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A new process for customized patient-specific aortic stent graft using 3D printing technique
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A new process for customized patient-specific aortic stent graft using 3D printing technique Yang Lei a,b,#, Xin Chen a,c,#, Zhen Li a, Lei Zhang a,∗, Wei Sun a, Lei Li d, Feng Tang d a
Department of Mechanical Engineering, Tsinghua University, Beijing 100084, China Beijing Institute of Aeronautical Materials, Beijing 100095, China c China Electronics Technology Group Corporation, Beijing 100846, China d The First Hospital of Tsinghua University, Beijing 100016, China b
a r t i c l e
i n f o
Article history: Received 5 January 2019 Revised 14 November 2019 Accepted 10 December 2019 Available online xxx Keywords: 3D printing Aortic dissection Personalized aortic stent graft Rapid prototyping sacrificial core-coating forming technique
a b s t r a c t Endovascular aneurysm repair (EVAR) is a popular and effective treatment for descending aortic disease. However, the majority of existing coating stents used in EVAR are of a standard design which may not meet the size or structural requirements of different patients. Therefore, in this paper, we propose using 3D printing and controlled deposition as a patient-specific aortic stent graft manufacturing technique. The methodology involves the use of a rapid prototyping study sacrificial core-coating forming (RPSC-CF) technique to develop an aortic stent graft that consists of a film and metallic stent. Polyether polyurethane and nickel-titanium alloys were chosen due to their shape memory properties and good biocompatibility. The resulting customized stent grafts meet the demands of personalized therapy and invasive surgery, and perform well as demonstrated from burst pressure testing and the degree of radial support provided and radial support force tests, laying the foundation for precise aortic dissection treatment. © 2019 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction Endovascular aneurysm repair (EVAR) is the primary method of treating aortic dissection [1]. In the EVAR process, a stent graft is implanted into the lesion to close the rupture of the aortic intima in order to reconstruct the aortic wall. The blood is blocked from the false lumen in case the false lumen further expands or ruptures [2]. Existing methods for forming aortic stent grafts rely on standardized manufacturing processes, producing stents with a fixed diameter and length [3]. However, along the axial direction when moving from proximal points to more distal areas, the diameter of the human thoracic descending aorta decreases. Therefore, standard diameter stents cannot meet personalized size and shape requirements [4].
Abbreviations: EVAR, endovascular aneurysm repair; RPSC-CF, “rapid prototyping” and “sacrificial core” and “coating” forming; PET, polyethylene terephthalate; e-PTFE, expanded polyethylene terephthalate; Ni-Ti alloy, nitinol; CT, computed tomography; PU, polyurethane; FDA, The Food and Drug Administration; Af, austenitetransition temperature; PVA, polyvinyl alcohol; HG-stained, histologicalgradingstained. ∗ Corresponding author. E-mail address: [email protected] (L. Zhang). # These authors contributed equally to this work.
Stent placement that does not correspond to the specific size and structure of the patient’s aorta may affect the adherence of the stent [5–7]. A stent graft diameter that is less than the patient’s aortic diameter will lead to poor stent adhesion, which can result in blood leakage or stent displacement. A stent diameter greater than the patient’s aortic diameter will cause the vessel to break, resulting in treatment failure. Additionally, the actual structure of the human thoracic aorta is not straight but curves; consequently, some aortic stent scaffolds cannot fit to the curved structure of blood vessels, which causes deflation and treatment failure. The branched structure of the aortic arch presents great challenges for surgical treatment. In current clinical practice, the open window method is usually adopted for the treatment of branch vessels, and the stent graft is processed and connected during surgery. This kind of surgery requires the joint implantation of multiple stent grafts, which is extremely risky. Therefore, there is a clinical demand for bending and branch customization of aortic stent grafts. At present, the most common stent graft on the market consists of a polymer film material like polyethylene terephthalate (PET) or expanded polyethylene terephthalate (e-PTFE) and a nitinol alloy [11,12]. Biopolymer materials are the primary materials for the film. Polymer film forming processes are mainly divided into textile forming and non-textile forming techniques. Textile forming
https://doi.org/10.1016/j.medengphy.2019.12.002 1350-4533/© 2019 IPEM. Published by Elsevier Ltd. All rights reserved.
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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uses a loom to weave polyester fibre into a tubular film. Because of its compact structure, stable size, and controllable thickness, this method is widely used by stent graft manufacturers. Non-textile forming methods mainly include extrusion forming, electrostatic spinning, and micro-nano-composition. The metal components of stents can be made from stainless steel, cobalt-chromium alloy, or nickel-titanium alloy, according to the surgical release methods to be used [8–10]. Metal stents are structured into ‘W’ or ‘Z’ shapes formed using laser cutting or weaving. The metal stent and film are then bound by methods including stitching and weaving. To a degree, weaving and stitching methods can meet the requirements of personalized manufacturing. Using a model of an actual aorta, Wang et al. [13] formed a polyester film using a suturing method to ensure similarity to the true structure of the aorta; then, a stereotyped nitinol (Ni-Ti alloy) stent was adhered to the film to prepare the personalized aortic stent. Zhang et al. [14] utilized a machine weaving method to weave film onto a nitinol stent in order to achieve customization. Cook Medical LLC can customize stent grafts, but its production cycle requires approximately one month. Patients with acute aortic dissection need to be treated promptly; after diagnosis, the mortality rate of acute aortic dissection is as high as 36–72% within 48 h and increases to 62–91% within one week [15]. Therefore, the one-month production period cannot guarantee timely treatment for patients with acute aortic dissection. For these reasons, the idea of using 3D printing to manufacture personalized stent grafts has been advanced [16,17] in order to produce a personalized stent graft that matches the true shape of a patient’s aorta in a very short time frame. 3D printing technology can transform a digital model directly into a solid sample, allowing for inexpensive and highly efficient customization. This technique not only ensures better treatment outcomes due to the stents’ longitudinal flexibility and compatibility with the vascular wall, but also avoids the disadvantage of the long production cycle of current stent grafts, indicating that this technique has good prospects for clinical application. At present, no research has applied 3D printing technology to the manufacture of stent grafts. 3D printing has been primarily used as an assistive technique for interventional treatments. Valverde et al. [18] used 3D printing to build a model of a patient’s abdominal aortic aneurysm in order to select the most suitable stent-graft system and simulated the graft implantation in vitro to improve the accuracy of the interventional therapy. Because 3D printing technology allows for rapid personalization, we wanted to apply 3D printing technology to the manufacture of stent moulding. However, it is difficult to use 3D printing technology to directly integrate and form the composite materials that comprise the covering film and metal stent. To address this problem, a sacrificial core-coating forming technique first proposed by Zhang [19] was used: the Rapid Prototyping Sacrificial Corecoating Forming (RPSC-CF) technique involves the formation ofTo solve this problem, this project adopts the rapid prototyping sacrificial core-coating forming (RPSC-CF) technique based on 3D printing technology proposed by Zhang . RPSC-CF employs the formation of a water-soluble non-toxic solvent core using rapid prototyping to ensure the stent lumen structure. The multilayer tube wall structure can then be formed layer-by-layer using coating, spraying, or other material deposition processes, and finally, the inner core is dissolved to obtain the stent. Using RPSC-CF technology, 3D printing is used in this study to print customized aortic stent cores with soluble materials, then the core is coated with polymer multiple times to form a multi-layer film. Next, a conventional metal stent is wrapped between the two layers of film to integrate the alloy stent and the film. This study investigates materials for the film, metal stent, and water-soluble sacrificial core, as well as the stent formation process, ultimately
