Patent Application
20230390090
The present invention relates, in general terms, to stents formed from Nitinol. The present invention also relates to methods of 3D printing Nitinol stents.
Stents are small tube-like surgical devices used by surgeons to unblock or widen clogged arteries to restore regular blood flow for treatment of patients with vascular diseases. Traditionally, stents are made of a biocompatible Stainless Steel or metal alloy.
Stent sizing and apposition have been shown to be important determinants of clinical outcome. Undersized stents tend to induce thrombosis in the longer term whereas oversizing increases the vessel wall stress and may induce inflammatory response, which, in turn, contributes to neointimal hyperplasia. Failure to achieve predicted stent diameter is a common problem for stents made of stainless steel and cobalt chrome when deploy using semi-compliant balloons. Therefore, selection of proper stent size relative to the target vessel should be considered as important as post-deployment optimization strategy.
In the pre-stent era restenosis ranged between 32-55% of all angioplasties and drop to 17-41% in the bare metal stent (BMS) era. A further step to reduce restenosis was undertaken with the introduction of drug-eluting stents (DES), with a reduction to numbers <10%. DES are used to counter In-Stent Restenosis by improving blood flow and decrease the likelihood of repeating procedures to reopen blocked blood vessels compared to uncoated devices. However, in recent times, the U.S Food and Drug Administration (FDA) has expressed concerns on whether Paclitaxel-coated balloons and stents have a long-term adverse effect such as late mortality. Drugs use serve as a preventive solution not rectifying the root cause of In-stent restenosis and stent thrombosis which is mainly due to vascular injuries caused by over-expanded stents over straining the vascular walls.
Although polymer bioresorbable stents have previously been made, polymer-based scaffolds display inferior mechanical performance compared with metallic drug eluting stents (DES). Polymeric surfaces often offer less than ideal conditions for endothelial cell migration, thrombogenic resistance, inflammation, and vessel wall healing.
Other challenges remain. For example, failure of vessel wall healing has been attributed either to the use of polymeric stent coatings, or due to the effects of the eluted drug.
Hemodynamics is the dynamics of blood flow, and in medical contexts it often refers to basic measures of cardiovascular function, such as arterial pressure or cardiac output. Stents, which are deployed to reopen stenotic regions of arteries and to restore blood flow, have risks of causing inflammation and localised stent thrombosis that would result in a stent failure. Stent edge restenosis, the formation of a neointima that gradually re-narrows the arterial lumen, is recurrent in 30-40% of patients receiving BMS. Even though there are more advanced DES that release anti-proliferative drugs to successfully address the problem of restenosis, DES recipients are still significantly associated to late-stent thrombosis (LST).
Stent thrombosis is a thrombotic occlusion of a coronary stent, which is the formation of blood clot inside the blood vessels, resulting in the obstruction of blood flow, and it is an acute process in contrast to restenosis. Studies showed strong correlations between LST and the lack of endothelial coverage. The endothelium is a single layer of endothelial cells lining the vascular walls and plays an integral part in maintaining vascular homeostasis. Stenting causes significant damage to the vascular wall and endothelium, resulting in inflammation, repair and the development of neointimal hyperplasia. The ability of the endothelial to repair itself is dependent on the migration of neighbouring mature endothelial cells and the attraction of circulating endothelial progenitor cells to the injured area, which then differentiate into endothelial-like cells.
Endothelialization of the stent's strut surface is inversely proportional to the thickness of the stent and areas with largest flow separation zone. The shape of individual struts promotes blood flow separation that creates recirculation zones, and slower flow in a recirculation zone yields lower shear rates that retards endothelialization. Recirculation zones can also serve as micro-reaction chambers where procoagulant and pro-inflammatory elements from the blood and vessel wall accumulate.
Accordingly, there is a need to develop new stent designs. There is also a need to develop stents using other materials which are more suitable.
Stents can be fabricated with various techniques, such as etching, micro-electro discharge machining, electro-forming and die-casting. In the market currently, most stents are fabricated using laser cutting technology such as micro-laser machining technology. The process involves a high energy density laser beam focusing on the workpiece surface, where the thermal energy that is absorbed heats and transforms the workpiece volume into a molten, vaporized or chemically changed state that can be easily removed by the flow of a high pressure assist gas jet.
However, laser cutting affects the quality issue of the workpiece due to long pulse duration and melting. This long pulse machining results in heat affected zone (HAZ) generation at the vicinity of the cut regions, as well as significant slog, dross, recast and oxide layer, and back wall damage that will require post-processing. Advancement of the laser industry has driven the development of shorter pulse lasers that would offers opportunities to minimise and eliminate HAZ. However, fabricating stents through laser cutting still requires system optimisation to achieve high yield, high quality and low cost.
There has been recent interest in using Nitinol for making stents. Nitinol is a metal alloy of Nickel and Titanium, where the two elements are present in roughly equal atomic percentages. It exhibits shape memory, superelasticity, good corrosion resistance and good biocompatibility. As such, it is highly sought after for medical applications. The global market for Nitinol-based medical devices is expected to grow at a compound annual growth rate of nearly 8.2% from 2017 to 2025. Reason for the rising global Nitinol medical devices market is the prevalence of cardiovascular diseases, growing population susceptible to peripheral artery diseases and increasing demand for minimally invasive surgical procedures. As per statistics of the World Health Organization, almost 17.7 million individuals suffer from cardiovascular diseases each year; cardiovascular diseases account for 31% deaths each year globally. Advantageously, the treatment of iliac artery occlusive disease with self-expandable stent as compared with Balloon expanded stent resulted in a lower 12-month restenosis rate and a significantly reduced TLR rate.
However, much difficulty is faced with processing Nitinol structures which are sufficiently thin for use in biomedical application such as stenting. Some reasons include surface quality, mechanical properties and biocompatibility. Partially unsintered powder, especially Nickel, could remain on the stent surface or be released into the bloodstream, which could have an adverse biological effect.
Additionally, as laser machining employs a high energy laser beam to precisely heat, melt and vaporise a tube of Nitinol material, heat affected zones are created. Not only does this create surface defects such as burrs and dross formation, but it also affects the microstructure.
While an ultrashort pulse laser can produce a dross-free cut of Nitinol, ultrashort pulse laser machining processes have low cutting efficiency. In addition, debris and recast formation from the vaporized material is still required to be removed by other methods.
Moreover, the tube-based cross section patterns and non-streamlined laser-moving paths in laser-machining have further reduced the design freedom of stents, raising the risk of stents thrombosis originated from blood flow separation.
Accordingly, there is a need to develop new technologies to produce stents. Even when a new technology is used to produce stents, there are concerns that the stents will not perform as expected due to the change in fabrication process.
It would be desirable to overcome or ameliorate at least one of the above-described problems, or at least to provide a useful alternative.
The present invention is predicated on the understanding that additive manufacturing (AM) can be a more economical solution to fabricate a high cost stent. In this method, instead of using high heat energy to vaporise the material, a laser is used as a heat source to precisely fuse metal powder particles together, layer by layer. In particular, while selective laser melting (SLM) has been used for making structures from metals such as steels, aluminium and titanium, a stable process for Nitinol has yet to be established. There is also a considerable mismatch in the fabrication parameters with feature sizes of earlier reported SLM processed nitinol stents. There is also a need to develop methodologies for creating features of about 100-300 μm, or wires with a diameter of less than 1 cm, or preferentially less than 0.5 mm. Without wanting to be bound by theory, the inventors have found that stents having wires with less than 1 cm can be reliably printed with superelastic properties and/or shape memory properties by controlling at least the laser power and scanning speed when printing. Further, additional considerations (such as boundary, downskin and upskin processing, support fabrication, and special geometries, e.g., thin-walled constructs) need to be taken into account to create a stent that has Nitinol's inherent unique properties of shape memory effect and superelasticity, as well as extrinsic minimal surface roughness and porosity, apart from the basic mechanical properties. With proper selection of process parameters, heat affected zones can be eliminated or minimized. To this end, SLM offers great potential for producing Nitinol stents.
Additionally, stents fabricated by AM have the capability to deliver customized parts which may not be feasible and cost effective using conventional manufacturing methods. This is favourable for medical applications in which patient-specific implant design can be realised to ensure a better anatomically fit thus promotes faster healing.
The present invention provides a method of 3D printing a stent, comprising: performing selective laser melting on a Nitinol powder in order to form the stent, wherein selective laser melting is performed with:
It was found that when 3D printing using these conditions, the resultant stents have shape memory and/or superelastic properties. In particular, the 3D printed stent can be customised such that it is curved and which retains its curved configuration due to the shape memory property. The 3D printed stent can be crimped into a linear configuration and inserted into a delivery tube for deployment at a target site. When deployed at about human body temperature, the crimped stent reverts back to its original printed size and shape, and is further superelastic in the sense that it can maintain its size and shape after removal of external forces such as muscle contractions. In contrast, conventional stents are produced by laser cutting, and can only be made with superelastic properties. Conventional stents are also made with a linear configuration due to the difficulties in producing curved stents.
In some embodiments, the selective laser melting is performed with:
In some embodiments, when the selective laser melting is performed using conditions in (i), the 3D printed stent is characterised by a As temperature of about −45° C. to about −25° C.
In some embodiments, when the selective laser melting is performed using conditions in (i), the 3D printed stent is characterised by a Ms temperature of about −10° C. to about 10° C.
