WO2019099080A1 - Flexible polymer anti-migration stent - Google Patents

Abstract

A stent includes a flexible, shape-retaining tubular body defining an internal lumen and having first and second ends and an outer surface. A helical spiral projects from the outer surface and extending around the tubular body from the first end to the second end. A method for supporting the passageway of a client and a method of making a stent are also disclosed.

Description

FLEXIBLE POLYMER ANTI-MIGRATION STENT

CROSS-REFERENCE TO RELATED APPLICATIONS.

[1] This application claims priority to US Provisional Patent Application Serial Number 62/586,609 filed on November 15, 2017, entitled“Novel 3D- Printed Flexible Polymer Esophagus Stent For Esophagus Cancer”, the disclosure of which is incorporated herein by reference in its entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH AND DEVELOPMENT

[2] This invention was made with government support under Contract No. 5R03CA201960-02 awarded by the National Cancer Institute of the National Institutes of Health. The government has certain rights in this invention.

FIELD OF THE INVENTION

[3] This invention relates to stents, and more particularly to stents for reducing migration in the body.

BACKGROUND OF THE INVENTION

[4] Esophageal cancer is the sixth leading cause of cancer-related deaths worldwide and recognized as one of the most difficult and challenging malignancies to cure. The five-year survival rate of these patients remain very low, less than 20%. For those patients with inoperable esophageal

malignancies, using esophageal stents is the palliative therapy for regular drinking and feeding.

[5] Currently, clinically used stents include self-expanding plastic stents (SEPS) and self-expanding metal stents (SEMS). SEMS have been widely used in the past decades for malignancies. These metal stents utilize the advantages of the metal mesh structure to self-expand after they are deployed in the esophagus, thus opening the blocked esophagus caused by tumor tissue. The mesh structure makes the stent easily compressed into a small dimension so that it can be loaded into a stent guide catheter before being deployed into the esophagus. Although the clinical use of these self expanding stents has grown immensely over the past decade due to the ease of placement, complications associated with these self-expanding esophageal stents have occurred at an early or late stage of the procedure. These complications include chest pain, fever, tumor tissue ingrowth, bleeding, recurrence of strictures, reflux disease, and stent migration into the stomach. The ends of metallic stents occasionally lead to complications such as fistulas or perforations. Although large-diameter stents and flared flange ends may reduce the risk of stent migration, increasing stent diameter and the size of stent ends increases the risk of stent-related esophageal complications.

Currently, there is not an ideal stent available in clinics to overcome the complications associated with SEMS stents. [6] SEPS have been used for the treatment of benign esophageal conditions, as they are thought to have several advantages over standard SEMS, for example, low cost, ease of placement and retrieval, limited local tissue reaction, effectiveness for treating benign esophageal conditions.

However, researchers have reported that SEPS have a high overall rate of stent migration, high complication rate (hemorrhage, tumor overgrowth), and limited use in alleviating dysphagia caused by benign strictures.

[7] The mesh structure of SEPS and SEMS allows the SEPS or SEMS to be easily compressed into a catheter to deploy and self-expand, but this mesh structure also causes the complications because the mesh structure cannot prevent new tissue from growing through the mesh. Although partially-covered or fully-covered stents may prevent the tissue ingrowth, the covered stents increase the migration rate and other complications. Tumor tissue ingrowth into the stent and stent migration remain unsolved problems.

SUMMARY OF THE INVENTION

[8] A stent includes a flexible, shape-retaining tubular body defining an internal lumen and having first and second ends and an outer surface. A helical spiral projects from the outer surface and extends around the tubular body from the first end to the second end.

[9] The stent can be made from a material having a hardness of from Shore A 55 and Shore A 90. The stent can include a combination of a hard material and a soft material. The hard material has a Shore A hardness greater than the Shore A hardness of the soft material. The combination can have a Shore A hardness between Shore A 55 and Shore A 90.

[10] The stent can further include end flanges at the first and second ends. The end flanges can extend axially and laterally outwardly from the first and second ends. The tubular body of the stent can be nonporous.

[11] The stent can comprise a stent mixture of a thermoplastic soft material and thermoplastic hard material. The hard material can be at least one selected from the group consisting of polylactic acid (PLA, PDLA, PLLA isomers), polycaprolactone, polyvinyl acetate, and polyethylene glycol. The soft material can be at least one selected from the group consisting of thermoplastic polyurethane (TPU) and thermoplastic copolyester. The stent mixture can comprise 3-20% hard material and 97-80% soft material, based on the total weight of the mixture. The hard material can be biodegradable and biodegrades at a faster rate than the soft material. [12] The stent mixture can further include a drug capable of release from the stent. The stent can further comprise a drug-loaded layer on the surface of the stent.

[13] The spiral pitch of the helix can be from 2 to 7 mm. The thickness of the helix can be from 0.3 mm -1.5 mm. The depth of the helix can be from 0.3 mm -1.5 mm. The thickness of the stent can be from 200 miti-800 pm. The length of the stent is from 10 mm-100 mm. The diameter of the internal lumen can be from 4 mm-26 mm. Other values and ranges are possible.

[14] The stent can provide a frictional force in a pig esophagus of from 0.4 to 4 N. The stent can provide a self-expansion force of from 2 N to 15 N. The stent can provide a compression force of from 1 N to 30 N. The stent can provide an anti-migration force of from 1 N to 10 N. The stent migration distance can be from 0 mm to 5 mm. Other values and ranges are possible.

[15] A method for supporting a passageway of a patient includes the step of providing a stent, where the stent comprises a flexible, shape-retaining tubular body defining an internal lumen and has first and second ends and an outer surface. A helical spiral projects from the outer surface and extends around the tubular body from the first end to the second end. The stent is positioned in the passageway. The stent can be customized for the patient.

[16] A method of making a stent can include the step of heating and mixing a hard thermoplastic material and a flexible thermoplastic material to create a stent precursor composition. The stent precursor composition is extruded from a thread maker to form a stent precursor thread. The stent precursor thread is cooled. The stent precursor thread can be fed into a 3D printer. The 3D printer heats and melts the stent precursor thread. A stent is printed with the 3D printer and the melted stent precursor thread.

BRIEF DESCRIPTION OF THE DRAWINGS

[17] There are shown in the drawings embodiments that are presently preferred it being understood that the invention is not limited to the

arrangements and instrumentalities shown, wherein:

[18] Figure 1 is a perspective view of a stent according to the invention.

[19] Figure 2 is a side elevation, partially in phantom.

[20] Figure 3 is a cross section of the stent as inserted into an esophagus.

[21] Figure 4 is a side elevation and schematic depiction of spiral geometry.

[22] Figure 5 is a plot of Maximum von Mises Stress (^*10⁶ Pa) for acute, right and obtuse spiral geometry.

[23] Figure 6 is a side elevation of a stent having a first spiral number.