developing a technology for forming and manufacturing personalized stent grafts. 2. Methods and materials 2.1. Stent graft fabrication Preparation of a personalized stent grafts should include two primary considerations: ensuring that the stent accurately matches the true structure of the patient’s aorta and, because the stent is composed of film and a metal stent, determining how to combine these two materials stably and effectively. These two aspects are fundamental to the success of the stent graft fabrication. Aortic 3D model restoration was performed with Mimics 10.01 software. Based on computed tomography (CT) data, a 3D aortic model file was obtained after eliminating noise from other tissues and smoothing the surface. The stent core was designed using the SolidWorks 2016 platform. Vascular bending angles can vary due to individual differences, and the diameter of the descending thoracic artery is irregular along the axial direction. To accurately extract size information from the descending thoracic aorta, diameters were taken from the 3D aortic model every 5 mm along the axial direction. Additionally, to ensure effective fixation of the stent graft into the lesion, clinical experience has shown that the core diameter of the stent should be 10% greater, and the top and bottom ends 15% greater, than the diameter of the blood vessel. Existing methods to integrate the metal and film moulding include suturing and deposition. The suture method for stent graft integration uses special stitch and suture techniques to combine the metal stent and film. In this method, it is easy to penetrate or even rupture the film. Moreover, this method’s low efficiency (due to the need for manual sewing) is not compatible with the requirement for rapid preparation [1,20]. In the deposition method, the outer film is coated onto the surface of the inner film of the stent graft. After the solution has volatilized, the metal stent is tightly wrapped by the inner and outer films. This method avoids film infiltration or rupture caused by hand sewing and has higher efficiency than the suture method. Therefore, in this study, the metal stent is combined with the film using the deposition method. The fabrication process depends on RPSC–CF technology (Fig. 1). 2.2. Film material Polyurethane (PU) has been chosen as the preferred material in this study. PU is polymerized by polyol (polyether and polyester), isocyanate, and a chain extender (diol and amine) that contains carbamate in its backbone [21]. Polyether PU produced by Zhejiang Huafeng Chemistry Co., Ltd, has been used as a film in aortic stent grafts. The US Food and Drug Administration (FDA) has certified that this material has good biocompatibility and resistance to hydrolysis. After it was dissolved in tetrahydrofuran and deposited into a film, its biocompatibility was detected in vitro: it exhibited a haemolysis rate of 0.46% and a relative cell productivity rate, obtained from cytotoxicity tests, of 93.5%. Platelet adhesion test was implemented. Human whole blood with anticoagulant was centrifuged for 5 min at 10 0 0r/min and platelet-rich supernatant was obtained. The PU samples were placed in the platelet-rich serum in a cell incubator at 37 °C and 5% CO2 for 1 h. The samples were immersed in glutaraldehyde aqueous solution with a volume ratio of 2.5% v/v for 12 h after rinsing with saline several times and then dried in a vacuum after dehydration of alcohol. After gold spraying treatment, platelet adhesion on the surface of the sample was observed by environmental scanning electron microscopy. No obvious agglomeration
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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Fig. 1. Technical Route.
Table 1 Components of Nitinol. Element wt% Nb <0.005
Ni 55.65 Fe <0.005
C 0.0043 O 0.037
Co <0.005 N 0.0044
Cu <0.005 H <0.0010
Cr <0.005 Ti remainder
core can be rapidly moulded using fused deposition manufacturing techniques. In this paper, PVA produced by Shenzhen Esun Industrial Co. with a 1.75 mm diameter and a printing temperature of 180–210 °C was used for 3D printing. A Raise3D N2 3D printer was used to form the soluble core. The following print parameters were used: a layer thickness of 0.1 mm, a surface thickness of three layers, a high precision surface, and a filling rate of 3%. After printing, the size tolerance of the core diameter was 0.1 mm and the core met the design requirements 3. Results and discussion 3.1. Stent graft design Fig. 2. Platelet adhesion test. No obvious agglomeration or pseudopodium were observed by environmental scanning electron microscopy.
or pseudopodium were observed in the test (Fig. 2). Therefore, the PU film can be used as a biomedical material. 2.3. Metal stent material Nitinol is typically selected as the material for metal stents due to its super-elasticity, shape memory, and biocompatibility [22,23]. The super-elasticity of nitinol provides good compliance with the vessel wall. In the field of interventional treatment, more than 4/5 of products rely on nitinol’s super-elasticity. In this paper, the nitinol stent is shaped by a heat-treatment process. The nitinol wire used in the experiment was provided by the Jiangyin RongBang New Material Technology Co. The wire diameter was 0.30 mm, it had a bright surface, its tensile strength was 1453 MPa, its fracture elongation rate was 14%, and its austenite-transition temperature (Af ) was 15 °C. The alloying components are shown in Table 1. 2.4. Sacrificial core material In this study, we use polyvinyl alcohol (PVA) as the watersoluble sacrificial material. PVA has good biocompatibility. Moreover, PVA is soluble in water and insoluble in most organic solvents such as toluene and tetrahydrofuran [24]. The entire soluble
For the purpose of clinical animal experiments, CT data of Chinese mini-pigs from a Philips Brilliance 64 were provided by the First Hospital of Tsinghua University. Fig. 3 shows the results of patient-specific aortic stent graft design. 3.2. Film moulding The stent uses a vertical dip coating method. According to the requirements of the process, we built a uniform vertical pulling platform (Fig. 4). The vertical motion axis is driven by a stepper motor with a motor step angle of 1.8° (corresponding to a subdivision of 32), and single machine single-turn motion z-axis vertical stroke of 8 mm. The PU film was fabricated using a dip coating technique. When the viscosity of the fluid and the pulling speed are quick enough, the film deposit thickness for flat substrates is related to the viscous drag (ηU0) and gravity (ρ g) [25,26]:
h0 = c1 (ηU0 /ρ g )1/2 ,
(1)
where ρ is the liquid density, g is the gravitational acceleration, η is the fluid viscosity, U0 is the pulling speed and c1 is 0.8 for Newtonian fluids. When the pulling speed is between 1 mm/s and 10 mm/s and the fluid viscosity η is low, surface tension γ LV affects the thickness of the film; formula (1) can be written as follows:
h0 = 0.94
ηU0 2/3 . γLV 1/6 (ρ g)1/2
(2)
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Fig. 3. Stent Graft Design (A) Structure reduction of the descending aorta; (B) For the experimental design of the straight stent graft, in order to ensure the bracket’s good geometric compliance, an inwardly curved ‘pleat’ is designed every 10 mm along the bracket’s axial direction with an axial length of 5 mm, so that the stent can be bent in any direction. This design prevents distortion when the stent graft is bent; (C) Bent aortic stent graft sacrifice core design.