In some embodiments, when the selective laser melting is performed using conditions in (ii), the 3D printed stent is characterised by a As temperature of about −30° C. to about 0° C.
In some embodiments, when the selective laser melting is performed using conditions in (ii), the 3D printed stent is characterised by a Ms temperature of about −10° C. to about 25° C.
In some embodiments, when the selective laser melting is performed using conditions in (i) or (ii), the 3D printed stent is characterised by columnar grains due to inter-layer over melt.
In some embodiments, when the selective laser melting is performed using conditions in (iii), the 3D printed stent is characterised by a As temperature of about 10° C. to about 75° C.
In some embodiments, when the selective laser melting is performed using conditions in (iii), the 3D printed stent is characterised by a Ms temperature of about 60° C. to about 100° C.
In some embodiments, when the selective laser melting is performed using conditions in (iv), the 3D printed stent is characterised by a As temperature of about 10° C. to about 80° C.
In some embodiments, when the selective laser melting is performed using conditions in (iv), the 3D printed stent is characterised by a Ms temperature of about 20° C. to about 70° C.
In some embodiments, when the selective laser melting is performed using conditions in (iii) or (iv), the 3D printed stent is characterised by fully merged layer boundaries due to re-melt of an underlying layer.
In some embodiments, the selective laser melting is performed with a hatch distance of about 0.1 mm to about 0.5 mm.
In some embodiments, the selective laser melting is performed with a layer thickness of about 0.01 mm to about 1 mm.
In some embodiments, the 3D printed stent has a wire diameter of less than 1 cm, preferably less than 0.5 mm.
In some embodiments, the method further comprises a step of heat treating the stent.
In some embodiments, the method further comprises a step of heat treating the stent when the stent is printed using condition ii, iii or iv.
Conventionally, stents fabricated using laser cutting require a heat treatment process called shape setting to obtain its superelastic property. Using selective laser melting and particularly the above parameters, it is not necessary for the stents to undergo heat treatment to achieve the desired properties. Rather, heat treatment may be used to further fine-tune the mechanical properties of the 3D printed stents. For example, the heat treating step causes the distribution of nickel and titanium within the stent to be re-distributed. This lowers the Ms temperature, thus allows for a transition from a crimped state to an original uncrimped state at or near human body temperature.
In some embodiments, the heat treating step comprises heating the stent from about 200° C. to about 800° C.
In some embodiments, the 3D printed stent is characterised by a austenite finish temperature (Af) of about 25° C. to about 50° C.
In some embodiments, the 3D printed stent is characterised by wires of the 3D printed stent having a partially flat cross sectional shape.
In some embodiments, the 3D printed stent is characterised by wires of the 3D printed stent having an elliptical, tear drop, partially flattened tear drop or circular cross section shape.
In some embodiments, the 3D printed stent is characterised by a curvature along its longitudinal dimension when in the expanded state.
In some embodiments, the 3D printed stent is characterised by a curvature of about 1° to about 160°.
In some embodiments, the 3D printed stent is characterised by a radius of curvature of about 1 mm to about 200 cm.
In some embodiments, the method further comprises providing a template of the stent; wherein the stent template comprises:
In some embodiments, the wave-like structure is a sinusoidal wave-like structure or a helical wave-like structure.
In some embodiments, each flex unit has a wave number of about 0.5 unit to about 2 units.
In some embodiments, the wave-like structure in each flex unit has a peak characterised by an angle of about 15° to about 90° relative to a local radial plane at the peak.
In some embodiments, when in the expanded state, each flex unit has a transverse breadth of about 2 mm to about 12 mm.
In some embodiments, when in the expanded state, each flex unit has a longitudinal length of about 5 mm to about 15 mm.
In some embodiments, a first end of at least one flex unit is connected to one of two adjacent circumferential sections by a first extension, and/or a second end of at least one flex unit is connected to the other of the two adjacent circumferential sections by a second extension.
In some embodiments, the first extension has a length of about 0.1 mm to about 5 mm and/or the second extension has a length of about 0.1 mm to about 5 mm.
The present invention also provides a 3D printed stent printed using the method as disclosed herein, the stent comprising Nitinol having a nickel content of about 54 wt % to about 57 wt % of the composition and a titanium content of about 43 wt % to about 46 wt % of the composition;
In some embodiments, when the stent has a As temperature of about −45° C. to about 0° C. and a Ms temperature of about −10° C. to about 25° C., the stent is characterised by columnar grains due to inter-layer over melt.
In some embodiments, when the stent has a As temperature of about 10° C. to about 80° C. and a Ms temperature of about 20° C. to about 100° C., the stent is characterised by fully merged layer boundaries due to re-melt of an underlying layer.
In some embodiments, the 3D printed stent is characterised by a austenite finish temperature (Af) of about 25° C. to about 50° C.
In some embodiments, the nickel is about 54.5 wt % to about 55.8 wt % the composition.
In some embodiments, the nickel is about 55.2 wt % of the composition.
The present invention also provides a stent delivery device, comprising:
As the crimped stent has shape memory properties, the stent is revertible back to its original uncrimped state when at least exposed to a temperature above its As temperature.
In some embodiments, the stent in its original uncrimped state is adapted to revert back to its original configuration after release of an external force and when exposed to a temperature of about 25° C. to about 50° C.
In some embodiments, the stent delivery device further comprises ejecting means for ejecting the crimped stent out from the tube.
The present invention also provides a method of delivering a stent in a stent delivery device into a channel, the stent delivery device comprising:
Embodiments of the present invention will now be described, by way of non-limiting example, with reference to the drawings in which:
The inventors envisioned that the stents of the present invention can be deployed in many areas within the human body. However, as a starting point, the femoropopliteal/femoral artery (FPA) is chosen for investigations for two main reasons. The FPA is a large artery located in the thigh provides majority of the arterial blood supply to the lower part of the extremity. First, the FPA undergoes one of the most extensive mechanical deformation in the human body during limb flexion, with twisting, bending and compression. The FPA experiences 2-4°/cm twist, 4-13% axial compression and has 22-72 mm bending radius during limb flexion. Measurements using intra-arterial markers showed even more severe deformation that were 2 to 7 times larger. Therefore, if the presently disclosed stent are made to handle the harsh conditions within the FPA, it is expected the stents are also suitable for use in other parts of the body. Second, as one of the largest arteries in the human body, the average common FPA has a diameter of 6.6 mm (3.9-8.9 mm), whereas the superficial FPA and deep FPA have average vessel diameters of 5.2 mm (2.5-9.6 mm) and 4.9 mm (2.7-7.6 mm) respectively. Current stents used in the FPA generally have diameters between 5 mm and 8 mm and are usually oversized compared to the diameter of the artery. Thus, stents with a range of diameters can be designed to ascertain its feasibility before scaling down for considerations in other parts of the body.
The material used to fabricate stents of the present invention is Nitinol. Nitinol is a metal alloy of nickel and titanium, where the two elements are present in roughly equal atomic percentages. Different alloys are named according to the weight percentage of Nickel, e.g. Nitinol 55 and Nitinol 60. The nickel content can be from about 40% to about 65%, and the titanium content can be correspondingly be from about 35% to about 60%. As used herein, Nitinol is used as a powder. The powder can have a particle size from about 10 μm to about 60 μm. The powder shows phase transformation peaks, i.e., the martensite start (Ms) and austenite start (As) temperatures at −17.9° C. and −5.6° C., respectively.
Nitinol alloys exhibit two closely related and unique properties: the shape memory effect and superelasticity (also called pseudoelasticity) at different temperatures. Shape memory is the ability of nitinol to undergo deformation at one temperature (generally a lower temperature), stay in its deformed shape when the external force is removed, then recover its original, undeformed shape upon heating above its “transformation temperature”. This is due to the reversion of martensite to austenite by heating, causing the original austenitic structure to be restored or reversed regardless of whether the martensite phase was deformed. The name “shape memory” refers to the fact that the shape of the high temperature austenite phase is “remembered” even though the alloy is severely deformed at a lower temperature.
Superelasticity is the ability for the metal to undergo large deformations and immediately return to its undeformed shape upon removal of the external load. Superelastic properties are generally observed when the temperature is above austenite temperature. Nitinol can deform 10-30 times as much as ordinary metals and return to its original shape. At high temperatures, nitinol assumes an interpenetrating simple cubic structure (austenite). At low temperatures, nitinol spontaneously transforms to a more complicated monoclinic crystal structure (martensite). There are four transition temperatures associated to the austenite-to-martensite and martensite-to-austenite transformations. Starting from full austenite, martensite begins to form as the alloy is cooled to the martensite start temperature (Ms), and the temperature at which the transformation is complete is the martensite finish temperature (Mf). When the alloy is fully martensite and is subjected to heating, austenite starts to form at the austenite start temperature (As), and finishes at the austenite finish temperature (Af). This is commonly shown in a cooling/heating cycle as a thermal hysteresis. The hysteresis width depends on the precise nitinol composition and processing. Its typical value is a temperature range spanning about 20-50 K (20-50° C.).
In designing and/or fabricating the stent of the present invention, a self-expanding deployment method was considered. The stent will undergo shape setting in the expanded form before being crimped at/under room temperature for insertion into the catheter for deployment. The mechanical hysteresis behaviour of nitinol results in ‘biased stiffness’, whereby a stent that recovers from the crimped position or state will be much more resistant to compression than to expansion, which is useful for reducing chronic outward force (chronic outward force correlates to a measure of the radial force the stent projects outwards in its deployed configuration).