[24] Figure 7 is a side elevation of a stent having a second spiral number.

[25] Figure 8 is a plot of Frictional Force (N) vs. Revolutions (Number) of the spiral.

[26] Figure 9 is a schematic diagram of a process for making a stent according to the invention.

[27] Figure 10 is a plot of Compression Force (N) vs. strain for different stent compositions and a SEMS.

[28] Figure 11 is a plot of Self-expansion Force (N) for different stent compositions and a SEMS, and with and without spirals.

[29] Figure 12 is a schematic diagram illustrating an anti-migration force testing apparatus and procedure.

[30] Figure 13 is a plot of Anti-migration Force for different stent

compositions and a SEMS, and with and without spirals. [31] Figure 14 is a schematic diagram illustrating a migration distance testing apparatus and procedure.

[32] Figure 15 is a plot of Migration Distance (mm) for different stent compositions and a SEMS, and with and without spirals.

[33] Figure 16 is a plot of Diameter of esophagus in the middle (mm) for different stent compositions and a SEMS, and with and without spirals.

[34] Figure 17 is a plot of Diameter of esophagus at the ends (mm) for different stent compositions and a SEMS, and with and without spirals.

[35] Figure 18 is a plot of remaining mass (g) vs. time (week) for degradation in neutral phosphate buffered saline (PBS) solution and degradation in acidic simulated gastric fluid (SGF) solution.

[36] Figure 19 is a plot of Weight loss (%) vs. Time (week) for degradation in neutral phosphate buffered saline (PBS) solution and degradation in acidic simulated gastric fluid (SGF) solution.

[37] Figure 20 is a plot of transmittance (%transmittance) vs. Wavenumbers (cm ¹) for several different periods (weeks) in PBS solution.

[38] Figure 21 is a plot of transmittance (% transmittance) vs. Wavenumbers (cm-1 ) for several different periods (weeks) in SGF solution.

[39] Figure 22 is a plot of Compression force (N) vs. Strain (%) before degradation and after degradation of 1 , 2, and 3 months.

[40] Figures 23 A - 23 D are depictions of live/dead staining of human primary esophagus epithelial cells for discs composed of TPU (Figure 23 A); 5PLA95TPU (Figure 23 B); 10 PLA90TPU (Figure 23 C); and 15PLA85TPU (Figure 23 D).

[41] Figure 24 is a plot of Intensity of Absorbance vs. Time (days) for different stent compositions.

[42] Figure 25 is a plot of Stress (^*10⁷ Pa) vs. stretch for different stent compositions, for experimental data (ED) and experimental stress (ES).

[43] Figure 26 is a plot of Stress (^*10⁵ Pa) vs stretch for Mucosa-Submucosa ED, Mucosa-Submucosa curve fit, Muscle ED, and muscle curve fit.

DETAILED DESCRIPTION OF THE INVENTION

[44] As shown in Figs. 1 -2, a stent 30 comprising a flexible, shape-retaining tubular body 34 defining an internal lumen 56 and having first end 38 and second end 42 and an outer surface is shown. A helical spiral 46 projects from the outer surface of the tubular body 34 and extends around the tubular body 34 from the first end 38 to the second end 42.

[45] The stent can further include first end flange 50 at the first end 38 and second end flange 54 at the second end 42. The end flanges 50 and 54 extend axially and laterally outwardly from the first and second ends 38 and 42, respectively, and act to secure the stent 30 in position when placed in the body.

[46] The stent of the invention can be placed into many different body canals or structures, such as the esophagus. Fig. 3 is a cross section of the stent 30 as inserted in an esophagus 60. This is commonly performed as a palliative treatment for esophageal cancer.

[47] The stent 30 can be made from a number of different materials. The stent can be formed from a stent composition comprising a mixture of a soft material and hard material. The stent composition can be thermoplastic to facilitate 3D printing. The hardness range of the composite stent composition can be Shore A 55-90, as tested by a Shore durometer based on ASTM D2240. The hardness of the stent composition can be Shore A 55, 56, 57, 58, 59, 60, 61 , 62, 63, 64, 65, 66, 67, 68, 69, 70, 71 , 72, 73, 74, 75, 76, 77, 78, 79, 80, 81 , 82, 83, 84, 85, 86, 87, 88, 89 and 90, or within a range of any high value and low value selected from these. The soft material can have a Shore A value less than The Shore A value of the hard material, so long as the mixture has a combined Shore A between 55-90.

[48] The hard material can be any of several suitable materials. The hard material can be at least one selected from the group consisting of polylactic acid (PLA, PDLA, PLLA isomers), polycaprolactone, polyvinyl acetate, and polyethylene glycol. The hard material can be thermoplastic. Other hard materials are possible.

[49] The soft material can be any of several suitable materials. The soft material can be at least one selected from the group consisting of

thermoplastic polyurethane (TPU) and thermoplastic copolyester. The soft material can be thermoplastic. Other soft materials are possible.

[50] The soft material provides the needed flexibility for ease of deployment. The hard material provides enough mechanical expansion for opening the blocked esophagus. Flexible TPU and crystalline PLA are both widely used in biomedical applications. The combination of a hard material and a soft material using adjustable amounts of soft/rigid composite polymers to replace metallic materials decreases the bleeding risk.

[51] The biodegradable elastomeric and stiff composite polymer materials used in the stent make it sufficiently rigid yet flexible enough to expand and flexibly contact with the esophagus. The radial force of the stent can be modulated by the adjustment of different ratios of the soft/rigid polymers. [52] Elastomeric TPU as the soft polymer has a unique combination of toughness, durability, biocompatibility, and biostability. The hard polymer can be stiff poly(lactic acid) (PLA) that provides enough mechanical strength for opening the blocked esophagus at the time of placement, and then provides slow degradation to adapt to the mechanical stress of the esophagus.

Elastomeric TPU and crystalline PLA are both widely used in biomedical applications. Research has shown that molten TPU and PLA are miscible. Through modulating the different ratios of PLA to TPU, the flexibility, mechanical strength, and biodegradation rate of the PLA/TPU stent

composition can be manipulated. Combinations of more than two polymers, or a single polymer, can also be used for the stent composition.

[53] The hard material and soft material can be thermoplastic and combined for example by heating and mixing. The heated mixture is then cooled and solidified. The hard material and soft material can be miscible to facilitate homogeneous mixing.

[54] The relative proportions of the hard material and the soft material can vary. In one aspect, the stent composition comprises 3-20 wt. % hard material and 80-97 wt. % soft material, based on the total weight of the stent

composition. The stent composition can comprise 3, 4, 5, 6, 7, 8, 9, 10, 1 1 ,

12, 13, 14, 15, 16, 17, 18, 19 or 20 wt. %, or within a range of any high value and low value selected from these. The soft material can be 80, 81 , 82, 83,

84, 85, 86, 87, 88, 89, or 90 wt. % of the stent composition, or within a range of any high value and low value selected from these. [55] The hard material can be biodegradable and thereby biodegrades at a faster rate than does the soft material. This property can facilitate the stent becoming more flexible over time, and thereby more able to adjust to changes in the characteristics of the body part into which the stent has been placed.

