with 150 mm strut width and strut thickness, full polymerization through either postprocessing method lead to stiffness of the 3D fabricated BVSs exceeding the stiffness of the nitinol control stent. For comparison, in our previous study, B-Ink consisting of 50% mPDC and 0% EDAB present required 400 mm strut thickness and 200 mm strut width to achieve similar stiffness to the nitinol control stent [28].

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4.3. Effects of [mPDC] and [EDAB] on BVS radial stiffness after the post-curing process BVSs with UV post-fabrication-cure of 2 min were mechanically tested via parallel plate compression to determine the effects of mPDC and EDAB concentrations on final BVS mechanical properties. mPDC and EDAB concentrations were both observed to have large effects on the final mechanical properties of printed BVSs. To understand the effect of mPDC concentration, the EDAB concentration was held at 3% for these B-Ink formulation variations. Fig. 5(a) relates the stiffness in parallel plate compression of fabricated BVSs to mPDC concentration. Per standard protocol, all BVSs were UV post-fabrication-cured for 2 min on each side. Within the region of 50% and 60% mPDC the stiffness increases with increasing mPDC concentration. This is expected, as higher prepolymer volume fraction could favorably strengthen the photopolymer, which results in the increase in the stiffness. However, the decrease in stiffness from 60% to 70% mPDC concentration is rather unexpected. We believe the observed phenomena might be associated with the increased viscosity and its effect on the printing process, which will need to be investigated further. This is unique to the mCLIP at relative high stage speed. The horizontal red bar in Fig. 5(a) and (b) reflects the stiffness measured for the nitinol control stent (2.58 N/(mm/mm)). BVSs printed using B-Ink containing 55% and 60% mPDC were the only concentrations that exceeded the stiffness of the Nitinol stent. BVSs fabricated with 60% mPDC had a stiffness value equal to 3.08 N/(mm/mm). When testing the effect of EDAB concentration on final BVS stiffness, the mPDC concentration was held constant at 60%. Although EDAB did not have a large effect during the printing process, Fig. 5(b) shows that EDAB concentration had a large effect on the final BVS stiffness after UV post-fabrication-curing. As the final BVS geometry was exposed to relatively long periods of UV light compared to each individual fabrication layer, this allowed the EDAB present within the ink to be activated and increased strength was obtained. BVSs made without EDAB were the weakest of the tested scaffolds, whereas those with 3% EDAB had the maximum stiffness values within the tested concentrations (3.08 N/(mm/ mm)). As EDAB concentration increased from 0% to 3%, the BVS stiffness increased linearly but this trend peaked at 3%. The tested concentrations of EDAB were within typical EDAB usage concentrations as a co-photoinitiator (between 0.5 and 5%) [38]. EDAB concentration of approximately 1.9% and 3.7% gives comparable stiffness to the nitinol control BVS. Following the EDAB and mPDC

Fig. 4. (a) Convection heat post-fabrication-cure effect on BVS stiffness. (b) UV post-fabrication-cure effect on BVS stiffness. Scaffolds tested had strut width and strut thickness of 150 mm.

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Fig. 5. (a) Effect of [mPDC] on 3D fabricated BVS stiffness. (b) Effect of [EDAB] on 3D fabricated BVS stiffness. Values shown are slopes of parallel plate compression tests. Epsilon is the compressive strain (mm/mm) of the BVS during testing.