In contrast, balloon-expandable stents which are usually made from stainless steel or cobalt-chromium has the disadvantage that when the stent is expanded, it is plastically deformed and retains a permanent geometry. There is also a perceived risk for balloon-expandable stents in arteries to be permanently deformed through outside pressure resulting in a partially or completely block vessel, once the buckling strength of the stent is exceeded.
In the stent market currently, there is no industry standard on the specifications of a stent and the properties it should have. Stents manufacturers in the United States of America (US) would have to gain approval from the Food and Drug Administration (FDA) before they can release their products commercially. The FDA has provides a guidance regarding its current thinking on non-clinical engineering test that are submitted in investigational device exemption applications and premarket approval applications to support the safety and effectiveness of intravascular stents and their associated delivery systems. The comprehensive non-binding guidance includes an array of tests such as mechanical properties and stress/strain analysis, and the manufacturer has to provide details such as the test method, accept/reject criteria, sample size and results.
Therefore, the lack of an industry standard due to differing testing results and testing conditions of different manufacturers makes it difficult to verify if the designed stent is up to par or comparable with commercially available stents in terms of mechanical properties.
Accordingly, to verify the stent's mechanical properties, comparisons with commercially available stents under different mechanical loading conditions were made. Common testing modes include radial compression, crush resistance, bending, axial tension and compression, as well as torsion.
Stents can be classified with an open-cell or closed-cell design, which is dependent on the density of the struts (
The inventors are of the opinion that the limitations of the stent fabricating method should also be considered. When using AM or SLM, the inventors have found that open-cell stents are not feasible for 3D printing due to the large overhang surfaces when printing from in any orientation. For example, when printing in orientation Y, there are excessive overhang areas which are not supported by the layer below, and which would result in a print failure (circled region in
To improve the flexibility of the closed-cell stent, it was postulated that the free cell area between each cell (in a closed cell stent) can be maximized to mimic an open-cell design, with bending features incorporated into the cell design. Not only will this provide the scaffolding required, it also promotes flexibility in this hybrid stent design. Additionally, since the strut thickness (wire diameter) and geometry can play an important role in the hemodynamic properties, the strut thickness can be minimised and the strut geometry can be optimized.
To investigate features that allowed closed-cell stents to achieve flexibility which measures up to that of an open-cell stent, features of commercial stents were studied.
The inventors have found that the flex sections should be designed such that the overhang areas are reasonable for 3D printing. The flex sections in commercial stents cannot be adopted for 3D printing as it has excessive overhang areas and small features that are spaced very closed together.
In an aspect, the present invention provides a 3D printed stent having a longitudinal dimension and a radial plane, the stent movable from a collapsed state to an expanded state, the stent comprising:
In some embodiments, the 3D printed stent comprising:
The circumferential sections 406 and 410 are connected to each other by a flex section 408. Accordingly, flex section 408 is sandwiched between or disposed between circumferential sections 406 and 410. The flex section 408 extends between the two adjacent circumferential sections 406 and 410. The flex section 408 is movable along the longitudinal dimension 402 when transiting between the collapsed state and the expanded state. In this sense, the flex section is longitudinally expandable in order for the stent to move from the collapsed state to the expanded state.
It should be noted that the stent in moving from a collapsed state to an expanded state has an overall increase in length in the longitudinal dimension. Thus, if the circumferential section is contracting longitudinally, an increase in the longitudinal length of the flex section is greater than the decrease in the longitudinal length of the circumferential section.
As is shown in
The flex section 408 comprises a plurality of flex units. The flex units are circumferentially arranged, with its length 418 parallel to the longitudinal axis of the stent.
At this angle, the flex units are 3D printable without experiencing overhang issues. The flex unit 408a is able to expand or contract in the longitudinal direction of 418 for expanding or contracting the stent along a longitudinal axis. The flex unit 408a is also able to expand or contract in the transverse direction of 416. It is not expected that this expansion and/or contraction in the direction of 416 will change the width of the stent.
The flex unit can alternatively be characterised as having at least three bends, for example with at least two of the bends in contact with the circumferential sections. In this example, the flex unit has 4 bends 412a, 412b, 412c and 412d. In this expanded state, the bends (412a, 412b, 412c and 412d) can each independently have an angle (downskin angle) 414 relative to the radial plane 404 of about 15° to about 90°.
The “downskin angle” is illustrated as δ in
The stent is fabricated from Nitinol, a metal alloy of nickel and titanium.
To determine the flexibility and suitability of flex section for 3D printing, 7 different flex units were fabricated for comparison.
Simulation results of 7 different flex units are shown in Table 1. When a greater amplitude is provided in the model (corresponding to the transverse breadth of a flex unit as shown in for example 416) or downskin angle of the flex segment (FS) is increased, a greater maximum displacement (which translates to a longitudinal expansion) without significantly changes to the maximum stress experienced by the part is observed. The largest displacement of the FS occurs at the point where the force was applied. This translates to a maximum displacement in the longitudinal dimension of the stent. Another observation is that sharp peaks (or smaller downskin angle) in the FS reduces the maximum stress incurred without substantial reduction in the maximum stress, and this observation is more distinct in FS with greater height or downskin angle. For example, and referring to 412b and 412c, if the curvature at these bends is made sharper or smaller, the maximum stress in the flex unit can be reduced.
Therefore, flex units with sharp peaks (or smaller downskin angle) would be preferred since they produces better results by offering flexibility without incurring as much stress. Flexibility of the stent can also be altered by varying the height (breadth in the transverse plane) and/or downskin angle of the flex units, and the variation has to be within the limitations of the 3D printer since excessive overhang might cause the print to fail. Hence, incorporation of flex units of similar designs that are within the 3D printing boundaries have great potential of improving the flexibility of the initial stent design.
In some embodiments, the downskin (or upskin) angle is about 15° to about 80°, about 15° to about 70°, about 15° to about 60°, about 15° to about 50°, about 15° to about 45°, about 15° to about 40°, about 15° to about 35°, about 15° to about 30°, or about 15° to about 25°. In other embodiments, a sharp peak (or smaller downskin angle) as discussed herein refers to an angle of about 15° to about 30°. While a small downskin angle is particularly advantageous, a downskin angle of about 30° to about 60° can also be favourable for use in blood vessels with less demanding conditions.
In some embodiments, the flex section comprises a plurality of flex units. In other embodiments, each flex section comprises 5 to 12 flex units. In other embodiments, each flex section comprises 5 to 11 flex units, 5 to 10 flex units, 6 to 10 flex units, 7 to flex units or 8 to 10 flex units. In other embodiments, the flex section comprises 6 to 12 flex units, 6 to 11 flex units, 7 to 12 flex units, 8 to 12 flex units, 9 to 12 flex units, or 10 to 12 flex units.
In some embodiments, the wave-like structure is a sinusoidal wave-like structure or a helical wave-like structure.
In some embodiments, each flex unit has a wavenumber of about 0.5 unit to about 2 units. The wavenumber (also wave number or repetency) is the spatial frequency of a wave.
An angle is also present within the wave-like structure of the flex unit. When the flex unit has a wavenumber of at least 0.5 unit, a peak is present. The peak can be characterised by an angle of about 15° to about 90° relative to a local radial plane at the peak. The local radial plane of the stent is thus formed with the highest point of the peak within the radial plane.
The flex units can each independently have a periodic structure. In this case, the flex unit can have 4 bends 412a, 412b, 412c, 412d oriented along the longitudinal dimension as shown in
The flex units can each independently have a period of about 0.5 to about 2. The flex units can have a period of about 0.5, about 1, about 1.5, or about 2.
Each flex unit can comprise at least three bends. In other embodiments, each flex unit can independently comprise 3 or 4 or 5 or 6 bends. The bends in the flex units can also be considered as peaks and troughs (dependent on their orientation). The peak and trough can refer to the highest and lowest point of the periodic structure. When the flex unit comprises 4 bends, it can be considered to substantially be of a single sine wave.
When in the expanded state, the flex unit can have a transverse breadth 416 of about 2 mm to about 15 mm. The transverse breadth 416 is measured as the maximum distance of the flex unit perpendicular to the longitudinal axis of the stent. As shown in
When in the expanded state, the flex section and/or the flex units can have a length along the longitudinal axis of the stent of about 1 mm to about 15 mm, about 2 mm to about 15 mm, about 4 mm to about 15 mm, about 5 mm to about 15 mm, about 7 mm to about 15 mm, or about 10 mm to about 15 mm. The longitudinal length 418 is measured as the distance of the flex unit parallel to the longitudinal axis of the stent. As shown in
Stents with varying flex section and circumferential sections were studied to understand how these components affect stent performance. The following factors affecting stent performance were studied:
As shown in Table 2, different stent designs were modelled with varying features. For consistency in comparison, all designs have a stent diameter of 8 mm with 8 circumferential sections (or strut segments) and 7 flex sections (or flex segments). Designs 1-3 are used to compare how the cross section diameter of the flex unit (strut thickness) affects the properties of a stent, designs 1 and 4 for comparison of how the overhang angle affects the stent properties. Designs 5-7 allows a comparison between stent properties and the number of circumferential units.