[56] The stent can be constructed to release a drug or other medicament. The stent composition can further include a drug capable of release from the stent. The stent provides a versatile platform for loading different drugs in polymers with adjustable biodegradation rates to control drug release. The biodegradation of at least one of the components of the stent composition can act to release the drug. The outside surface of the stent can be further coated with a drug-loaded polymer layer, thus the stent can provide the unidirectional release of drugs to the mucosa tissue of tumorous esophagus. The coating/3D printing process will not compromise the drug bioactivity, and the coating process is reproducible. The stent can serve as a platform for loading many different drugs using a drug-loaded polymer solution to coat the 3D-printed stent, thus enabling the polymer coating to have potentials to load different drugs and control the drug release rate.

[57] The dimensions and characteristics of the stent can vary. The spiral pitch of the helix can vary. The spiral pitch of the helix can be from 3 to 12 per 20 mm, or 2 mm to 7 mm. The spiral pitch can be 2, 2.25, 2.5, 2.75, 3, 3.25, 3.5, 3.75, 4, 4.25, 4.5, 4.75, 5, 5.25, 5.5, 5.75, 6, 6.25, 6.5, 6.75 or 7 mm, or within a range of any high value and low value selected from these. The thickness of the helix can vary. The thickness of the helix can be from 0.3 mm -1.5 mm. The thickness of the helix can be 0.3, 0.4, 0.5, 0.6, 0.7, 0.8, 0.9, 1.0,

1.1 , 1.2, 1.3, 1.4, or 1.5 mm, or within a range of any high value and low value selected from these. The depth of the helix can vary. The depth of the helix can be from 0.3 mm -1.5 mm. The depth of the helix can be 0.3, 0.4, 0.5, 0.6, 0.7, 0.8, 0.9, 1.0, 1.1 , 1.2, 1.3, 1.4, or 1.5 mm, or within a range of any high value and low value selected from these. The diameter of the stent can vary. The thickness of the stent can be from 200 miti-800 pm. The thickness of the stent including the helix can be 200, 250, 300, 350, 400, 450, 500, 550, 600, 650, 700, 750, or 800 pm, or within a range of any high value and low value selected from these. The length of the stent can vary. The length of the stent can be from 10 mm-100 mm. The length of the stent can be 10, 15, 20, 25,

30, 35, 40, 45, 50, 55, 60, 65, 70, 75, 80, 85, 90, 95 or 100 mm, or within a range of any high value and low value selected from these. The diameter of the internal lumen of the stent can vary. The diameter of the internal lumen can be from 4 mm-26 mm. The diameter of the internal lumen of the stent can be 4, 5, 6, 7, 8, 9, 10, 1 1 , 12, 13, 14, 15, 16, 17, 18, 19, 20, 21 , 22, 23, 24, 25, or 26 mm, or within a range of any high value and low value selected from these.

[58] The stent can provide a frictional force in a pig esophagus of from 0.4 to 4 N. The stent provides a frictional force in a pig esophagus that can be 0.4, 0.6, 0.8, 1.0, 1.2, 1.4, 1.6, 1.8, 2.0, 2.2, 2.4, 2.6, 2.8, 3, 3.2, 3.4, 3.6, 3.8, or 4 N, or within a range of any high value and low value selected from these. [59] The stent can provide a self-expansion force of from 2 N to 15 N. The stent can provide a self-expansion force of 2, 3, 4, 5, 6, 7, 8, 9, 10, 11 , 12, 13, 14 or 15 N, or within a range of any high value and low value selected from these.

[60] The stent can provide a compression force of from 1 N to 30 N. The stent can provide a compression force of 1 , 2, 3, 4, 5, 6, 7, 8, 9, 10, 11 , 12, 13, 14, 15, 16, 17, 18, 19, 20, 21 , 22, 23, 24, 25, 26, 27, 28, 29, or 30 N, or within a range of any high value and low value selected from these.

[61] The stent can provide an anti-migration force of from 1 N to 10 N. The stent can provide an anti-migration force of 1 , 1.5, 2, 2.5, 3, 3.5, 4, 4.5, 5, 5.5, 6, 6.5, 7, 7.5, 8, 8.5 9, 9.5 or 10 N, or within a range of any high value and low value selected from these.

[62] The stent can limit the stent migration distance less of from 0 mm to 5 mm. The stent can provide a stent migration distance of 0, 0.5, 1 , 1.5, 2, 2.5,

3, 3.5, 4, 4.5 or 5 mm, or within a range of any high value and low value selected from these.

[63] Computational simulations were performed to determine the parameters of the spiral helix included the angles, thickness, and interval distance between two helix spirals. In order to determine these parameters, a finite element method was used to computationally analyze the self-expansion and anti-migration forces, based on a simulated model where the stent was placed in the esophagus (Fig. 3). A CAD (Computer-aided design) program was used to design the shape and parameters of the dumbbell-like PLA/TPU materials, and a customized 3D printer with modified nozzles was used to print the dumbbell-like PLA/TPU specimens with standard size for mechanical tests. Tensile tests on the PLA/TPU materials were performed adopting the guidelines indicated in the American Society for Testing and Materials methods (ASTM-D638). Specimens with standard size were printed on the 3D printer. A universal mechanical tester (TestResources Inc., Shakopee, MN) was used to test all the samples in dry conditions at room temperature.

Samples were tested until failure to obtain tension strength. Elastic moduli, maximum strength, and maximum strain were determined. These parameters were used for computational modeling through a finite element analysis software, Comsol Multiphysics™. The main parameters of the PLA/TPU stent included the angles and thickness of the helix spiral grooves and the interval distance (pitch) between two spirals. The angles of the spiral grooves were set at three types of angles: acute, right, and obtuse. The thickness of spirals was set at T1. The length of the stent was set at H1. The thickness, inner and outer radii, and middle stem length of the stent were set at T2, R1 , R2, and H2, respectively. The internal distance between the two spiral grooves was set at D1. The settings of R1 , H1 , and H2 were based on the diameter of a pig esophagus. Stents with 100TPU, 5PLA/95TPU, 10PLA/90TPU and

15PLA/85TPU (wt. %/wt.%) were considered, along with a SEMS stent (a representative of metal stents from Micro-Tech (Nanjing) Co. (Nanjing,

China)). Finite element analysis of PLA/TPU stents was performed for stent models with obtuse, right and acute angle, and local stress versus acute, right and obtuse angle, for a 15PLA/85TPU (wt. %) stent composition. The following parameters of the stents were determined for purposes of

experimentation. The inner diameter of the stent is 10 mm (outer diameter was 10.4 to 10.48 mm), the thickness is 0.4 mm, the diameter of the two flanges is 16 mm, the distance between two spirals (spiral pitch) is 3.5 mm and there are 6 spirals in the stem part of the stent. Stents with spiral thickness of 0.3 mm, 0.6 mm, 0.9 mm, 1.2 mm and 1.5 mm were prepared. The circumferential stress versus circumferential deformation with

15PLA/85TPU stent was analyzed. Stents with spiral numbers 3, 4, 5, 6, 7, 8, 9, 10 were tested, and the frictional force versus spiral number with 100TPU, 5PLA/95TPU, 10PLA/90TPU and 15PLA/85TPU stent was determined.