decreases the uniformity of the dip coating. Therefore, we investigated the effects of film thickness to find the optimal coating process. The PVA sacrificial cores used in the experiments had 10 mm diameter cylinders. A pulling speed of 5 mm/s, a solution concentration of 10%, and a two-dip coating process were found to be the optimum dip coating conditions for the inner layer of the film (Fig. 5). Due to the high concentration of the solution, the outer layer of the film was not flat. To rectify this situation, a pulling speed of 5 mm/s, a solution concentration of 5%, and two-dip coating process found to be optimal for the formation of the outer film layer. 3.3. Nitinol stent moulding In this study, the nitinol stent was shaped by heat treatment. The stainless steel mould was produced with a high-speed computer numerical control milling process; the current mould production time is approximately 2 days. When including the other stent graft fabrication steps, the total production cycle takes 3 days. After shaping, the nitinol stent should have good super-elasticity and exhibit proper radial support. However, the heat-treatment temperature and duration affect the nitinol’s Af and mechanical properties. Therefore, we investigated the appropriate heat-treatment parameters to optimize nitinol stent performance. The temperature at which nitinol is heat treated is generally between 460 °C and 500 °C [27]. Optimal heat treatment was achieved through cooling by quenching.
Fig. 4. Uniform vertical pulling platform for deposition processing.
The PU is dissolved in tetrahydrofuran. After the tetrahydrofuran has been completely volatilized, the PU wraps around the water-soluble core to form the stent graft. To ensure optimal film performance, its thickness should be reduced as much as possible. Furthermore, considering that the stent should be manufactured quickly, a low dip coating time, a high solution concentration, and a relatively high pulling speed should be adopted. However, an excessive concentration increases the solution’s viscosity, which is not conducive to its volatilization. In addition, a high speed pulling
3.3.1. Effect of heat treatment on af At temperatures greater than nitinol’s Af , the nitinol is in a stable austenite state. Under applied stress, a martensitic transformation occurs, which results in an unstable state. The unstable martensitic phase will undergo an inverse phase transformation to a stable austenite phase after stress unloading. Therefore, nitinol is super-elastic on the macro scale [28]. For the aortic stent graft, the Af should be lower than 37 °C to ensure that nitinol exhibits super-elasticity in the human body. In addition, the radial support force of the nitinol stent increases with a decreasing Af . A low Af leads to strong external forces, which may cause vascular damage and hyperplasia. Results from differential scanning calorimetry (Fig. 6A) showed that heat treatment at 480 °C resulted in an Af less than 37 °C and better stability.
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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Fig. 5. Effects of film thickness. (A) When testing solution concentrations (w/v) of 7.5%, 10%, 12.5%, and 15% at a pulling speed of 8 mm/s for 5 dip coats, we found that the film thickness was exponentially related to the solution concentration. (B) When testing pulling speeds of 2 mm/s, 5 mm/s, 8 mm/s, and 10 mm/s at a solution concentration of 12.5% (w/v) for 5 dip coats, we found the film thickness exhibited a logarithmic relationship with respect to the pulling speed. (C) For a pulling speed of 8 mm/s and a solution concentration of 12.5% (w/v), the relationship between film thickness and number of dip coatings was exponential.
Fig. 6. Effects of heat treatment parameters. (A) Effect of different heat treatment conditions on Af . The initial Af was 15 °C. At 460 °C, the Af decreased slightly, then rose rapidly with time. At 500 °C, the Af declined over time. At 480 °C, the Af was approximately 10 °C and was relatively stable; (B, C, D) With an increased heat-treatment time, the stress on the upper and lower platforms tended to decrease, but the residual strain rose. When the time exceeded 10 min, the residual strain was greater than 0.5%, indicating that irreversible plastic deformation had occurred. Therefore, the heat-treatment time should not exceed 10 min, and the heat-treatment temperature should be greater than 460 °C.
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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Fig. 7. Shaped nitinol wire after 20 min, 25 min, and 30 min of heat treatment (left to right).
Fig. 8. Personalized aortic stent grafts. (A) Combined nitinol stent and stent film; (B) Personalized aortic stent graft; (C) Aortic stent graft with branch; (D) Straight aortic stent graft.
3.3.2. Effect of heat treatment on the stress–strain curve The material’s stress–strain curve was tested according to the ASTM F2516-14 standard. The initial gauge length was 10 mm, the travel rate of the testing machine transom was 2 mm/s and the test temperature was 26 °C. The results showed that the upper platform stress, lower platform stress, and residual strain of the original nitinol were 397 MPa, 50 MPa and 0.3%, respectively. Due to the mechanical property changes caused by heat treatment, 12 group experiments were prepared for heat-treatment parameter determination (Figs. 6B, C, and D). The chosen heat treatment conditions were 480 °C for 10 min.
3.3.3. Moulding the stent The stainless steel mould has thermal inertia; after weaving the nitinol to the mould, the shaping effect of the nitinol is difficult to maintain without the mould. Therefore, heat treatment was extended to 20 min, 25 min, and 30 min at 480 °C (Fig. 7). After 30 min of heat treatment at 480 °C, nitinol achieved a good shaping effect. Moreover, under this heat treatment condition, the Af of the Ni-Ti stent was 20 °C and its residual strain was 0.4%, indicating that the stent can exhibit super-elasticity in the human body. Therefore, a 30 min heat treatment at 480 °C was the most suitable for the shaping process.
Fig. 9. Mechanical properties test. (A, B) Radial support force measuring device; (C) Explosive pressure detection device.
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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Fig. 10. Preliminary in vivo test results. (A) Implantation Day 0; (B) Implantation Day 30. Marker 1 to marker 2 10 mm.
3.4. Combining the metal stent and film The key to preparing an optimal stent graft lies in the firm combination of the film and the metal stent. Here, the deposition method was used for this step. The heat-treated nitinol stent was wrapped in the inner coating surface of a 5% PU solution by dipping, and when the solution evaporated, the metal stent was wrapped by the inner and outer membranes (Fig. 8A). Finally, the core was dissolved in water due to the high water-solubility of the PVA material. The stent graft was obtained when the core was fully dissolved (Figs. 8B, C, and D). 3.5. Mechanical properties 3.5.1. Radial support force In this paper, the stress required by reducing the diameter of the stent by half is used as the radial support force. GORE® TAG® Thoracic Endoprosthesis, Medtronic ValiantTM Thoracic Stent Graft, Cook Zenith Alpha Thoracic Endovascular Graft are used as reference standards and their single-ring radial support forces are measured between 0.28 N and 0.34 N. The experimental results showed that the supporting force of the single nitinol ring was 0.3 N, which is equal to the supporting force of an aortic stent graft that has already been clinically applied (Figs. 9A and B). Therefore, the stent graft can support an appropriate radial support force.