Designs 1-5 and designs 5-7 have different patterns, which are patterns A and B respectively. Pattern A is illustrated in
Pattern B is illustrated in
In particular,
A flex unit 604a is illustrated in
Alternatively, the flex unit 604a comprises 4 bends 608a-d. The bends (608a-d) can each independently have an angle (downskin angle) 610 relative to the radial plane of about 15° to about 90°.
In some embodiments, the first extension and the second extension each independently has a length of about 0.1 mm to about 5 mm. In other embodiments, the length is about 0.1 mm to about 4.5 mm, about 0.1 mm to about 4 mm, about 0.1 mm to about 3.5 mm, about 0.1 mm to about 3 mm, about 0.1 mm to about 2.5 mm, about 0.1 mm to about 2 mm, about 0.5 mm to about 2 mm, or about 1 mm to about 2 mm.
In some embodiments, the stent comprises a combination of Pattern A flex section and Pattern B flex section. In other embodiments, one flex section comprises a plurality of circumferentially arranged flex units with no extensions, while another flex section comprises a plurality of circumferentially arranged flex units with a first extension 612. In another example, one flex section comprises a plurality of circumferentially arranged flex units with no extensions, while another flex section comprises a plurality of circumferentially arranged flex units with a first extension 612 and a second extension 614. This advantageously provides greater flexibility in personalising a stent for a patient based on his condition.
The flex units 704 can be connected to circumferential units 702 and 706 in a periodic structure. For example, the flex units in flex section 704 can be in a wave pattern, which can be in the form of a spiral, helical or double helical pattern. Each end of the flex unit in flex section 704 can be connected respectively to a circumferential unit in the circumferential sections 702 and 706. For example, the connection can be at a turning point (peak or trough) of the circumferential unit.
The various examples as disclosed herein illustrates the flexibility of the stent design when it comprises circumferential sections and at least one flex section. Further by varying, for example, the number of bends, the transverse breadth of the flex units, the number of flex units, the number of circumferential units, the longitudinal length of the sections, various stents can be fabricated to suit the specific requirements of a patient.
In some embodiments, the plurality of flex units each independently has a wire with a cross sectional diameter of about 0.2 mm to about 0.4 mm. For example, the wires of the plurality of flex units can have a cross sectional diameter of about 0.2 mm to about 0.4 mm.
Based on simulation studies, the inventors have found that a wire with a thinner cross section diameter of the flex unit (strut thickness) result in a larger displacement of the model when placed under bending, compression and tension, which translates to greater flexibility. This allows the stent to be subjected to a bending and/or compression force easily without breaking.
Similarly, flexibility can also be imparted by having wires with thinner cross section diameter of the circumferential units. In this regard, the wires forming the circumferential units can have a thinner cross sectional diameter. Further, lesser circumferential units in the circumferential sections also results in greater flexibility when placed under bending, compression and tension conditions. No distinct correlations can be observed from the displacement of the models when placed under torsion. When placed under a torsional force, designs 5 to 7 have a more even distribution of areas with higher stress concentrations across the entire stent structure, whereas the areas of higher stress concentrations are concentrated at areas closer to where the torsional force is acting for designs 1 to 4. Similarly, designs 5 to 7 have stress concentrations that are better distributed across all the flex segments across the entire structure, which is probably attributed to their uniform and symmetrical designs. The designs that are uniform also have a more even deformation model as observed from the simulation models. Therefore, it appears that a uniform circumferential section (strut) design (as provided using a periodic design) can result in a more uniform stress distribution and thus more predictable and favourable results. Flexibility of the stent in terms of bending, compression and torsion can also be manipulated by varying the strut longitudinal thickness as well as the number of circumferential units.
In some embodiments, each of the at least two circumferential sections comprises a plurality of circumferential units. The circumferential units are circumferentially arranged within the circumferential sections.
In some embodiments, the at least two circumferential sections each independently comprises about 5 circumferential units to about 12 circumferential units. In other embodiments, the circumferential section comprises about 5 circumferential units to about 11 circumferential units, about 5 circumferential units to about 10 circumferential units, about 6 circumferential units to about 10 circumferential units, about 7 circumferential units to about 10 circumferential units, or about 8 circumferential units to about 10 circumferential units. The number of units will depend on the desired diameter of the expanded stent, and to this end, the circumferential units can be adjusted to fabricate stents of different diameters.
Each circumferential section can be made up of circumferential units having a wave-like structure. For example, the circumferential units can have a sinusoidal wave-like structure. The circumferential units can be characterised by a length which can be a multiple of a wavenumber, and a peak to peak amplitude. The length forms a closed loop and extends circumferentially relative to the longitudinal axis of the stent. The circumferential units can be spaced apart from each other at regular intervals.
When the circumferential units have a sinusoidal wave-like structure, each circumferential section can have a wavenumber of about 5 unit to about 12 units.
An angle is also present within the wave-like structure of the circumferential section. The peak can be characterised by an angle of about 15° to about 90° relative to a local transverse plane at the peak. The local transverse plane of the stent is thus formed with the highest point of the peak within the transverse plane.
In some embodiments, the circumferential units each independently has a wire having a cross sectional diameter of about 0.2 mm to about 0.4 mm. In other embodiments, the circumferential units have a cross sectional diameter of about 0.22 mm to about 0.4 mm, about 0.24 mm to about 0.4 mm, about 0.26 mm to about 0.4 mm, about 0.28 mm to about 0.4 mm, about 0.3 mm to about 0.4 mm, about 0.32 mm to about 0.4 mm, about 0.34 mm to about 0.4 mm, or about 0.36 mm to about 0.4 mm.
In some embodiments, the at least two circumferential sections each independently has a periodic structure. In other embodiments, the at least two circumferential sections each independently has a periodic wave structure.
As shown in
In some embodiments, the circumferential units in the at least two circumferential sections are arranged such that the circumferential units in one circumferential section are anti-phase relative to the circumferential units in the other circumferential section. It was found that such an arrangement can provide additional support and strength to the stent, and can further reduce torsional breakage. Alternatively, the circumferential units are arranged such that they are in phase with respect to each other, or arranged such that they are substantially out of phase with respect to each other.
Depending on the circumferential sections, the number of flex units in the flex section may vary. In some embodiments, when the at least two circumferential sections comprises periodic or wave-like structures and are anti-phase with respect to each other, the stent is formed such that one end of a flex unit connects to a peak in a circumferential section and another end of the flex unit connects to a trough in another circumferential section. In other embodiments, when the circumferential section comprises periodic or wave-like structures and are in phase with respect to each other, the stent is formed such that one end of a flex unit connects to a peak in a circumferential section and another end of the flex unit connects to a peak in another circumferential section. In other embodiments, when the circumferential section comprises periodic or wave-like structures and are in phase with each other, the stent is formed such that one end of a flex unit connects to a trough in a circumferential section and another end of the flex unit connects to a trough in another circumferential section.
In some embodiments, when the circumferential section comprises periodic or wave-like structures, the stent is formed such that each peak are connected to a flex unit. Alternatively, alternate peak can be connected to a flex unit, or a third of the peaks are connected to flex units. In other embodiments, when the circumferential section comprises periodic structures, the stent is formed such that each trough are connected to a flex unit. Alternatively, alternate trough can be connected to a flex unit, or a third of the troughs are connected to flex units. In other embodiments, with the exception of the terminal circumferential section, when the circumferential section comprises periodic structures, the stent is formed such that each peak and each trough are independently connected to a flex unit. In other embodiments of the stent, when adjacent circumferential sections comprise different circumferential units, a peak of a circumferential unit on a circumferential section is connected via two (or more) flex units to two (or more) peaks or troughs of two (or more) circumferential units on the adjacent circumferential section. To this end, one end of the stent is wider than the other end. This stent variation can advantageously provide support at, for example, an intersection between a vein (or artery) and a capillary.
In some embodiments, circumferential units in the at least two circumferential sections are alternatively connected to flex units in the flex section. In other embodiments, at least 20% of the circumferential units in the at least two circumferential sections are connected to flex units in the flex section, In other embodiments, at least 30%, at least 33%, at least 40%, at least 45%, at least 50%, at least 60%, at least 66%, at least 70%, at least 80%, or at least 90% of the circumferential units in the at least two circumferential sections are connected to flex units in the flex section.
The stent comprises at least two circumferential sections. In some embodiments, the stent comprises 2 to 10 circumferential sections. The stent can comprises 3 to 10 circumferential sections, 4 to 10 circumferential sections, 5 to 10 circumferential sections, 6 to 10 circumferential sections, or 7 to 10 circumferential sections. The stent can comprise 2, 3, 4, 5, 6, 7, 8, 9 or 10 circumferential sections. The number of circumferential sections depends on the desired length of the stent, which depends on its application in the body. As the presently disclosed stent is fabricated using an additive method, the number of circumferential sections can be tuned to suit its desired application.
In some embodiments, the at least two circumferential sections each independently has a length along the longitudinal axis of about 2 mm to about 15 mm. This is also referred to as the peak to peak amplitude when the circumferential section has a wave-like structure. In other embodiments, the length is about 2 mm to about 14 mm, about 2 mm to about 13 mm, about 2 mm to about 12 mm, about 2 mm to about 11 mm, about 2 mm to about 10 mm, about 2 mm to about 9 mm, about 2 mm to about 8 mm, about 2 mm to about 7 mm, about 2 mm to about 6 mm, or about 2 mm to about 5 mm.