[64] After running the computational simulation, the stent parameters were obtained. These parameters were then used to print the stents. The

parameters were used to print several stents, including several stent prototypes with different sizes. The printed stents were pliable. The stents can be compressed and recovered to the original shape. This result showed that the 3D-printed PLA/TPU stent could self-expand. Four types of stents with different PLA weights (0, 5%, 10%, and 15%) were printed, named as

100TPU, 5PLA95TPU, 10PLA90TPU, and 15PLA85TPU. Medical grade soft elastomeric TPU with excellent resilience and recoverability was obtained from Lubrizol LifeScience Polymers Co. (Cleveland, Ohio). PLA biopolymer was obtained from Sigma-Aldrich. Several stent prototypes were printed with different sizes. A kind of SEMS stent (Nitinol mesh with atraumatic ends), as a representative of metal stents, was provided as a control. The Nitinol mesh stents with atraumatic ends were obtained from MICRO-TECH(Nanjing) Co., Ltd (Nanjing, China). The middle diameter of the stent is 20 mm, the diameter of the flared ends is 26 mm, and the length is 100 mm.

[65] Mechanical tests on the printed PLA/TPU stents were performed according to the guidelines indicated in the American Society for Testing and Materials (ASTM-D4762) methods. For compression, stents were placed between two plates and testing was performed at a strain rate of 1 %/s until the sample was compressed to 50% strain. Three samples of each group were measured. The forces were recorded. For the expansion force test, the stent first was guided and inserted into a freshly harvested pig esophagus, and then the esophagus with the inserted catheter was placed between the two compression plates of the instrument. The sample was compressed around 20-30% strain. Afterward, the stent was released from the catheter in the esophagus. The expansion forces that the stent gave were recorded when the stent expanded in the pig esophagus.

[66] In order to observe whether the 3D-printed polymeric stent has a similar self-expansion property to the SEMS used in this study, the printed stent was compressed into a clinically used catheter. Then the stent was pushed out from the catheter. The stent gradually self-expanded. After the stent was fully pushed out, it recovered to the original shape. The size changed from the compressed status of 8.32 mm to its original size of 10.48 mm. In order to further verify if the stent could function the same in a constraint lumen, the stent was inserted into the fresh pig esophagus using the catheter. After the catheter was removed, the stent self-expanded in the pig esophagus. The stent enlarged the size of the esophagus, compared to the esophagus without the inserted stent. It was observed that the esophagus was opened and maintained the lumen of the esophagus, compared to the lumen of the esophagus without an inserted stent. The inserted stent opened the lumen of the esophagus gradually after several minutes’ post-release. This functional property of the new 3D-printed polymer stent is similar to the functional property of the SEMS.

[67] In order to determine the optimal angle of the helix spirals on the stent, in terms of the minimal local stress at the cross point of the angle and the dense tubular body, three angles, including acute (45°), right (90°), and obtuse (135°) were analyzed. Fig. 4 is a side elevation and schematic depiction of spiral geometry. The stent 64 has a helix 66 forming an acute angle 68. The stent 70 has a right angle helix 74 forming a right angle 78 with the tubular body. The stent 80 has a helix 84 forming an obtuse angle 88 with the tubular body. Fig. 5 is a plot of Maximum von Mises Stress (^*10⁶ Pa) for acute, right and obtuse spiral geometry. The acute geometry produces the greatest stress. Results shown in Fig. 5 indicate that the acute angle induced a 1.99-fold higher local stress of maximum von Mises than the obtuse angle (p=0.0002). The right angle also induced a 1.72-fold higher local stress of maximum von Mises than the obtuse angle (p=0.0006). The obtuse angle induced the lowest local stress, and also had the lowest risk of being broken when the stent was subjected to an external force.

[68] The pitch or helix spiral spacing is also a significant parameter of the stent. Fig. 6 is a side elevation of a stent having a first spiral number. Fig. 7 is a side elevation of a stent having a second spiral thickness. In order to determine the spiral numbers in a certain length of stent, the effects of different spiral pitch on the simulated friction force were compared. Fig. 8 is a plot of Frictional Force (N) vs. Revolutions (Number) of the spiral on a 20 mm tubular body. It can be seen that 4-8 revolutions produce a peak in the frictional force, corresponding to a pitch of 20/4 or 5 mm to 20/8 or 2.5 mm. Results indicated that the friction force of the stent increased with the number of the spiral grooves until 6 per 20 mm, after which, the friction force decreased when the spiral continued to increase, regardless of the composite polymers. All four groups of materials showed that the anti-migration force peaked at the number of 6 spirals per 20 mm, which means that the pitch was 20/6 or 3.5 mm (Fig. 8). The spiral number at 6/20 mm could provide the optimal frictional force of about 1.5 N. The other parameters, such as the diameter of the middle part and flange part of the stent, were determined at the same time.

[69] Figure 9 is a schematic diagram of a process for making a stent according to the invention. A method of making a stent, comprising the steps of: heating and mixing a hard thermoplastic material and a flexible

thermoplastic material to create a stent precursor composition 112. The stent precursor composition is extruded to form a stent precursor thread 116. The stent precursor thread is fed into a 3D printer. The 3D printer heats, melts and then prints a stent with the melted stent precursor thread 116. Other 3D printing processes are possible with stent compositions that are not

thermoplastic. The stent includes a body 132 and a helix 128. The 3D printer can be controlled to customize the stent including the helix 128.

[70] The stent can be customized by the polymer 3D printing technique. The 3D-printing technique can print stents for the treatment of patients with different tumor contours at much lower cost than the current methods. The dimensions and compositions can be customized. This technique can print personalized stents for the special needs of different patients, for example, according to the contours of esophageal tumors based on exporting individual patient CT or other graphic data.

[71] The mixture of PLA powders (stiff material) and TPU powders (soft materials) (PLA/TPU) are thermoplastic. They are mixed and melted in a beaker and then extruded into a thread that is subsequently fed into the 3D printer. The 3D printer heats the materials at the tip of the nozzle, prints them under the software control. They are printed on a flat platform and cooled to solidify. There is no chemical curing process, and the PLA/TPU also do not chemically bond with each other. They can form a homogeneous mixture.