3.5.2. Burst pressure Generally, the normal range of human systolic blood pressure is 12.0–18.7 kPa and the normal range for diastolic blood pressure is 8.0–12.0 kPa. Using deionized water with a similar blood density as the pressure medium, the liquid bursting pressure was measured for the stent graft (Fig. 9C) and indicated that the bursting pressure of the personal stent graft was greater than 70 kPa, which is safe and reliable. This result shows that the stent could be implanted into the descending aorta. 3.6. In vivo implantation As the stent graft in this study was found to meet the basic requirements for in vivo implantation, we then conducted in vivo testing. In vivo implantation was carried out on a Chinese minipig in the Department of Vascular Surgery at the First Hospital of Tsinghua University. Preliminary test results indicate that 30 days after implantation (Fig. 10), the stent graft was compliant and effectively supported the descending aorta with no displacement. However, there is a slight atrophy in the middle of the stent graft. In future work, the radial support force of the nitinol in the stent will be reinforced to support to the aorta wall. Blood tests at 24 h, 48 h, and 72 h after surgery show that the leukocyte counts are 19.08 × 109 /L, 19.99 × 109 /L, and
Fig. 11. Pathological sections. (A) Smooth muscle of upper convexity of aortic arch; (B) Smooth muscle of concavity of aortic arch.
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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24.21 × 109 /L (reference range: 10.20 × 109 /L to 30.00 × 109 /L). The percentages of neutrophils at these time points were 35.7%, 26.4%, and 36.2%, respectively (reference range 20% to 60%). The blood test results show that the blood index of the experimental Chinese mini-pig is normal and the inflammatory response is small after the intervention. Histologicalgrading-stained (HG-stained) aortic sections (Fig. 11) show that although there is partial oedema in the smooth muscle tissue on both sides of the aortic arch, the degree of damage is similar, indicating that the support on both sides of the aortic arch is roughly equivalent. Accordingly, the stent graft meets the design requirements. In summary, the stent graft performed well in vivo. Further results on the in-vivo performance will be published in due course. 4. Conclusion Customized stent grafts with proper biocompatibility and mechanical properties were successfully fabricated based on rapid prototyping sacrificial core-coating forming (RPSC-CF) technique. Film was prepared from polyether PU. The film’s thickness was 0.185 mm, increasing the aortic stent scaffold burst pressure to greater than 70 kPa, which is a safe and reliable level for a device to be implanted into the human descending aorta. Besides, the stent’s single ring radial support force was 0.3 N with 0.3 mm nitinol wire. The stent graft can support an appropriate radial support force. Meanwhile, using a real descending thoracic aortic structure, curved and branched stents were manufactured, demonstrating the ability of these materials and production methods to meet the needs of individual patients. For the last, the entire stent graft preparation cycle was 3–4 days. With continued modification of the preparation process, the entire stent preparation cycle is expected to be compressed to within 48 h to ensure that patients with aortic dissection receive timely treatment. This new process for customized patient-specific aortic stent graft provides a new idea for precise aortic dissection treatment. Declaration of Competing Interest None. Acknowledgments The Beijing Municipal Health Planning Commission 2016-2-4121 and the Beijing Life Science Frontier Technology Cultivation Study Z14110 0 0 0 0214012 supported this study. Ethical approval The in-vivo test carried out in this study is in accordance with Chinese National Laboratory Animal Standard and EU Directive 2010/63/EU. References [1] Dua A, Kuy SR, Lee CJ, Upchurch GR, Desai SS. Epidemiology of aortic aneurysm repair in the United States from 20 0 0 to 2010. J Vasc Surg 2014;59(6):1512–17.
[2] Baum RA, Carpenter JP, Cope C, Golden MA, Velazquez OC, Neschis DG, et al. Aneurysm sac pressure measurements after endovascular repair of abdominal aortic aneurysms. J Vasc Surg 2001;33(1):32–41. [3] Anonymous Research and markets offers report: aortic stent grafts - Global Pipeline analysis, competitive landscape and market forecasts to 2017. Wireless News 2011. [4] Kiyama H, Imazeki T, Kurihara S, Yoneshima H. Long-term follow-up of polyurethane vascular grafts for hemoaccess bridge fistulas. Ann Vasc Surg 2003;17(5):516–21. [5] Li ZZ, Xu SD, Liu T, Tao Y, Wang S, Ma LA, et al. Outcome of endovascular stent graft placement in patients with acute thoracic aortic syndromes. Chin J Cardiol 2007;35(6):555–8. [6] Greenhalgh RM. The EVAR trial participants. Comparison of endovascular aneurysm repair with open repair in patients with abdominal aortic aneurysm (EVAR trial I), 30-day operative mortality results: randomised controlled trial. Lancet 2004;364(9437):843–8. [7] Psacharopulo D, Ferrero E, Viazzo A, Ferri M, Ripepi M, Nessi F. Abdominal aortic endograft collapse due to false lumen radial force of an acute type b aortic dissection. Ann Vasc Surg 2014;28(8):1931.e9–1931.e12. [8] Sketch MH Jr, Ball M, Rutherford B, Popma JJ, Russell C, Kereiakes D. Evaluation of the Medtronic (Driver) cobalt-chromium alloy coronary stent system. Am J Cardiol 2005;95(1):8–12. [9] Kleinstreuer C, Li Z, Basciano CA, Seelecke S, Farber MA. Computational mechanics of Nitinol stent grafts. J Biomech 2008;41(11):2370–8. [10] Yang CPO, Hsu CP, Chen WY, Chen IM, Weng CF, Chen CK, et al. Aortic remodeling after endovascular repair with stainless steel-based stent graft in acute and chronic type b aortic dissection. J Vasc Surg 2012;55(6):1600–10. [11] Demanget N, Duprey A, Badel P, Orgéas L, Avril S, Geindreau C, et al. Finite element analysis of the mechanical performances of 8 marketed aortic stent– grafts. J Endovasc