The stent can be formed with wires of circumferential units of varying cross sectional diameters. For example, a cross sectional diameter of the circumferential units in a circumferential section at an end portion of the stent can be thinner than a cross sectional diameter of the circumferential units in a circumferential section at a middle portion of the stent. Advantageously, this allows for stent to collapse into a more compact state such that it is easier to insert into a blood vessel.
In some embodiments, the at least two circumferential sections are two terminal circumferential sections. In other embodiments, the at least two circumferential sections comprises two terminal circumferential sections. The terminal circumferential sections can have a different morphology compared to the non-terminal circumferential sections. The two terminal circumferential sections can have has a wave-like structure with peaks which are not connected to the flex units, wherein the non-connected peaks are around. Advantageously, this allows for a reduction of sharp edges so as to reduce damage to the blood vessel.
In some embodiments, the stent comprises 1 to 9 flex sections. The stent can comprises 2 to 9 flex sections, 3 to 9 flex sections, 4 to 9 flex sections, 5 to 9 flex sections, or 6 to 9 flex sections. The stent can comprise 2, 3, 4, 5, 6, 7, 8, or 9 flex sections. The number of flex sections depends on the desired length of the stent, which depends on its application in the body. As the presently disclosed stent is fabricated using an additive method, the number of flex sections can be tuned to suit its desired application.
In some embodiments, the flex sections each independently comprise a plurality of flex units. The number of flex units in a flex section can be different from the number of flex units in another flex section. For example, the flex sections can each have a different number of flex units. In this regard, a flex section can have 8 flex units while a neighbouring flex section can have 7 flex units.
In some embodiments, the circumferential sections each independently comprise a plurality of circumferential units. The number of circumferential units in a circumferential section can be different from the number of circumferential units in another circumferential section. For example, the circumferential sections can each have a different number of circumferential units. In this regard, a circumferential section can have 10 flex units while a neighbouring circumferential section can have 9 flex units.
In some embodiments, the cross sectional diameter of wires of the plurality of flex units at an end portion of the stent is smaller than a cross sectional diameter of wires of the plurality of flex units at a middle portion of the stent. Advantageously, this allows for stent to collapse into a more compact state such that it is easier to insert into a blood vessel.
Strut geometry plays an important role in determining blood recirculation zones and shear rates. The inventors have found that a non-streamlined strut deployed at the arterial surface in contact with flowing blood, regardless of the height of the strut, promotes creation of recirculation zones, low shear rates as well as prolonged particle residence time. Based on computational studies, a significant recirculation region is present both downstream and upstream of the non-streamlined rectangular geometry, and the regions increases with increasing height. Whereas for the streamlined circular arc geometry, recirculation zone is observed only for the arc with largest height while the other arcs do not demonstrate any flow separation. Taken altogether, a streamlined geometry with smaller slopes of cross section morphology (less than about 90°) will be more favourable for endothelialization which decreases the rate of both restenosis and LST.
Conventional stents have wires with a rectangular cross section. Additionally, one limitation of 3D printing is the resultant surface quality of the printed product. For 3D printed products such as stents where the hatch distance is the diameter of the wire for diameter of less than 0.2 mm, and for parts which require low surface roughness, post processing such as electropolishing (EP) is required. In most cases, the surface finish after EP will still not be able to meet the low roughness requirements if the particle adherence to the structure is severe. The inventors have found that the 3D printed stent can be further improved when the wires forming the stent have a partially curved cross sectional shape. Particularly advantageously, besides improving the blood flow, these cross sectional shapes (for example arc, elliptical or aerofoil) also reduce particles adherence to the structure (or less balling). This reduces the risk of tearing or rupturing of the blood vessel during insertion.
As the partially curved cross sectional shape is of an anisotropic shape, it can be characterised by a cross sectional thickness and a cross sectional width. In some embodiments, the cross sectional thickness is about 0.01 mm to 0.5 mm, about 0.01 mm to 0.4 mm, about 0.01 mm to 0.3 mm, about 0.01 mm to 0.2 mm, about 0.1 mm to 0.5 mm, or about 0.1 mm to 0.4 mm. In other embodiments, the cross sectional width is about 0.1 mm to 0.5 mm, about 0.1 mm to 0.4 mm, or about 0.2 mm to 0.4 mm.
The stent can have a diameter of about 4 mm to about 12 mm. In other embodiments, the diameter is about 5 mm to about 12 mm, about 6 mm to about 12 mm, about 7 mm to about 12 mm, about 8 mm to about 12 mm, or about 9 mm to about 12 mm.
The stent can have a length of about 7 mm to about 50 mm. In other embodiments, the length is about 7 mm to about 40 mm, about 7 mm to about 30 mm, or about 7 mm to about 20 mm. The length of stent can be extended depending on a patient-specific scale and requirement.
The stent can have a length to diameter aspect ratio of about 15:1 to about 30:1. In some embodiments, the aspect ratio is about 16:1 to about 30:1, about 17:1 to about about 18:1 to about 30:1, about 19:1 to about 30:1, about 20:1 to about 30:1, about 22:1 to about 30:1, about 24:1 to about 30:1, about 26:1 to about 30:1, or about 28:1 to about 30:1.
Current stents in the market are limited to straight stents. When a straight stent is placed in the curved blood vessel, a mechanical stress will be induced on the vessel wall. Overtime, this stress will cause injury to the vessel wall leading to restenosis. Curved stents with wires of less than 1 cm are seldom used due to the difficulties in fabricating and in delivering the curved stents to the site of repair in the vessel. It was found that by manipulation of the shape memory property as disclosed herein, the stent can have a straight profile when in a crimped state but adopt a curvature in an expanded state upon deployment in the vessel. In the crimped state, the straight profile facilitates the transfer of the stent in a catheter and positioning of the stent at the target vessel site. In the expanded state, the stent conforms to the curvature of the vessel thus alleviating stress on the vessel and on the stent. Accordingly, the stent, when in an expanded state, can have a curvature between its terminal ends (stent edges) and along its longitudinal dimension (
The curved stent is able to be crimped (and thus straightened) for insertion into the catheter at room temperature and expanded to return to its original curved form at body temperature (when inside the vessel). This can be achieved without heat treatment post-fabrication of the stent by controlling the properties of the shape memory alloy.
To impart a curvature to a stent, the units forming a section can have different longitudinal lengths. In some embodiments, the lengths of respective flex units in a flex section are different from each other in the longitudinal dimension. For example, the length of a flex unit in a flex section can be different from its neighbouring flex unit. This can provide a curvature to the stent when expanded.
In some embodiments, the lengths of respective circumferential units in a circumferential section are different from each other in the longitudinal dimension. For example, the length of a circumferential unit in a circumferential section can be different from its neighbouring circumferential unit. This provides a curvature to the stent when expanded.
In some embodiments, the flex units in a flex section have a length in the longitudinal dimension which is different from the flex units in an adjacent flex section. In other embodiments, a flex section has a longitudinal length which is different from the longitudinal length of an adjacent flex section.
In some embodiments, the circumferential units in a circumferential section have a length in the longitudinal dimension which is different from the circumferential units in an adjacent circumferential section. In other embodiments, a circumferential section has a longitudinal length which is different from the longitudinal length of an adjacent circumferential section.
The transverse breadth of the flex units can also be varied to create suitable driven force, displacement and geometrical curvatures for stents. In some embodiments, the transverse breadth of respective flex units in a flex section are different from each other. In other embodiments, the transverse breadth of flex units in a flex section are different from those of flex units in an adjacent flex section. For example, the transverse breadth can be about 1 mm to about 7 mm, about 1 mm to about 6 mm, about 1 mm to about 5 mm, about 2 mm to about 5 mm, or about 3 mm to about 5 mm.
In some embodiments, the stent has a curvature between its terminal ends of about 1° to about 160°. In a stent with a single curve, the curvature refers to an angle between two convergent lines extending from an edge of the stent to an opposite edge of the stent. In a stent with multiple curves, the curvature can be measured between an edge to an inflexion point along the length of the stent or between two adjacent inflexion points. In other embodiments, the curvature is about 1° to about 150°, about 1° to about 140°, about 1° to about 130°, about 1° to about 120°, about 1° to about 110°, about 1° to about 100°, about 1° to about 90°, about 1° to about 80°, about 1° to about 70°, about 1° to about 60°, about 1° to about 50°, about 1° to about 45°, about 1° to about 40°, about 1° to about 35°, about 1° to about 30°, about 1° to about 25°, about 1° to about 20°, about 1° to about 15°, or about 1° to about °.
In some embodiments, the stent is characterised by a radius of curvature of about 1 mm to about 100 mm. In other embodiments, the radius of curvature is about 1 mm to about 90 mm, about 1 mm to about 80 mm, about 5 mm to about 80 mm, about 10 mm to about 80 mm, about 20 mm to about 80 mm, about 30 mm to about 80 mm, about 40 mm to about 80 mm, about 50 mm to about 80 mm, or about 60 mm to about 80 mm.