[72] The dense layer of the tubular body will provide stronger mechanical support to open the esophagus for drinking and food while also preventing the ingrowth of new tumor tissue and restenosis. The stiff but soft polymer will decrease the potential complications associated with the rigid metal stents, such as chest pain and bleeding. Additionally, current metal stents used in clinics lack the ability to load chemotherapy drugs. The 3D-printed polymer stent can load any drugs by coating or printing process in the next steps. The tubular body can be solid and nonporous. This will prevent tissue ingrowth, as occurs with mesh.

[73] Pig esophagi were obtained from a local farm slaughterhouse (Mary’s Ranch, Miami, Florida). The mucosa and muscular layer of the fresh

harvested esophagi were separated for the tensile tests. The structural parameters of the stent including the angle, thickness, and interval distance, were determined based on the simulated anti-migration force and radical expansion force from the finite element analysis when the stent was placed in the esophagus.

[74] To quantify the expansion force, two approaches were used, one to directly measure the compression force with a predetermined strain rate under a mechanical tester, the other to measure the expansion force in a pig esophagus between the two plates of the mechanical tester.

[75] The stents were compressed into a clinically used catheter. The stent was then pushed out of the catheter. The stent gradually self-expanded. After the stent was fully pushed out, it recovered the original shape. To further verify if the stent functioned the same in a constraint lumen, the stent was inserted into a fresh pig esophagus using the catheter. After the catheter was removed, the stent self-expanded in the esophagus. The stent obviously enlarged the size of the esophagus, compared to the esophagus without the inserted stent. The inserted stent opened the lumen of the esophagus gradually after several minutes’ post release.

[76] To quantify the expansion force, a test of self-expansion was performed by compressing the stent by 12.5%, 31.25%, and 50% strain. The

compression force with spiral stent was higher than that of the stent without spirals, and furthermore, the 3D-printed stent had at least twice higher expansion force than the SEMS stent, which had 3 N. In the harvested pig esophagus, the expansion force has a similar pattern. The results showed that the compression force in all the stents increased with the strain. Fig. 10 is a plot of Compression Force (N) vs. strain for different stent compositions and the SEMS. When the stent was compressed by 50%, the expansion force of the SEMS was 5.4 N, but the 15PLA85TPU stent without spirals was 13.8 N. The expansion force further decreased to 12.8 N on 10PLA90TPU, 11.4 N on 5PLA95TPU, and 8.2 N on 100TPU. The expansion force of the

15PLA/85TPU stent with spirals increased to 17.7 N (15.7 N on 10PLA90TPU, 12.6 N on 5PLA85TPU, 10.4 N on TPU). The compression force with the spiral stent was higher than that of the stent without spirals, and furthermore, the 3D-printed stent had at least twice as high of an expansion force as that of the SEMS stent at every stain degree. [77] Fig. 11 shows the stent compression force when the stent was expanded in a pig esophagus. The 3D-printed PLA/TPU stent, regardless of the percentage of PLA and spirals, had higher compression force than the SEMS (1.9 N). These results indicate that the 3D-printed elastic/stiff polymeric stent can open a blocked esophagus similarly to the SEMS used in this study.

[78] In the ex vivo pig esophagus, the expansion force had a similar pattern. Fig. 11 is a plot of Self-expansion Force (N) for different stent compositions and the SEMS, and with and without spirals. Fig. 11 shows the stent expansion force when the stent was released from the guiding catheter in the esophagus. After release, the 15PLA85TPU stent with spiral had a 6.8 N self expansion force. With the content of PLA decreasing, the self-expansion force decreased to 5.2 N on 10% PLA, 3.9 N on 5% PLA, and 2.8 N on TPU alone. The self-expansion force of the spiral stents with spirals was higher than that of the stents without spirals. The maximum self-expansion force of 15% PLA at the set strain degree was about 5.7 N, while it was 4.8 N on 10% PLA, 2.5 N on 5% PLA, and 2.1 N on 0% PLA. The 3D-printed PLA/TPU stent, regardless of the percentage of PLA and spirals, had higher self-expansion force than that of the SEMS (1.9 N). These results indicate that the 3D-printed elastic/stiff polymeric stent can open a blocked esophagus as the SEMS does.

[79] Fig. 12 is a schematic diagram illustrating an anti-migration force testing apparatus and procedure. To measure the anti-migration force of the stent 140, fresh pig esophagi of 10 cm length were used. One end of a pig esophagus 144 was held by the bottom grip 148 of the tensile tester. The stent was deployed and released in the esophagus by the esophagus catheter. Afterwards, the other end of the esophagus was pulled up and fixed on a supporting scaffold by threads 160 to maintain the lumen of the

esophagus open and straight. Three holes were made in the upper flange of the stent so that the stent could be hung by threads 152 to the upper grip 156 of the tensile tester showed in Fig. 12. The stent was pulled up from the esophagus at 200 mm/min set from the machine, and the force was recorded.

[80] Results show all the polymer stents with/without spirals have significant higher anti-migration forces than the SEMS stent. Fig. 13 is a plot of Anti migration Force for different stent compositions and the SEMS, and with and without spirals. In the polymer stent with spirals, the maximal anti-migration force was 3.6N (10PLA/90TPU group), and the maximal force in the non-spiral stents was 3.1 N (10PLA/90TPU group), but the SEMS stent has only 0.75 N. The 3D-printed polymer stents had at least four-fold higher anti-migration force than the SEMS stent. In the spiral stents, with the content of PLA increasing, the anti-migration force significantly increased until the PLA reached 10%, and then decreased. The same trend was seen in the polymer stents without spirals. Furthermore, the spiral stent significantly increased the anti-migration force, compared to the forces produced by stents with the same composition without spirals (Fig. 13).

[81] To further test if the spiral helix can prevent the slippage, the slipping distance of the stent was measured in the pig esophagus when exposed to the external contraction force. Fig. 14 is a schematic diagram illustrating a migration distance testing apparatus and procedure. The stent 170 was positioned in the esophagus 174. The esophagus 174 was secured by a bottom grip 178 and by threads 182. To measure the migration distance of the stent when exposed to external radical forces, a blood pressure gauge 186 was used to wrap the stent tightly. A compressive force 192 was then applied to it. The stent was radically compressed around 8 mm at the frequency of once per 4 seconds to mimic the peristaltic contractions often occurring in the esophagus during a series of wet swallows. The times of contractions was 30 and the total time lasted 2 minutes. The displacement of the stent in the esophagus was then measured.