Therap 2013;20(4):523–35. [12] Catanese J III, Cooke D, Maas C, Pruitt L. Mechanical properties of medical grade expanded polytetrafluoroethylene: the effects of internodal distance, density, and displacement rate. J Biomed Mater Res 1999;48(2):187–92. [13] Wang L. Design and process investigation on a new type of thoracic stent. Dalian University of Technology; 2013. [14] Zhang C, Wang G, Mu S, Zheng H, Li Z. Observation of efficacy and complications of personalized weaving machine type covered stent for advanced esophageal carcinoma in elderly patients. J China Clinic Med Imaging 2014;25(7):507–10. [15] Anagnostopoulos CE, Prabhakar MJ, Kittle CF. Aortic dissections and dissecting aneurysms. Am J Cardiol 1972;30(3):263–73. [16] Lee VK, Kim DY, Ngo H, Lee Y, Seo L, Yoo SS, et al. Creating perfused functional vascular channels using 3D bio-printing technology. Biomaterials 2014;35(28):8092–102. [17] Boland T, Xu T, Damon B, Cui X. Application of inkjet printing to tissue engineering. Biotechnol J 2006;1(9):910. [18] Valverde I, Gomez G, Coserria JF, Suarez-Mejias C, Uribe S, Sotelo J, et al. 3D printed models for planning endovascular stenting in transverse aortic arch hypoplasia. Catheter Cardiovasc Interv 2015;85(6):1006. [19] Zhang L. Fabrication and fundamental research of multi-branched blood vessel scaffolds with multi-layered wall. Tsinghua University; 2006. [20] Krohg-Sørensen K, Brekke M, Drolsum A, Kvernebo K. Periprosthetic leak and rupture after endovascular repair of abdominal aortic aneurysm: the significance of device design for long-term results. J Vasc Surg 1999;29(6):1152–8. [21] Davis FJ, Mitchell GR. Polyurethane based materials with applications in medical devices. In: Bártolo P, Bidanda B, editors. Bio-materials and prototyping applications in medicine. New York: Springer Science+Business Media; 2008. p. 27–48. [22] Stoeckel D, Pelton A, Duerig T. Self-expanding nitinol stents: material and design considerations. Eur Radiol 2004;14(2):292–301. [23] Azaouzi M, Lebaal N, Makradi A, Belouettar S. Optimization based simulation of self-expanding Nitinol stent. Mater Des 2013:917–28. [24] Holloway JL, Spiller K, Lowman AM, Palmese G. Analysis of the in vitro swelling behavior of poly(vinyl alcohol) hydrogels in osmotic pressure solution for soft tissue replacement. Acta Biomater 2011;7(6):2477. [25] Landau L, Levich B. Dragging of a liquid by a moving plate. Dyn Curved Fronts 1942;17(2):141–53. [26] Scriven LE. Physics and applications of Dip coating and spin coating. Proc MRS 1988;121:717. [27] Vojteˇ ch D, Michalcová A. Influence of heat-treatment on mechanical properties and transformation temperatures of Nitinol. Key Eng Mater 2011;465:471–4. [28] Cui Y, Wan J, Man J, Zhang J, Rong Y. Mechanism of reorientation and redistribution of multi-variants in shape memory alloys under external field. Mater China 2012;31(3):39–43.
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A new process for customized patient-specific aortic stent graft using 3D printing technique Yang Lei a,b,#, Xin Chen a,c,#, Zhen Li a, Lei Zhang a,∗, Wei Sun a, Lei Li d, Feng Tang d a
Department of Mechanical Engineering, Tsinghua University, Beijing 100084, China Beijing Institute of Aeronautical Materials, Beijing 100095, China c China Electronics Technology Group Corporation, Beijing 100846, China d The First Hospital of Tsinghua University, Beijing 100016, China b
a r t i c l e
i n f o
Article history: Received 5 January 2019 Revised 14 November 2019 Accepted 10 December 2019 Available online xxx Keywords: 3D printing Aortic dissection Personalized aortic stent graft Rapid prototyping sacrificial core-coating forming technique
a b s t r a c t Endovascular aneurysm repair (EVAR) is a popular and effective treatment for descending aortic disease. However, the majority of existing coating stents used in EVAR are of a standard design which may not meet the size or structural requirements of different patients. Therefore, in this paper, we propose using 3D printing and controlled deposition as a patient-specific aortic stent graft manufacturing technique. The methodology involves the use of a rapid prototyping study sacrificial core-coating forming (RPSC-CF) technique to develop an aortic stent graft that consists of a film and metallic stent. Polyether polyurethane and nickel-titanium alloys were chosen due to their shape memory properties and good biocompatibility. The resulting customized stent grafts meet the demands of personalized therapy and invasive surgery, and perform well as demonstrated from burst pressure testing and the degree of radial support provided and radial support force tests, laying the foundation for precise aortic dissection treatment. © 2019 IPEM. Published by Elsevier Ltd. All rights reserved.
1. Introduction Endovascular aneurysm repair (EVAR) is the primary method of treating aortic dissection [1]. In the EVAR process, a stent graft is implanted into the lesion to close the rupture of the aortic intima in order to reconstruct the aortic wall. The blood is blocked from the false lumen in case the false lumen further expands or ruptures [2]. Existing methods for forming aortic stent grafts rely on standardized manufacturing processes, producing stents with a fixed diameter and length [3]. However, along the axial direction when moving from proximal points to more distal areas, the diameter of the human thoracic descending aorta decreases. Therefore, standard diameter stents cannot meet personalized size and shape requirements [4].
Abbreviations: EVAR, endovascular aneurysm repair; RPSC-CF, “rapid prototyping” and “sacrificial core” and “coating” forming; PET, polyethylene terephthalate; e-PTFE, expanded polyethylene terephthalate; Ni-Ti alloy, nitinol; CT, computed tomography; PU, polyurethane; FDA, The Food and Drug Administration; Af, austenitetransition temperature; PVA, polyvinyl alcohol; HG-stained, histologicalgradingstained. ∗ Corresponding author. E-mail address: [email protected] (L. Zhang). # These authors contributed equally to this work.