In some embodiments, the stent is characterised by a radius of curvature of about 1 cm to about 200 cm. In other embodiments, the radius of curvature is about 1 cm to about 180 cm, about 1 cm to about 160 cm, about 1 cm to about 150 cm, about 1 cm to about 140 cm, about 1 cm to about 130 cm, about 1 cm to about 120 cm, about 1 cm to about 110 cm, about 1 cm to about 100 cm, about 1 cm to about 90 cm, about 1 cm to about 80 cm, about 1 cm to about 70 cm, about 1 cm to about 60 cm, about 1 cm to about cm, about 1 cm to about 40 cm, about 1 cm to about 30 cm, about 1 cm to about 20 cm, or about 1 cm to about 10 cm.
As used herein, “radius of curvature” means the radius of a circle that touches a curve at a given point and has the same tangent and curvature at that point. The radius of curvature of a stent may refer to the radius of curvature of either side of the stent, or the radius of curvature of the longitudinal axis of the stent graft.
The stent can have a curvature at more than one location along its length.
In some embodiments, the stent comprises an elemental composition of:
In some embodiments, the stent comprises an elemental composition of:
The elemental composition of Nitinol stent was found to fall within medical grade Nitinol, as defined by ASTM F2063.
The present invention also provides a method of fabricating a 3d printed stent.
Accordingly, the present invention provides a method of 3D printing a stent, comprising:
In some embodiments, the method of 3D printing a stent, comprises:
Advantageously, the method allows for greater flexibility in fabrication as process parameters can be used to focus on manipulating the Nitinol properties to achieve the desired outcome. Further advantageously, by 3D printing the stent, the stent can be customised to adapt to a patient's blood vessel conditions, by for example varying radical dimensions.
The inventors have found that to ensure the expanded stent has a definite cylindrical shape, it is particularly advantageous to fabricate the stent in its expanded form before shape-setting it through heat treatment.
In some embodiments, the stent is printed in its expanded state.
While the present invention is illustrated using SLM, other AM approaches such as laser engineered net shaping (LENS) E-beam melting (EBM), direct metal laser sintering (DMLS) and selective laser sintering (SLS) can also be used.
As mentioned, the stent is 3D printed from Nitinol, a shape memory alloy. Nitinol is an intermetallic compound with approximately equiatomic nickel and titanium, exhibits the unique properties of shape-memory and superelasticity which were originated from reversible phase change from an austenitic to a martensitic microstructure when subjected by temperature or external stress. At lower temperatures, nitinol have a readily deformable crystalline arrangement termed martensite. This structure can be achieved by cooling nitinol below the martensite start and finish temperatures, Ms and Mf respectively, which restructures the material into the low-temperature stable martensitic phase. By reheating the nitinol through its austenite start and finish temperatures, As and Af respectively, the alloy passes through a characteristic transformation temperature range (TTR), causing the realignment of atomic planes that has occurred to be reversed. The crystal structure alters to a rigid and ordered cubic-like configuration known as austenite. This reversible process describes the shape-memory effect of nitinol.
Current stents in the market rely only on its superelastic properties for their deployment in a subject. The stent when subjected to large deformations are able to immediately return to its undeformed shape upon removal of the external load. It is this function that allows the stent to be compressed for insertion (via a delivery tube) and which subsequently ‘open’ when exiting the delivery tube for supporting a blood vessel. The shape memory effect of Nitinol is not used as there is an industry scalable method to control the Af temperature within a suitable range.
The inventors have found a method that allows deployment of stents of the present invention (via shape memory effect) to be achieved by heating the nitinol stent to a suitable temperature, such as a body temperature. To this end, the nitinol stent can be cooled and crimped (in the martensitic phase). The crimped stent maintains its collapsed state and can subsequently actuate or expand to its expanded state (austenite phase) when, for example, placed at body temperature (
In particular, the inventors have found that specific printing parameters can be used to vary the austenite and martensite temperatures. By doing so, the temperatures at which superelasticity takes effect can be controlled such that it is suitable for use in the human body. Additionally, by controlling the austenite and martensite temperatures, the shape memory properties of Nitinol can be advantageously utilised in a way to allow for easy insertion of a stent into a vessel.
Without wanting to be bound by theory, the inventors postulated that additive manufacturing processes can be used to fabricate stents. Additive manufacturing, known 3D printing, is defined as “a process of joining materials to make objects from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing methodologies” (ASTM international). Powder bed fusion (PBF) method can be used to melt or fuse powders together in a layer-by-layer approach, and the printing process was named according to the energy types, i.e., E-beam melting (EBM), selective laser melting (SLM), direct metal laser sintering (DMLS) and selective laser sintering (SLS). SLM is a laser-powered powder bed fusion process that can produce highly dense metallic parts with delicate geometrical features. Metallurgical bonding of originally loose powders can be achieved by laser melting followed by rapid solidification. This invention has introduced the optimized printing parameter regimes in a selective laser melting system, to achieve superelasticity, shape memory effect, and surface finishing using process parameters under different regimes.
To optimize the printing process parameters for stents, wires having aspect ratio exceed 20:1 comparing the height to the diameter, were printed at angles ranged from 90°, 60°, 45° and 30°, respectively. Functional material characterizations were conducted, wherein the porosity of wires were examined by observing the microstructurel features and high resolution X-ray computed tomography, the elemental distribution of nickel and titanium was analyzed using energy dispersive X-ray detector (EDX) attached to scanning electron microscope (SEM), the phase formed was characterized using differential scanning calorimetry (DSC), and the mechanical performances were characterized using mechanical tensile tests.
The volumetric energy density can be defined as:
E
v
=P/(V×h×t)
Where, Ev is energy density (J/mm3), P is the laser power (W), v is the scan speed (mm/s), and h is scan spacing or hatch distance (mm), and t is layer thickness (mm).
Without wanting to be bound by theory, it was found that the energy level, which is commonly used in the art, is not sensitive enough to predict the printing quality. This is especially so for Nitinol, as the shape memory and superelasticity properties are highly sensitive to powder composition and printing parameters. For example, it was found that even if using the same energy level, samples show different mechanical properties when different combinations of parameters were applied. Additionally, it was found that keeping the term of (v×h×t) as constant while changing any of three individual parameters does not necessarily lead to fabrication of identical parts in terms of microstructure and mechanical responses. It is believed that to improve consistency of printing, the relationship between the printing energy levels, elemental composition, the transformation temperatures and mechanical performances must be understood. This is aided using microscope, DSC, EDX and mechanical analysis. As disclosed herein, the energy levels of printing nitinol wires were tested from 6.67 to 2333.33 J/mm3 by changing the laser power and scanning speed, with fixed hatch distance of 0.1 mm and layer thickness of 0.03 mm. It was found that with adjustment of parameters, two printing regimes can be obtained. To this end, by selectively changing the printing conditions, optimized combinations for printing stents of different mechanical properties can be fabricated for different applications.
Several factors can be modulated to 3D print stents from Nitinol.
To elucidate desirable printing parameters, the microstructural features of printed samples at different parameters were compared to observe the microstructural evolution path with respect to the SLM process parameters.
The three samples all showed full density (no porosity in the structure) (
In contrast,
Three struts were printed to examine the microstructures formed when printed at P100 v100, P200 v500, and P300 v1000, corresponding to the energy density of 333.3, 133.32, and 99.9 J/mm3. The Ms temperatures was at −1.24° C. for P100 v100, at 35.43° C. for P200 v500, and at 30.11° C. for P300 v1000. As shown in
These three trends suggested melt pool boundary evolutions. Firstly, at low laser power printing, decreasing the scanning speed to 150 mm/s or below has caused inter-layer over melt, thus formed columnar grains. However, this is not likely to induce significant elemental redistribution in the strut. As a result, the phase transformation temperatures of struts processed under this regime were close to that of the powder. Secondly, high laser power (320 W) tended to re-melt the previous layers and likely caused redistribution of nickel elements in the grain boundaries. Moreover, nickel evaporation may be induced at high laser power. This facilitated secondary phase formation further, thus degraded the superelasticity of struts. Lastly, the laser power was the dominant indicator of the microstructural evolution in thin nitinol struts printing, whereas the energy density was not.
The Ms temperature of the SLM processed samples was generally higher than the powder (Ms: −17.9° C.). When printed at a low scanning speed of 50 mm/s and 150 mm/s, the Ms temperature of samples showed the same trend with respect to the increasing laser power from 50 W to 350 W. Specifically, the Ms temperature of the samples was kept at about −7.4° C. to 1.2° C. at the low scanning speed and low laser power region (50 W and 125 W). In contrast, it reached about 24.3° C. to 67.3° C. when the laser power was increased from 200 W to 280 W, respectively. This observation indicated that the Ms temperature met a threshold of transition at low scanning speed printing when the laser power was increased from 125 W to 280 W, corresponding to the energy density varied from 833 J/mm3 to 1867 J/mm3, respectively.
Accordingly, in some embodiments, the stent can have a temperature at which it initially transits from martensite to austenite (As (stent)) and a temperature at which it initially transits from austenite to martensite (Ms (stent)).
Correspondingly, the Nitinol powder has a temperature at which it initially transits from martensite to austenite (As (Nitinol powder)) and a temperature at which it initially transits from austenite to martensite (Ms (Nitinol powder)).
In some embodiments, As (stent) is lower than As (Nitinol powder) and Ms (stent) is higher than Ms (Nitinol powder).
In some embodiments, As (Nitinol powder) is about −6° and Ms (Nitinol powder) is about −18° C. In other embodiments, As is about −5.6° C. In other embodiments, Ms is about −17.9° C.