[82] Fig. 15 is a plot of Migration Distance (mm) for different stent

compositions and the SEMS, and with and without spirals. Fig. 15 shows that the SEMS migrated 8 mm. The TPU stent without spirals migrated about 3.5 mm. With the addition of PLA into TPU, the migration distance decreased to 1.7 mm in 5% PLA, 2.4 mm in 10% PLA, and 1.9 mm in 15% PLA. However, all the PLA/TPU stent with spirals migrated less than 1 mm. Compared with the SEMS stent, the PLA/TPU stent with spirals decreased the migration distance at least eight times. Spiral significantly prevented the slippage in the pig esophagus (Fig. 15).

[83] After the stent was released into the pig esophagus, the diameter changes of the esophagus in the middle part and at the two ends of the stent were measured. Fig. 16 is a plot of the Diameter of the esophagus with the stent inserted in the middle (mm) for different stent compositions and the SEMS, and with and without spirals. Fig. 17 is a plot of the Diameter of the esophagus at the ends (mm) with the stent inserted for different stent compositions and a SEMS, and with and without spirals. The results show that the diameter of the esophagus in the middle part and at the two ends of the stent increased with the increase of the PLA content (Fig. 17). The 15PLA85TPU stent expanded the esophagus larger than the SEMS and all other stent groups. The 10PLA90TPU stent with spirals opened almost a similar diameter of the pig esophagus as the 10PLA90TPU stent without spirals (Fig. 16). Other groups including 5PLA95TPU and TPU alone with/without spirals had less space to open the same diameter as the SEMS did. From these results, although the diameter of the 3D-printed stent for this study was smaller than that of the SEMS, the lumen size of the esophagus opened by the 3D-printed stent was similar to or larger than the lumen size opened by the SEMS. Thus, the 3D-printed stent with this size can be delivered by using the currently available stent delivery system.

[84] In vitro degradation of the stent is important. Experiments were performed to measure the degradation properties of the stent materials. Since the stent will be used in the body, the degradation rate of the stents in body fluid (neutral) and gastric acid (GC, the pH of GC is 1.5 to 3.5) was tested.

The degradation of PLA/TPU composite materials in phosphate buffered saline (PBS, pFI=7.4) and simulated gastric fluid (SGF, without pepsin) at 37 ^QC was characterized over a 16-week period. Four different composition ratios of PLA/TPU were 3D printed into 5x5 mm square plates. To measure the hydrolytic degradation, dry PLA/TPU plates were weighed and placed in closed tubes containing 2 ml_ of a PBS and SGF at 37 °C for 16 weeks under a sterile condition. At the end of the designed time point, samples were taken out from the degradation medium and washed. The samples were then freeze-dried for five days. The remaining mass was weighed, and the weight loss was calculated as weight loss (%) = 100 c (Wi - W₂)/Wi, where Wi and W₂ are the weights of the printed square plates before and after degradation, respectively. Fig. 18 is a plot of remaining mass (g) vs. time (week) for degradation in PBS and SGF solutions. Fig. 19 is a plot of Weight loss (%) vs. Time (week) for degradation in PBS and SGF.

[85] Results showed that there was no observable difference in weight between the pre and post degradation after the first 6 weeks. Afterwards, the materials started to degrade slowly until the end of the experimental period (Fig. 18). After 16 weeks’ degradation, the total weight loss was around 1.5% (Fig. 19). There is no significant difference in the degradation rate and weight loss between neutral PBS and acid SGF. The subtle change of weight between samples indicates that degradation was very slow over the 16-week period. The mean and standard deviation results were compared using the T- test and it was concluded that there is no statistically significant difference between the samples before 12 weeks (p> 0.05), and the degradation seemed to start a little quicker after 12 weeks.

[86] Fourier Transform Infrared (FTIR) spectra analyses of the stent materials before and after degradation in PBS and SGF were performed using a Jasco FT/IR-4100 spectrometer. Figs. 20 and 21 show the FTIR spectra of the degraded polymer materials. The band around 3300 cm^-1 (C=0 overtone in PLA and TPU) was weakened over time, but it remained. Since PLA and TPU have C=0 stretches at around 1750 cm ¹, it could be hard to distinguish one from another. The peak did not decrease or shift at all with time. Also, the band at around 1070 cm^-1 was indicative of a C-O-C stretch in an ester. An ester was in both TPU and PLA, and that did not degrade or shift with time. If TPU and PLA were undergoing hydrolytic degradation, then the C-O-C stretch should shift to a lower frequency and that did not happen. This indicates that the TPU/PLA combination was not undergoing dramatic hydrolytic degradation and it was quite stable up to 16 weeks. Scanning electronic microscopy (SEM) observation showed that the surface of the printed stent was smooth, and the printed polymer had isotropic structure. After 16 weeks of degradation, the surface started to form nano-scale cracking lines due to the degradation. This surface morphological change may potentially affect the mechanical properties. Figure 22 is a plot of Compression force (N) vs. Strain (%) before degradation and after degradation of 1 , 2, and 3 months. The compression strength test result showed that the compression force of the stent had no significant decrease after the degradation of one month (1 M), two months (2 M), and three months (3M) (Fig. 22), but it showed that there was a trend to slightly decrease. According to the mass loss, the mechanical properties of the stent would potentially slowly decrease with time, but not dramatically.

[87] To examine whether the stent materials have no cytotoxicity, primary human esophagus epithelial cells were used. The viability of human primary esophageal epithelial cells (isolated from normal human esophageal tissue), which were obtained from Cell Biologies Co. (Chicago, USA), was studied. Four different composition ratios of TPU/PLA plates were printed. The plates were sterilized by ethanol and PBS followed by exposure to UV light source for 1 h. After sterilization, the plates were soaked in a culture medium

overnight. Epithelial cells were cultured in basal medium containing 10% fetal bovine serum, 1 % L-Glutamine and Antimycotic, 0.1 % EGF and

Hydrocortisone (Human Epithelial Cell Medium Kit, CellBiologics). After cells reached confluence, 4x10⁴ cells were suspended and seeded onto a tissue culture plate (TCP) as a control and four types of PLA/TPU discs (diameter 15 mm, thickness 1 mm) were placed in a 24-well plate. After 1 , 3, and 7 days of culturing, the viability and proliferation of epithelial cells were determined by a MTT (3-[4, 5-dimethylthiazol- 2-yl]-2,5 diphenyltetrazolium bromide) assay. Firstly, 50 mI of MTT solution (5 mg/ml_) were added to the culture wells and they were incubated at 37 °C and 5% CO2 for 4 h. The upper medium was removed cautiously, and wells were washed by PBS two times. Then the intracellular formazan was solubilized by adding 200 pi DMSO to each well. After removing the discs from each well, the absorbance was measured at 570 nm with a spectrophotometer (Molecular

Devices Spectra MAX 190). At the same time, Live/Dead staining (Live/Dead Kit, ThermoFisher Scientific) was performed after 3 days. A statistical ANOVA with Tukey’s procedure as the post hoc test was used to evaluate the statistical differences for all assays. When the p value is less than 0.05, significant differences were considered. [88] Figs. 23 A - 23 D are depictions of live/dead staining of human primary esophagus epithelial cells for a TPU stent (Fig. 23 A); a 5PLA95TPU stent (Fig. 23 B); a 10 PLA90TPU stent (Fig. 23 C); and a 15PLA85TPU stent (Fig. 23 D). Fig. 23 A-D show the results of cell proliferation assay of TPU/PLA discs at different incubation times using a MTT assay. Results also showed the effects of different groups on cell viability using a Live/Dead staining. Figs. 23 A - 23 D show that very few dead cells were observed. On 10PLA/90TPU discs, almost no dead cells were observed. Cells grew well on the surface. MTT results show that cells proliferated on the four groups of stents with time. Both MTT and Live/Dead staining results show that human esophagus epithelial cells grew and proliferated on the PLA/TPU discs. This further implied that the stent materials are biocompatible.