Stent placement that does not correspond to the specific size and structure of the patient’s aorta may affect the adherence of the stent [5–7]. A stent graft diameter that is less than the patient’s aortic diameter will lead to poor stent adhesion, which can result in blood leakage or stent displacement. A stent diameter greater than the patient’s aortic diameter will cause the vessel to break, resulting in treatment failure. Additionally, the actual structure of the human thoracic aorta is not straight but curves; consequently, some aortic stent scaffolds cannot fit to the curved structure of blood vessels, which causes deflation and treatment failure. The branched structure of the aortic arch presents great challenges for surgical treatment. In current clinical practice, the open window method is usually adopted for the treatment of branch vessels, and the stent graft is processed and connected during surgery. This kind of surgery requires the joint implantation of multiple stent grafts, which is extremely risky. Therefore, there is a clinical demand for bending and branch customization of aortic stent grafts. At present, the most common stent graft on the market consists of a polymer film material like polyethylene terephthalate (PET) or expanded polyethylene terephthalate (e-PTFE) and a nitinol alloy [11,12]. Biopolymer materials are the primary materials for the film. Polymer film forming processes are mainly divided into textile forming and non-textile forming techniques. Textile forming
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uses a loom to weave polyester fibre into a tubular film. Because of its compact structure, stable size, and controllable thickness, this method is widely used by stent graft manufacturers. Non-textile forming methods mainly include extrusion forming, electrostatic spinning, and micro-nano-composition. The metal components of stents can be made from stainless steel, cobalt-chromium alloy, or nickel-titanium alloy, according to the surgical release methods to be used [8–10]. Metal stents are structured into ‘W’ or ‘Z’ shapes formed using laser cutting or weaving. The metal stent and film are then bound by methods including stitching and weaving. To a degree, weaving and stitching methods can meet the requirements of personalized manufacturing. Using a model of an actual aorta, Wang et al. [13] formed a polyester film using a suturing method to ensure similarity to the true structure of the aorta; then, a stereotyped nitinol (Ni-Ti alloy) stent was adhered to the film to prepare the personalized aortic stent. Zhang et al. [14] utilized a machine weaving method to weave film onto a nitinol stent in order to achieve customization. Cook Medical LLC can customize stent grafts, but its production cycle requires approximately one month. Patients with acute aortic dissection need to be treated promptly; after diagnosis, the mortality rate of acute aortic dissection is as high as 36–72% within 48 h and increases to 62–91% within one week [15]. Therefore, the one-month production period cannot guarantee timely treatment for patients with acute aortic dissection. For these reasons, the idea of using 3D printing to manufacture personalized stent grafts has been advanced [16,17] in order to produce a personalized stent graft that matches the true shape of a patient’s aorta in a very short time frame. 3D printing technology can transform a digital model directly into a solid sample, allowing for inexpensive and highly efficient customization. This technique not only ensures better treatment outcomes due to the stents’ longitudinal flexibility and compatibility with the vascular wall, but also avoids the disadvantage of the long production cycle of current stent grafts, indicating that this technique has good prospects for clinical application. At present, no research has applied 3D printing technology to the manufacture of stent grafts. 3D printing has been primarily used as an assistive technique for interventional treatments. Valverde et al. [18] used 3D printing to build a model of a patient’s abdominal aortic aneurysm in order to select the most suitable stent-graft system and simulated the graft implantation in vitro to improve the accuracy of the interventional therapy. Because 3D printing technology allows for rapid personalization, we wanted to apply 3D printing technology to the manufacture of stent moulding. However, it is difficult to use 3D printing technology to directly integrate and form the composite materials that comprise the covering film and metal stent. To address this problem, a sacrificial core-coating forming technique first proposed by Zhang [19] was used: the Rapid Prototyping Sacrificial Corecoating Forming (RPSC-CF) technique involves the formation ofTo solve this problem, this project adopts the rapid prototyping sacrificial core-coating forming (RPSC-CF) technique based on 3D printing technology proposed by Zhang . RPSC-CF employs the formation of a water-soluble non-toxic solvent core using rapid prototyping to ensure the stent lumen structure. The multilayer tube wall structure can then be formed layer-by-layer using coating, spraying, or other material deposition processes, and finally, the inner core is dissolved to obtain the stent. Using RPSC-CF technology, 3D printing is used in this study to print customized aortic stent cores with soluble materials, then the core is coated with polymer multiple times to form a multi-layer film. Next, a conventional metal stent is wrapped between the two layers of film to integrate the alloy stent and the film. This study investigates materials for the film, metal stent, and water-soluble sacrificial core, as well as the stent formation process, ultimately
developing a technology for forming and manufacturing personalized stent grafts. 2. Methods and materials 2.1. Stent graft fabrication Preparation of a personalized stent grafts should include two primary considerations: ensuring that the stent accurately matches the true structure of the patient’s aorta and, because the stent is composed of film and a metal stent, determining how to combine these two materials stably and effectively. These two aspects are fundamental to the success of the stent graft fabrication. Aortic 3D model restoration was performed with Mimics 10.01 software. Based on computed tomography (CT) data, a 3D aortic model file was obtained after eliminating noise from other tissues and smoothing the surface. The stent core was designed using the SolidWorks 2016 platform. Vascular bending angles can vary due to individual differences, and the diameter of the descending thoracic artery is irregular along the axial direction. To accurately extract size information from the descending thoracic aorta, diameters were taken from the 3D aortic model every 5 mm along the axial direction. Additionally, to ensure effective fixation of the stent graft into the lesion, clinical experience has shown that the core diameter of the stent should be 10% greater, and the top and bottom ends 15% greater, than the diameter of the blood vessel. Existing methods to integrate the metal and film moulding include suturing and deposition. The suture method for stent graft integration uses special stitch and suture techniques to combine the metal stent and film. In this method, it is easy to penetrate or even rupture the film. Moreover, this method’s low efficiency (due to the need for manual sewing) is not compatible with the requirement for rapid preparation [1,20]. In the deposition method, the outer film is coated onto the surface of the inner film of the stent graft. After the solution has volatilized, the metal stent is tightly wrapped by the inner and outer films. This method avoids film infiltration or rupture caused by hand sewing and has higher efficiency than the suture method. Therefore, in this study, the metal stent is combined with the film using the deposition method. The fabrication process depends on RPSC–CF technology (Fig. 1). 2.2. Film material Polyurethane (PU) has been chosen as the preferred material in this study. PU is polymerized by polyol (polyether and polyester), isocyanate, and a chain extender (diol and amine) that contains carbamate in its backbone [21]. Polyether PU produced by Zhejiang Huafeng Chemistry Co., Ltd, has been used as a film in aortic stent grafts. The US Food and Drug Administration (FDA) has certified that this material has good biocompatibility and resistance to hydrolysis. After it was dissolved in tetrahydrofuran and deposited into a film, its biocompatibility was detected in vitro: it exhibited a haemolysis rate of 0.46% and a relative cell productivity rate, obtained from cytotoxicity tests, of 93.5%. Platelet adhesion test was implemented. Human whole blood with anticoagulant was centrifuged for 5 min at 10 0 0r/min and platelet-rich supernatant was obtained. The PU samples were placed in the platelet-rich serum in a cell incubator at 37 °C and 5% CO2 for 1 h. The samples were immersed in glutaraldehyde aqueous solution with a volume ratio of 2.5% v/v for 12 h after rinsing with saline several times and then dried in a vacuum after dehydration of alcohol. After gold spraying treatment, platelet adhesion on the surface of the sample was observed by environmental scanning electron microscopy. No obvious agglomeration
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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Fig. 1. Technical Route.