At the intermediate scanning speed of 500 mm/s, the Ms temperature was at about 23.2° C. and 23.7° C. when the laser power was at 50 W and 150 W; thereafter the Ms temperature increased to 33.4° C. and 33.9° C. when the laser power reached 200 W and 280 W. It jumped to about 60.5° C. when the laser power was increased to 320 W and 350 W. Herein, the threshold of transition in the Ms temperature occurred at a scanning speed of 500 mm/s combined with laser power of 280 W to 320 W, corresponding to the energy density of 187 J/mm3 to 213 J/mm3.
At the high scanning speed of 1500 mm/s, the Ms temperature slowly increased from 3.2° C. to 27.2° C. when the laser power was increased from 125 W to 350 W. No significant trend of transition was observed at the high scanning speed printing.
X-ray Diffraction (XRD) spectra shows that the major peaks of austenitic phase (B2) and martensitic phase (B19′) were both shown when printed under P100 v100 and P200 v500. This was consistent with the phase transformation temperatures measured below RT. However, in the as printed P300 v1000 strut, a significant amount of peaks showing secondary phases, which were identified as the Ni4Ti, Ti2Ni, TiC, and NiCx, accordingly. By applying heat treatment on the stents, those secondary phases were disappeared from the XRD spectrum. The minor martensitic phase and secondary nickel-rich or titanium-rich phases could be attributed to the significant thermal instability in the high laser power combined with high scanning speed printing, and can be due to austenitic B2 phase dominated microstructures. Apparent noise and broadening at the base of the XRD peaks could have been a remnant from secondary phases with a low volume fraction.
The nitinol wt % of the SLM processed samples was further compared with the virgin nitinol powders. When printed at low laser power & low scanning speed (P100 v100), the energy density was 333.3 J/mm3. The nickel wt % varied from 54.64% to 55.63%. After heat treatment of newly printed samples, epitaxial columnar features showed similar nickel wt % with the surroundings, with nickel wt % varied from 54.66% to For the sample printed under P200 v500, the nickel wt % varied from 54.96% to 55.41% in the as printed state, and varied from 54.72% to 55.19% after heat treatment. The same XRD peaks in the as-printed and heat-treated struts further confirmed that the P200 v500 strut was dominated by the austenitic phase.
The strut printed at P300 v1000 represented the high laser power and high scanning speed combination, although the energy density was relatively low (100 J/mm3). The nickel wt % was at 53.74% to 54.77% in the as-printed state, whereas it has increased significantly after heat treatment, i.e., varied from 54.72% to 55.68%. The significant increase in the Ni wt % after heat treatment could originate from the dissolving of secondary Ni-rich phases. This was consistent with the XRD analysis of crystalline structures of samples. Apparently, as summarized in
The above indicates that besides the volumetric energy density, the combination of SLM process parameters (i.e., scanning speed and laser power) also influenced the nickel weight percentage remarkably, and the phase transformation behavior of SLM fabricated nitinol alloys.
The printing process window for superelastic nitinol was indicated using the Ms temperature as an indicator summarized in
In Region I (
In Region II (
Two aspects could be suggested from this observation. Firstly, the level of nickel evaporation has kept consistent when printed using the parameters in Region I and Region II, i.e., decreased by about 0.3% to 0.5%. Secondly, increasing the laser power and scanning speed has facilitated the formation of Ni-rich secondary phases in the SLM process, which has further increased the Ms temperature of the struts printed. Lastly, those secondary phases could be eliminated using heat treatment. Therefore, the Ms temperature could be brought down using appropriate heat treatment conditions, accordingly.
Region III (
Furthermore, samples were peeled off from the substrate during the powder removing process when printed using parameters in Region IV (
Based on the stent design as disclosed herein, stents was printed with parameters P100 v100. The printed stents were with strut radius varied from 0.2093-0.2342 mm, and the overprint ratio was ranged from 39%-56% (overprint ratio=(printed diameter−designed diameter)/designed diameter) (
The correlation between laser power and speed is also shown in
Under the regime of A1, low laser power combined with low scanning speed resulted high levels of energy density, however the local temperature may maintained below certain threshold which has favoured the Ti-rich phase formation. Furthermore, Ni-rich secondary phase may also be formed during this process due to the complex thermal history involved in different sections of wire. Similarly, under the regime of A2, intermediate power but high speed may result in low local temperature which favoured secondary phase formation.
Accordingly, the present invention provides a method of 3D printing a stent, comprising: performing selective laser melting on a Nitinol powder in order to form the stent, wherein selective laser melting is performed with:
In other embodiments, the selective laser melting is performed with a laser power of about 50 W to about 200 W, and a scanning speed of about 50 mm/s to about 150 mm/s. In other embodiments, the power is about 50 W to about 150 W, and a scanning speed of about 50 mm/s to about 900 mm/s, about 50 mm/s to about 800 mm/s, about mm/s to about 700 mm/s, about 50 mm/s to about 600 mm/s, about 50 mm/s to about 500 mm/s, about 50 mm/s to about 450 mm/s, about 50 mm/s to about 400 mm/s, about 50 mm/s to about 350 mm/s, about 50 mm/s to about 300 mm/s, about mm/s to about 250 mm/s, about 50 mm/s to about 200 mm/s, or about 50 mm/s to about 150 mm/s.
In some embodiments, the laser power is about 150 W to about 250 W, and the scanning speed is about 500 mm/s to about 2500 mm/s, about 500 mm/s to about 2000 mm/s, or about 500 mm/s to about 1500 mm/s.
In some embodiments, As (stent) is about −45° C. to about −25° C. In other embodiments, As (stent) is lower than As (Nitinol powder) but not below −45° C. In some embodiments, Ms (stent) is about −10° C. to about 10° C. These stents can, for example, be produced by Regime A1 of the method as disclosed herein.
In some embodiments, As (stent) is higher than As (Nitinol powder) and Ms (stent) is higher than Ms (Nitinol powder).
In some embodiments, As (stent) is about −30° C. to about 0° C. and Ms (stent) is about −10° C. to about 25° C. These stents can, for example, be produced by Regime A2 of the method as disclosed herein.
In some embodiments, the step of selective laser melting the Nitinol powder is such that the laser has a power of about 50 W to about 125 W, and a scanning speed is about mm/s to about 500 mm/s. This set of conditions falls within Regime A1 as disclosed above. In other embodiments, the power is about 50 W to about 120 W, about 50 W to about 110 W, or about 50 W to about 100 W. In other embodiments, the scanning speed is about 50 mm/s to about 450 mm/s, about 50 mm/s to about 400 mm/s, about 50 mm/s to about 350 mm/s, about 50 mm/s to about 300 mm/s, about 50 mm/s to about 250 mm/s, about 50 mm/s to about 200 mm/s, about 50 mm/s to about 150 mm/s, or about 50 mm/s to about 100 mm/s.
In some embodiments, the step of selective laser melting the Nitinol powder is such that the laser has a power of about 125 W to about 200 W, and a scanning speed of about 500 mm/s to about 1,500 mm/s. This set of conditions falls within Regime A2 as disclosed above. In other embodiments, the power is about 150 W to about 200 W, about 160 W to about 200 W, about 170 W to about 200 W, about 180 W to about 200 W, or about 190 W to about 200 W. In other embodiments, the scanning speed is about 600 mm/s to about 1500 mm/s, 700 mm/s to about 1500 mm/s, 800 mm/s to about 1500 mm/s, 900 mm/s to about 1500 mm/s, 1000 mm/s to about 1500 mm/s, 1100 mm/s to about 1500 mm/s, 1200 mm/s to about 1500 mm/s, 1300 mm/s to about 1500 mm/s, or 1400 mm/s to about 1500 mm/s.
Under regime B, all of the Ms and As transformation temperatures of printed wires were above the values of powders, when applying laser power at 50 W and 100 W combined with scanning speed of 500 mm/s, 1000 mm/s and 1500 mm/s, wherein the energy levels were varied from 11 to 67 J/mm3 (Regime B1-low power-high speed-low energy). Similar increased phase transformation temperature was observed when the laser power was increased to 150 W and 200 W while the scanning speed was decreased to 100 mm/s and 150 mm/s, resulted energy density of 133 to 1333 J/mm3 (Regime B2-high power-low speed-high energy), respectively. Furthermore, phase transformation temperatures were all above that of the powders when the laser energy was further increased to 280 W, 300 W, 320 W and 350 W, either under high speed of 500 mm/s to 1500 mm/s or under low speed of 50 mm/s to 150 mm/s, with an energy range from 62-233 J/mm3 for high speed and 622-2333 J/mm3 for low speed, respectively (Regime B3-high power-varies speed-varies energy, also correlating to Region II in
This could indicate a deficient of nickel in the printed wires. Firstly, regime B1 likely promote formation of nickel rich secondary phases such as Ni3Ti during SLM process, thus the nickel element was decreased in the intermetallic nitinol phase. Secondly, when under high energy B2 and high power B3 regime, nickel will likely be evaporated during SLM process, caused an increase in the phase transformation temperature. This behaviour is even clear when comparing the B3 regime under high powder but low speed and high speed. When printed under the same high laser power, lower scanning speed resulted higher transformation temperatures than that of the higher scanning speed, further supporting nickel evaporation when under high energy levels.