[89] In order to examine whether the stent materials have any cytotoxicity, primary human esophagus epithelial cells were used. Live cells were stained green, and dead cells were stained red. Fig. 23 A-D show that very few dead cells were observed. On 10PLA90TPU discs, few dead cells were observed. Cells grew well on the surface. Fig. 24 is a plot of Intensity of Absorbance vs. Time (days) for different stent compositions. The rate of cells proliferation on the 10PLA90TPU discs was significantly higher than the rates on other groups at day 3 and 7 (Fig. 24). Both MTT and Live/Dead staining results together showed that human esophagus epithelial cells grew and proliferated on the PLA/TPU discs. This further implied that the stent after the printing processes is biocompatible. [90] The experimental results show that the expansion force of the 3D- printed polymer stents was twice as high as the SMES stent. The anti migration force of the 3D-printed stent with upward spirals was four times higher than the anti-migration force of the metal stent. Furthermore, the new stent with spirals significantly decreased the migration distance compared to the migration distance of the SEMS and the 3D-printed polymer stents without spirals. Degradation study showed that the polymer materials started to degrade after 6 weeks and the compressive strength of the stent was not significantly changed. In vitro cell viability results further indicated that the polymer stent did not have any cytotoxicity. Together, these results showed that the 3D-printed stent with upward spirals has promising potential to treat esophageal malignancies with less risk of migration.

[91] Unlike to SEMS or SEPS, the biodegradable polymer stents do not require endoscopic removal. A polymer stent is inserted into the esophagus as a“bridge to surgery” or palliative treatment to allow enteral nutrition until neo adjuvant treatment is completed. In vitro study showed that the new

esophageal stent has the capacity to prevent stent migration.

[92] In order to determine whether the material properties of the stent could be affected by the potential hydrolysis, the biodegradable properties of the stent materials were also investigated. Furthermore, cell biocompatibility of the stent was characterized.

[93] The thickness of the helix spiral is another parameter that determines the expansion force. The simulation results showed that the expansion force increases with thickness. Spirals with a 1.5 mm thickness increased a 2-fold higher expansion force than the thinner spirals of 0.3 mm when the stent was compressed to 1 mm displacement. When the circumferential expansion force of the stent significantly increases beyond the bearing force of the epithelium, it may cause bleeding or other complications. To avoid the occurrence of this situation, the maximum resistance force of the esophagus was tested. The mucosa and muscular layer of the esophagus were separated first, and then their tensile strengths were tested individually. Results showed that the maximum tensile strength of the porcine esophagus mucosa-submucosa layer was 210 KPa stress. After computational simulation based on the Mooney- Rivlin model and finite element model, the maximum radial force that the esophagus mucosa layer could resist from the stent was 52 KPa. Fig. 25 is a plot of Stress (^*10⁷ Pa) vs. stretch for different stent compositions, for experimental data (ED) and engineering stress (ES). Fig. 26 is a plot of Stress (^*10⁵ Pa) vs stretch for Mucosa-Submucosa ED, Mucosa-Submucosa curve fit, Muscle ED, and muscle curve fit. To analyze these parameters using the finite element analysis software, pig esophagi were utilized to obtain the boundary conditions of this stent/esophagus system.

[94] The computational simulation result in Fig. 25 showed that the maximum expansion force of the stent 15PLA/85TPU (15 wt. % PLA, 85 wt. %TPU) with a thickness of 0.9 mm was about 52 KPa, while the force was 36 KPa when the thickness of the spiral groove was 0.6 mm. For safety consideration, therefore, 0.6 mm was selected as the thickness of the spiral groove. [95] To demonstrate the flexibility and recovery ability of the stent, the 3D- printed PLA/TPU stent was loaded in a clinically used esophagus delivery catheter. The catheter is currently used in clinics for deploying a SEMS. The inner diameter of the delivery system is around 8 mm. After the 3D-printed polymeric stent was compressed into the catheter, it was then pushed out.

The self-expanding ability was digitally recorded. As controls, a SEMS (Nitinol mesh with atraumatic ends, the middle diameter of the stent was 20 mm, and the flange diameter was 26 mm) and TPU/PLA stents without upward spiral grooves experienced the same tests.

[96] The PLA/TPU stent’s dense structure not only provides enough recovering energy, but also holds the released energy in the longitudinal direction, compared to the mesh structure of the SEMS stent. In addition to the different self-expansion mechanisms, the expansion force of the polymer stent can be adjusted by the ratio of PLA to TPU in the PLA/TPU stent. When the PLA concentration increases, the expansion force increases. Therefore, unlike the SEMS stent, this new 3D-printed polymer stent obtains the expansion force by the stent’s self-recovery energy and material strength, while a mesh stent obtains radial force through the shrinkage of the elongated lumen. Many studies have been performed to achieve enough expansion force of a stent; however, most of the stents obtain radial force by the expansion of the meshes. Once a mesh-like stent is exposed to an

esophagus that has frequent contraction forces, the SEMS stent may elongate along the longitudinal direction, thus resulting in a stent that cannot provide enough radial force after releasing the energy. It then migrates down to the stomach. Additionally, the expansion force of the stent cannot enlarge beyond the tolerance of the esophagus.