Table 1 Components of Nitinol. Element wt% Nb <0.005
Ni 55.65 Fe <0.005
C 0.0043 O 0.037
Co <0.005 N 0.0044
Cu <0.005 H <0.0010
Cr <0.005 Ti remainder
core can be rapidly moulded using fused deposition manufacturing techniques. In this paper, PVA produced by Shenzhen Esun Industrial Co. with a 1.75 mm diameter and a printing temperature of 180–210 °C was used for 3D printing. A Raise3D N2 3D printer was used to form the soluble core. The following print parameters were used: a layer thickness of 0.1 mm, a surface thickness of three layers, a high precision surface, and a filling rate of 3%. After printing, the size tolerance of the core diameter was 0.1 mm and the core met the design requirements 3. Results and discussion 3.1. Stent graft design Fig. 2. Platelet adhesion test. No obvious agglomeration or pseudopodium were observed by environmental scanning electron microscopy.
or pseudopodium were observed in the test (Fig. 2). Therefore, the PU film can be used as a biomedical material. 2.3. Metal stent material Nitinol is typically selected as the material for metal stents due to its super-elasticity, shape memory, and biocompatibility [22,23]. The super-elasticity of nitinol provides good compliance with the vessel wall. In the field of interventional treatment, more than 4/5 of products rely on nitinol’s super-elasticity. In this paper, the nitinol stent is shaped by a heat-treatment process. The nitinol wire used in the experiment was provided by the Jiangyin RongBang New Material Technology Co. The wire diameter was 0.30 mm, it had a bright surface, its tensile strength was 1453 MPa, its fracture elongation rate was 14%, and its austenite-transition temperature (Af ) was 15 °C. The alloying components are shown in Table 1. 2.4. Sacrificial core material In this study, we use polyvinyl alcohol (PVA) as the watersoluble sacrificial material. PVA has good biocompatibility. Moreover, PVA is soluble in water and insoluble in most organic solvents such as toluene and tetrahydrofuran [24]. The entire soluble
For the purpose of clinical animal experiments, CT data of Chinese mini-pigs from a Philips Brilliance 64 were provided by the First Hospital of Tsinghua University. Fig. 3 shows the results of patient-specific aortic stent graft design. 3.2. Film moulding The stent uses a vertical dip coating method. According to the requirements of the process, we built a uniform vertical pulling platform (Fig. 4). The vertical motion axis is driven by a stepper motor with a motor step angle of 1.8° (corresponding to a subdivision of 32), and single machine single-turn motion z-axis vertical stroke of 8 mm. The PU film was fabricated using a dip coating technique. When the viscosity of the fluid and the pulling speed are quick enough, the film deposit thickness for flat substrates is related to the viscous drag (ηU0) and gravity (ρ g) [25,26]:
h0 = c1 (ηU0 /ρ g )1/2 ,
(1)
where ρ is the liquid density, g is the gravitational acceleration, η is the fluid viscosity, U0 is the pulling speed and c1 is 0.8 for Newtonian fluids. When the pulling speed is between 1 mm/s and 10 mm/s and the fluid viscosity η is low, surface tension γ LV affects the thickness of the film; formula (1) can be written as follows:
h0 = 0.94
ηU0 2/3 . γLV 1/6 (ρ g)1/2
(2)
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Fig. 3. Stent Graft Design (A) Structure reduction of the descending aorta; (B) For the experimental design of the straight stent graft, in order to ensure the bracket’s good geometric compliance, an inwardly curved ‘pleat’ is designed every 10 mm along the bracket’s axial direction with an axial length of 5 mm, so that the stent can be bent in any direction. This design prevents distortion when the stent graft is bent; (C) Bent aortic stent graft sacrifice core design.
decreases the uniformity of the dip coating. Therefore, we investigated the effects of film thickness to find the optimal coating process. The PVA sacrificial cores used in the experiments had 10 mm diameter cylinders. A pulling speed of 5 mm/s, a solution concentration of 10%, and a two-dip coating process were found to be the optimum dip coating conditions for the inner layer of the film (Fig. 5). Due to the high concentration of the solution, the outer layer of the film was not flat. To rectify this situation, a pulling speed of 5 mm/s, a solution concentration of 5%, and two-dip coating process found to be optimal for the formation of the outer film layer. 3.3. Nitinol stent moulding In this study, the nitinol stent was shaped by heat treatment. The stainless steel mould was produced with a high-speed computer numerical control milling process; the current mould production time is approximately 2 days. When including the other stent graft fabrication steps, the total production cycle takes 3 days. After shaping, the nitinol stent should have good super-elasticity and exhibit proper radial support. However, the heat-treatment temperature and duration affect the nitinol’s Af and mechanical properties. Therefore, we investigated the appropriate heat-treatment parameters to optimize nitinol stent performance. The temperature at which nitinol is heat treated is generally between 460 °C and 500 °C [27]. Optimal heat treatment was achieved through cooling by quenching.
Fig. 4. Uniform vertical pulling platform for deposition processing.
The PU is dissolved in tetrahydrofuran. After the tetrahydrofuran has been completely volatilized, the PU wraps around the water-soluble core to form the stent graft. To ensure optimal film performance, its thickness should be reduced as much as possible. Furthermore, considering that the stent should be manufactured quickly, a low dip coating time, a high solution concentration, and a relatively high pulling speed should be adopted. However, an excessive concentration increases the solution’s viscosity, which is not conducive to its volatilization. In addition, a high speed pulling
3.3.1. Effect of heat treatment on af At temperatures greater than nitinol’s Af , the nitinol is in a stable austenite state. Under applied stress, a martensitic transformation occurs, which results in an unstable state. The unstable martensitic phase will undergo an inverse phase transformation to a stable austenite phase after stress unloading. Therefore, nitinol is super-elastic on the macro scale [28]. For the aortic stent graft, the Af should be lower than 37 °C to ensure that nitinol exhibits super-elasticity in the human body. In addition, the radial support force of the nitinol stent increases with a decreasing Af . A low Af leads to strong external forces, which may cause vascular damage and hyperplasia. Results from differential scanning calorimetry (Fig. 6A) showed that heat treatment at 480 °C resulted in an Af less than 37 °C and better stability.
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Fig. 5. Effects of film thickness. (A) When testing solution concentrations (w/v) of 7.5%, 10%, 12.5%, and 15% at a pulling speed of 8 mm/s for 5 dip coats, we found that the film thickness was exponentially related to the solution concentration. (B) When testing pulling speeds of 2 mm/s, 5 mm/s, 8 mm/s, and 10 mm/s at a solution concentration of 12.5% (w/v) for 5 dip coats, we found the film thickness exhibited a logarithmic relationship with respect to the pulling speed. (C) For a pulling speed of 8 mm/s and a solution concentration of 12.5% (w/v), the relationship between film thickness and number of dip coatings was exponential.
Fig. 6. Effects of heat treatment parameters. (A) Effect of different heat treatment conditions on Af . The initial Af was 15 °C. At 460 °C, the Af decreased slightly, then rose rapidly with time. At 500 °C, the Af declined over time. At 480 °C, the Af was approximately 10 °C and was relatively stable; (B, C, D) With an increased heat-treatment time, the stress on the upper and lower platforms tended to decrease, but the residual strain rose. When the time exceeded 10 min, the residual strain was greater than 0.5%, indicating that irreversible plastic deformation had occurred. Therefore, the heat-treatment time should not exceed 10 min, and the heat-treatment temperature should be greater than 460 °C.