optimizations, we determined that the B-Ink formulation resulting in the strongest scaffolds was 60% mPDC and 3% EDAB. EDAB and mPDC content had significant effects on the final BVS's mechanical properties after post-processing. While high mPDC concentrations allow for curing at low input energy, this didn't necessarily lead to higher mechanical strength. EDAB showed a similar trend within the realm of final BVS stiffness. The increase in BVS stiffness is expected, as the combination of Irgacure 819 and EDAB had previously been shown in the literature to increase the degree of polymerization [34]. The addition of EDAB and a 10% increase of mPDC relative to the ink reported in our earlier work (50% mPDC and 0% EDAB) has imparted greater mechanical properties to the printed BVSs. This newer B-Ink formulation of 60% mPDC and 3% EDAB allowed a reduction of strut thickness from 400 mm to below 200 mm with comparable stiffness to nitinol stents [28]. Ultimately, the introduction of EDAB as an additional crosslinker allowed this reduction in strut thickness. In addition, EDAB has been utilized in dental resin systems, indicating at least reasonable biocompatibility [34e36]. This addition, as well as the change of the diethyl fumarate solvent to ethanol, has allowed key reductions of the scaffold profile and increased its viability as an implantable device. It is worthwhile to note that this study is focused on one particular scaffold design optimized for high rigidity, which allowed us to optimize the process parameter as well as the B-ink compositions. Future study of the scaffold design should emphasis more on the proper balance of flexibility and rigidity to properly fit complex artery geometries. 4.4. Degradation and biocompatibility evaluation In vitro degradation kinetics of photopolymerized mPDC polymer films had previously been investigated with Irgacure 1173 and Irgacure 819 as the photoinitiator systems in 37
C PBS [28,32]. Both investigations suggested that ~25 and 26% of the polymer degrades in 6 months for Irgacure 819 and Irgacure 1173 photoinitiator systems, respectively. Our current iteration of the ink now incorporates EDAB in addition to Irgacure 819, which could affect potential degradation properties. In this work, a base accelerated hydrolytic degradation study in 0.1 M NaOH was performed to prove complete degradation of mPDC based 3D printed BVS occurs (Fig. 6). NaOH has previously been used in various concentrations and temperatures to accelerate hydrolytic degradation of polymers in several studies [31,39e41]. Fifty percent of the samples were degraded at 10 h and complete degradation occurred at 27 h.

Fig. 6. Accelerated degradation test of BVS. Mass loss (%) was normalized based on the initial weight of the samples. (n ¼ 3).

Biocompatibility of BVS incorporating Irgacure 819 and EDAB as the photoinitiator system was investigated using a model cell type, HUVECs. The cells treated with BVS extracts showed similar proliferation rate to control group (fresh media) at 3 days (Fig. 7). There was not any large statistical difference between control versus treated groups and among different treatments (24 h-72 h extracts). The previous iteration of the B-Ink formulation utilized solvents and UV absorbers which were not biocompatible [28]. There does not appear to be any cytotoxicity concerns regarding BVS and corresponding B-Ink based on these findings. 5. Conclusions In this study, a methodology for optimizing a high-resolution and high-speed 3D printing process is presented. The speed working curve method allowed for fast, in-process calibration of the mCLIP process parameters at which point a quantitative evaluation of different B-Ink formulations was carried out. The method presented allowed for 3D printing of BVSs with tight manufacturing tolerances with a variety of material concentrations. mPDC concentration was shown to influence the in-process curing of the BInk material, while EDAB concentration was found to have rather

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33

References

Fig. 7. Proliferation of HUVECs when they were incubated in fresh media or BVS extracts from day 1e3. Green dashed line indicates initial cell seeding number. (n ¼ 3 per group). (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article).

slight effect during printing process. However, both mPDC and EDAB were shown to have a large effect on BVS strength following post-fabrication-cure processing. Introduction of the cophotoinitiator EDAB allowed greater cross-linkage during the post-process step, which ultimately made possible strut geometries comparable to that of commercial polymer stents, such as the Absorb GT1. While UV post-fabrication-curing allowed quick postprocessing to achieve final properties for potential on-demand printing, heat based post-fabrication-curing was shown to be useful in situations where larger lead times for fabrication are acceptable. At 5 mm LST, BVSs were fabricated in 26.5 min whereas for 10 mm and 15 mm LST, BVSs were printed in 15.3 min and 11.3 min, respectively. Further, BVSs printed with the optimized BInk formula were shown to have radial stiffness exceeding that of nitinol stents. This work validates the possibility for the use of customized, 3D printed vascular scaffold for patients in emergency settings.

Acknowledgements This work was supported by the National Science Foundation (NSF) under Grant No. EEC-1530734. The development of the mCLIP system is supported by a generous donation from The Farley Foundation. Henry Oliver T. Ware would like to acknowledge the National Science Foundation Graduate Research Program as he is a recipient of the fellowship. This work used Northwestern University's Central Laboratory for Materials and Mechanical Properties to conduct mechanical testing. This work utilized Northwestern University Micro/Nano Fabrication Facility (NUFAB), which is partially supported by Soft and Hybrid Nanotechnology Experimental (SHyNE) Resource (NSF ECCS-1542205), the Materials Research Science and Engineering Center (DMR-1720139), the State of Illinois, and Northwestern University.

Appendix A. Supplementary data Supplementary data related to this article can be found at https://doi.org/10.1016/j.mtchem.2017.10.002.

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