[97] Stent migration is another complication with current stents, particularly fully covered self-expanding metal stents. This complication can be avoided by suturing these stents in place using an endoscopic sewing device (Apollo Overstitch device), but this is a time-consuming and technically difficult procedure. SEMS stents also use double-flared designs. Another tapered design is also used to theoretically reduce migration with a proximal flare (30 or 24 mm) and a gradual distal taper (20 or 16 mm). To prevent the slippage of the stent into the stomach, the spirals are used as anchors to increase the friction between the stent and the epithelial layer of the esophagus. Results proved that the new structural design increased the anti-migration force and significantly decreased the migration distance. These results imply that there is a promising potential to overcome the issue of stent migration. Most interesting, all of the 3D-printed PLA/TPU stents have at least four-fold higher anti-migration forces than that of the SEMS stent, and the printed stent with spirals significantly increases the anti-migration force, compared to the same composition stents without spirals. The main reason is that the spiral can provide a higher friction force than non-spiral stents. It seems that the spirals can grip tightly on the epithelium of the porcine esophagus ex vivo. The simulation results also show that the spiral thickness and spiral number showed a great effect on the anti-migration force. The optimized pattern of the spirals increases the friction between the stent and the internal wall of the esophagus, and the added friction decreases the potential risk of migration after placement. Additionally, the anti-migration force significantly increased until the PLA content reached 10%, and then the anti-migration force decreased. The reason may be due to the decrease in flexibility with the increasing ratio of the stiff PLA contents. Higher flexibility may make the stent pliable to provide more contact areas between the surface of the stent and the inner surface of the esophagus. When the esophagus compresses the stent due to contractions, the stent with an appropriate spiral density and material composition can provide the effective contact surface of the PLA/TPU stent contacting the esophagus. Thus, the migration issue was overcome by the spiral design and adjustable flexible/rigid polymer materials to accommodate the esophageal mobility.

[98] The biodegradation property of the new 3D-printed stent provides another advantage over SEPS and SEMS. Less than 15% PLA was added to obtain a certain degradation rate of the whole stent. The amount of PLA provides not only the slow degradation ability of the whole stent, but also the appropriate expansion force and anti-migration force. The slow degradation did not significantly deteriorate the compressive property since it did not dramatically induce the mass loss. However, whether the mechanical properties will dramatically decrease with time in the in vivo implantation of a living animal or in a clinical human patient remains unknown, although our degradation study was performed in simulated gastric acid. The slow degradation slowly decreasing the mechanical expansion force could efficiently make the stent fit compatibly with the epithelium of the esophagus, thus reducing the risks of bleeding and other complications. The biocompatible and biodegradable stent may not need the re-intervention for the removal of the stent if the stent migrates into the stomach as the materials can degrade; the other potential of slow degradation of the stent is to encapsulate anti-cancer drugs into the stent when the stent is printed, thus the degradation of the small amount of the PLA part of the polymer can slowly release anti-cancer drugs into the esophagus. Therefore, the degradable 3D- printed polymer stent has not only the mechanical support to open the blocked esophagus, but also the potential to provide chemotherapy at the same time. Furthermore, the stent has the potential to combine with other therapies, such as radiotherapy, as an early support for drinking and feeding to promote longer survival time.

[99] Additionally, the 3D printing technique brings potential to produce polymer stents of any size and structure in a short time. This technology provides an advantage over the current traditional technologies for the preparation of polymer stents, such as braided, knitted, segmented, and laser- cut. With the advantages of the 3D printing, this newly developed stent with the optimized structure can be modifed to accommodate different sizes for patients.

[100] The design of the PLA/TPU stent provides excellent expansion and anti migration properties, which brings promising potential not only for the treatment of malignant esophagus cancer, but also benign esophageal stricture. The expansion and migration experiments were performed on harvested pig esophagi ex vivo, which may not truly reflect the situation of the esophagus in a living animal where the esophagus has many random contraction forces, these experiments provide evidence of efficacy.

[101] This invention can be embodied in other forms without departing from the spirit or essential attributes thereof. Accordingly, reference should be made to the following claims to determine the scope of this invention.

Claims

CLAIMS We claim:

1. A stent comprising a flexible, shape-retaining tubular body defining an internal lumen and having first and second ends and an outer surface, a helical spiral projecting from the outer surface and extending around the tubular body from the first end to the second end.

2. The stent of claim 1 , wherein the stent is made from a material having a hardness of from Shore A 55 and Shore A 90.

3. The stent of claim 2, wherein the stent comprises a combination of a hard material and a soft material, the hard material having a Shore A hardness greater than the Shore A hardness of the soft material, the combination having a Shore A hardness between Shore A 55 and Shore A 90.

4. The stent of claim 1 , further comprising end flanges at the first and

second ends, the end flanges extending axially and laterally outwardly from the first and second ends.

5. The stent of claim 1 , wherein the stent comprises a stent mixture of a thermoplastic soft material and thermoplastic hard material.

6. The stent of claim 5, wherein the hard material is at least one selected from the group consisting of polylactic acid (PLA, PDLA, PLLA isomers), polycaprolactone, polyvinyl acetate, and polyethylene glycol.

7. The stent of claim 5, wherein the soft material is at least one selected from the group consisting of thermoplastic polyurethane (TPU) and thermoplastic copolyester.

8. The stent of claim 5, wherein the stent mixture comprises 3-20% hard material and 97-80% soft material, based on the total weight of the mixture.

9. The stent of claim 5, wherein the hard material is biodegradable and biodegrades at a faster rate than the soft material.

10. The stent of claim 1 , wherein the stent mixture further comprises a drug capable of release from the stent.

1 1. The stent of claim 1 , further comprising a drug-loaded layer on the

surface of the stent.

12. The stent of claim 1 , wherein the spiral pitch of the helix is from 2 to 7 mm.

13. The stent of claim 1 , wherein the thickness of the helix is from 0.3 mm - 1.5 mm.

14. The stent of claim 1 , wherein the depth of the helix is from 0.3 mm -1.5 mm.

15. The stent of claim 1 , wherein the thickness of the stent is from 200 prn- 800 pm.

16. The stent of claim 1 , wherein the length of the stent is from 10 mm-100 mm.

17. The stent of claim 1 , wherein the diameter of the internal lumen is from 4 mm-26 mm.

18. The stent of claim 1 , wherein the stent provides a frictional force in a pig esophagus of from 0.4 to 4 N.

19. The stent of claim 1 , wherein the stent provides a self-expansion force of from 2 N to 15 N.

20. The stent of claim 1 , wherein the stent provides a compression force of from 1 N to 30 N.

21. The stent of claim 1 , wherein the stent provides an anti-migration force of from 1 N to 10 N.

22. The stent of claim 1 , wherein the stent migration distance is from 0 mm to 5 mm.

23. The stent of claim 1 , wherein the tubular body is nonporous.

24. A method for supporting a passageway of a patient, comprising the steps of:

providing a stent, the stent comprising a flexible, shape-retaining tubular body defining an internal lumen and having first and second ends and an outer surface, a helical spiral projecting from the outer surface and extending around the tubular body from the first end to the second end;

positioning the stent in the passageway.

25. A method of making a stent, comprising the steps of:

heating and mixing a hard thermoplastic material and a flexible thermoplastic material to create a stent precursor composition;

extruding the stent precursor composition to form a stent precursor thread;

cooling the stent precursor thread; feeding the stent precursor thread into a 3D printer, the 3D printer heating and melting the stent precursor thread;

printing a stent with the 3D printer and the melted stent precursor thread.