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Fig. 7. Shaped nitinol wire after 20 min, 25 min, and 30 min of heat treatment (left to right).
Fig. 8. Personalized aortic stent grafts. (A) Combined nitinol stent and stent film; (B) Personalized aortic stent graft; (C) Aortic stent graft with branch; (D) Straight aortic stent graft.
3.3.2. Effect of heat treatment on the stress–strain curve The material’s stress–strain curve was tested according to the ASTM F2516-14 standard. The initial gauge length was 10 mm, the travel rate of the testing machine transom was 2 mm/s and the test temperature was 26 °C. The results showed that the upper platform stress, lower platform stress, and residual strain of the original nitinol were 397 MPa, 50 MPa and 0.3%, respectively. Due to the mechanical property changes caused by heat treatment, 12 group experiments were prepared for heat-treatment parameter determination (Figs. 6B, C, and D). The chosen heat treatment conditions were 480 °C for 10 min.
3.3.3. Moulding the stent The stainless steel mould has thermal inertia; after weaving the nitinol to the mould, the shaping effect of the nitinol is difficult to maintain without the mould. Therefore, heat treatment was extended to 20 min, 25 min, and 30 min at 480 °C (Fig. 7). After 30 min of heat treatment at 480 °C, nitinol achieved a good shaping effect. Moreover, under this heat treatment condition, the Af of the Ni-Ti stent was 20 °C and its residual strain was 0.4%, indicating that the stent can exhibit super-elasticity in the human body. Therefore, a 30 min heat treatment at 480 °C was the most suitable for the shaping process.
Fig. 9. Mechanical properties test. (A, B) Radial support force measuring device; (C) Explosive pressure detection device.
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Fig. 10. Preliminary in vivo test results. (A) Implantation Day 0; (B) Implantation Day 30. Marker 1 to marker 2 10 mm.
3.4. Combining the metal stent and film The key to preparing an optimal stent graft lies in the firm combination of the film and the metal stent. Here, the deposition method was used for this step. The heat-treated nitinol stent was wrapped in the inner coating surface of a 5% PU solution by dipping, and when the solution evaporated, the metal stent was wrapped by the inner and outer membranes (Fig. 8A). Finally, the core was dissolved in water due to the high water-solubility of the PVA material. The stent graft was obtained when the core was fully dissolved (Figs. 8B, C, and D). 3.5. Mechanical properties 3.5.1. Radial support force In this paper, the stress required by reducing the diameter of the stent by half is used as the radial support force. GORE® TAG® Thoracic Endoprosthesis, Medtronic ValiantTM Thoracic Stent Graft, Cook Zenith Alpha Thoracic Endovascular Graft are used as reference standards and their single-ring radial support forces are measured between 0.28 N and 0.34 N. The experimental results showed that the supporting force of the single nitinol ring was 0.3 N, which is equal to the supporting force of an aortic stent graft that has already been clinically applied (Figs. 9A and B). Therefore, the stent graft can support an appropriate radial support force.
3.5.2. Burst pressure Generally, the normal range of human systolic blood pressure is 12.0–18.7 kPa and the normal range for diastolic blood pressure is 8.0–12.0 kPa. Using deionized water with a similar blood density as the pressure medium, the liquid bursting pressure was measured for the stent graft (Fig. 9C) and indicated that the bursting pressure of the personal stent graft was greater than 70 kPa, which is safe and reliable. This result shows that the stent could be implanted into the descending aorta. 3.6. In vivo implantation As the stent graft in this study was found to meet the basic requirements for in vivo implantation, we then conducted in vivo testing. In vivo implantation was carried out on a Chinese minipig in the Department of Vascular Surgery at the First Hospital of Tsinghua University. Preliminary test results indicate that 30 days after implantation (Fig. 10), the stent graft was compliant and effectively supported the descending aorta with no displacement. However, there is a slight atrophy in the middle of the stent graft. In future work, the radial support force of the nitinol in the stent will be reinforced to support to the aorta wall. Blood tests at 24 h, 48 h, and 72 h after surgery show that the leukocyte counts are 19.08 × 109 /L, 19.99 × 109 /L, and
Fig. 11. Pathological sections. (A) Smooth muscle of upper convexity of aortic arch; (B) Smooth muscle of concavity of aortic arch.
Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002
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Y. Lei, X. Chen and Z. Li et al. / Medical Engineering and Physics xxx (xxxx) xxx
24.21 × 109 /L (reference range: 10.20 × 109 /L to 30.00 × 109 /L). The percentages of neutrophils at these time points were 35.7%, 26.4%, and 36.2%, respectively (reference range 20% to 60%). The blood test results show that the blood index of the experimental Chinese mini-pig is normal and the inflammatory response is small after the intervention. Histologicalgrading-stained (HG-stained) aortic sections (Fig. 11) show that although there is partial oedema in the smooth muscle tissue on both sides of the aortic arch, the degree of damage is similar, indicating that the support on both sides of the aortic arch is roughly equivalent. Accordingly, the stent graft meets the design requirements. In summary, the stent graft performed well in vivo. Further results on the in-vivo performance will be published in due course. 4. Conclusion Customized stent grafts with proper biocompatibility and mechanical properties were successfully fabricated based on rapid prototyping sacrificial core-coating forming (RPSC-CF) technique. Film was prepared from polyether PU. The film’s thickness was 0.185 mm, increasing the aortic stent scaffold burst pressure to greater than 70 kPa, which is a safe and reliable level for a device to be implanted into the human descending aorta. Besides, the stent’s single ring radial support force was 0.3 N with 0.3 mm nitinol wire. The stent graft can support an appropriate radial support force. Meanwhile, using a real descending thoracic aortic structure, curved and branched stents were manufactured, demonstrating the ability of these materials and production methods to meet the needs of individual patients. For the last, the entire stent graft preparation cycle was 3–4 days. With continued modification of the preparation process, the entire stent preparation cycle is expected to be compressed to within 48 h to ensure that patients with aortic dissection receive timely treatment. This new process for customized patient-specific aortic stent graft provides a new idea for precise aortic dissection treatment. Declaration of Competing Interest None. Acknowledgments The Beijing Municipal Health Planning Commission 2016-2-4121 and the Beijing Life Science Frontier Technology Cultivation Study Z14110 0 0 0 0214012 supported this study. Ethical approval The in-vivo test carried out in this study is in accordance with Chinese National Laboratory Animal Standard and EU Directive 2010/63/EU. References [1] Dua A, Kuy SR, Lee CJ, Upchurch GR, Desai SS. Epidemiology of aortic aneurysm repair in the United States from 20 0 0 to 2010. J Vasc Surg 2014;59(6):1512–17.
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Please cite this article as: Y. Lei, X. Chen and Z. Li et al., A new process for customized patient-specific aortic stent graft using 3D printing technique, Medical Engineering and Physics, https://doi.org/10.1016/j.medengphy.2019.12.002