The above printing was performed on vertical thin structures of diameter 0.3 mm and 1 cm height for characterisation. It would be understood that printing of stent structures using these parameters could have slightly different results due to the angulation of the stent structures during printing.
The present invention provides a method of 3D printing a stent, comprising: performing selective laser melting on a Nitinol powder in order to form the stent, wherein selective laser melting is performed with:
In some embodiments, the selective laser melting is performed with a laser power of about 150 W to about 250 W, and a scanning speed of about 50 mm/s to about 450 mm/s, about 50 mm/s to about 400 mm/s, about 50 mm/s to about 350 mm/s, about mm/s to about 300 mm/s, about 50 mm/s to about 250 mm/s, about 50 mm/s to about 200 mm/s, or about 50 mm/s to about 150 mm/s.
In some embodiments, the selective laser melting is performed with a laser power of about 250 W to about 350 W, and a scanning speed of about 500 mm/s to about 2500 mm/s, about 500 mm/s to about 2000 mm/s, or about 500 mm/s to about 1500 mm/s.
In some embodiments, As (stent) is about 10° C. to about 25° C. and Ms (stent) is about C. to about 100° C. These stents can, for example, be produced by Regime B1 of the method as disclosed herein.
In some embodiments, As (stent) is about 10° C. to about 75° C. and Ms (stent) is about 60° C. to about 100° C. These stents can, for example, be produced by Regime B2 of the method as disclosed herein.
In some embodiments, As (stent) is about 10° C. to about 80° C. and Ms (stent) is about ° C. to about 70° C. These stents can, for example, be produced by Regime B3 of the method as disclosed herein.
In some embodiments, As (stent) is about 20° C. to about 75° C. and Ms (stent) is about 25° C. to about 100° C. These stents can, for example, be produced by Regime B of the method as disclosed herein.
In some embodiments, As (stent) is about −45° C. to about 80° C. and Ms (stent) is about −18° C. to about 100° C.
Taken altogether, the presently disclosed nitinol stent has a self-expanding deployment system, high flexibility and conformability due to a hybrid closed-cell design, thinner and rounder struts as well as adequate radial strength.
In some embodiments, the step of selective laser melting the Nitinol powder is such that the laser has a power of about 50 W to about 100 W, and a scanning speed of about 500 mm/s to about 1,500 mm/s. This set of conditions falls within Regime B1 as disclosed above. In other embodiments, the power is about 50 W to about 90 W, about 50 W to about 80 W, about 50 W to about 70 W, or about 50 W to about 60 W. In other embodiments, the scanning speed is about 600 mm/s to about 1500 mm/s, about 700 mm/s to about 1500 mm/s, about 800 mm/s to about 1500 mm/s, about 900 mm/s to about 1500 mm/s, about 1000 mm/s to about 1500 mm/s, about 1100 mm/s to about 1500 mm/s, or about 1200 mm/s to about 1500 mm/s.
In some embodiments, the step of selective laser melting the Nitinol powder is such that the laser has a power of about 150 W to about 200 W, and a scanning speed of about 50 mm/s to about 500 mm/s. This set of conditions falls within Regime B2 as disclosed above. In other embodiments, the power is about 160 W to about 200 W, about 170 W to about 200 W, or about 180 W to about 200 W. In other embodiments, the scanning speed is about 50 mm/s to about 450 mm/s, about 50 mm/s to about 400 mm/s, about 50 mm/s to about 350 mm/s, about 50 mm/s to about 300 mm/s, about 50 mm/s to about 250 mm/s, about 50 mm/s to about 200 mm/s, about 50 mm/s to about 150 mm/s, or about 50 mm/s to about 100 mm/s.
In some embodiments, the step of selective laser melting the Nitinol powder is such that the laser has a power of about 200 W to about 350 W, and a scanning speed of about 50 mm/s to about 1,500 mm/s. This set of conditions falls within Regime B3 as disclosed above. In other embodiments, the power is about 210 W to about 350 W, about 220 W to about 350 W, about 230 W to about 350 W, about 240 W to about 350 W, about 250 W to about 350 W, about 260 W to about 350 W, about 270 W to about 350 W, about 280 W to about 350 W, about 290 W to about 350 W, about 300 W to about 350 W, about 310 W to about 350 W, about 320 W to about 350 W, or about 330 W to about 350 W. In other embodiments, the scanning speed is about 60 mm/s to about 1500 mm/s, about mm/s to about 1500 mm/s, about 80 mm/s to about 1500 mm/s, about 90 mm/s to about 1500 mm/s, about 100 mm/s to about 1500 mm/s, about 200 mm/s to about 1500 mm/s, about 300 mm/s to about 1500 mm/s, about 400 mm/s to about 1500 mm/s, about 500 mm/s to about 1500 mm/s, about 600 mm/s to about 1500 mm/s, about 700 mm/s to about 1500 mm/s, about 800 mm/s to about 1500 mm/s, about 900 mm/s to about 1500 mm/s, about 1000 mm/s to about 1500 mm/s, about 1100 mm/s to about 1500 mm/s, about 1200 mm/s to about 1500 mm/s, or about 1300 mm/s to about 1500 mm/s.
Regimes A and B provide an indication of nickel-rich or titanium-rich precipitation phase formation conditions when varying the scanning speed and laser powers in SLM process. Advantageously, based on the above, after receiving the nitinol powders from a vendor, nitinol devices manufacturer can decide the regimes of parameters to refine based on the functional requirements of products, i.e. refine under Regime A to obtain products with lower phase transformation temperatures than the raw powders, while refine under Regime B to obtain products with higher phase transformation temperature than the raw powders.
In some embodiments, the selective laser melting is performed with a hatch distance of about 0.1 mm to about 0.5 mm. In other embodiments, the hatch distance is about 0.1 mm to about 0.4 mm, or about 0.1 mm to about 0.3 mm. In some embodiments, the selective laser melting is performed with a hatch distance of about 0.1 mm.
In some embodiments, the selective laser melting is performed with a layer thickness of about 0.01 mm to about 1 mm. In other embodiments, the layer thickness is about 0.01 mm to about 0.9 mm, about 0.01 mm to about 0.8 mm, about 0.01 mm to about 0.7 mm, about 0.01 mm to about 0.6 mm, about 0.01 mm to about 0.5 mm, about 0.01 mm to about 0.4 mm, about 0.01 mm to about 0.3 mm, about 0.01 mm to about 0.2 mm, about 0.01 mm to about 0.1 mm. In some embodiments, the selective laser melting is performed with a layer thickness of about 0.03 mm.
In some embodiments, a stent with a wire diameter of less than about 1 cm is 3D printed. In other embodiments, stent with a wire diameter of about 0.1 mm to about 0.9 mm is 3D printed, or about 0.1 mm to about 0.8 mm, about 0.1 mm to about 0.7 mm, about 0.1 mm to about 0.6 mm, about 0.1 mm to about 0.5 mm, about 0.1 mm to about 0.4 mm, about 0.1 mm to about 0.3 mm, or about 0.1 mm to about 0.2 mm.
For example, a stent with a wire diameter of about 0.1 mm can be printed using the following parameters:
For example, a stent with a wire diameter of about 0.2 mm to about 0.5 mm can be printed using the following parameters:
In some embodiments, the 3D printed stent is characterised by a austenite finish temperature (Af) of about 25° C. to about 50° C. In other embodiments, the Af temperature is about 30° C. to about 50° C., about 30° C. to about 45° C., about 30° C. to about 40° C., or about 35° C. to about 40° C. In other embodiments, the Af temperature is about 37° C.
In a preferred embodiment, the parameters are selected from:
The above parameters are suitable for stents with wire diameter of about 0.1 mm.
In a preferred embodiment, the parameters are selected from:
The above parameters are suitable for stents with wire diameter of about 0.2 mm.
In some embodiments, the method further comprises providing a template of the stent as disclosed herein.
The present invention also provides a 3D printed stent printed using the method as disclosed herein, the stent comprising Nitinol having a nickel content of about 54 wt % to about 57 wt % of the composition and a titanium content of about 43 wt % to about 46 wt % of the composition;
In some embodiments, the stent has a martensite to austenite transition (As) temperature of about −45° C. to about 0° C. and a austenite to martensite transition (Ms) temperature of about −10° C. to about 25° C.
In some embodiments, when the selective laser melting is performed using the above conditions, the stent is characterised by columnar grains due to inter-layer over melt.
In some embodiments, the stent has a martensite to austenite transition (As) temperature of about 10° C. to about 80° C. and a austenite to martensite transition (Ms) temperature of about 20° C. to about 100° C.
In some embodiments, when the selective laser melting is performed using the above conditions, the stent is characterised by fully merged layer boundaries due to re-melt of an underlying layer.
In some embodiments, the 3D printed stent is characterised by a austenite finish temperature (Af) of about 25° C. to about 50° C. In other embodiments, the Af temperature is about 30° C. to about 50° C., about 30° C. to about 45° C., about 30° C. to about 40° C., or about 35° C. to about 40° C. In other embodiments, the Af temperature is about 37° C.
In some embodiments, the nickel is about 54 wt % to about 56.9 wt %, about 54 wt % to about 56.8 wt %, about 54 wt % to about 56.7 wt %, about 54 wt % to about 56.6 wt %, about 54 wt % to about 56.5 wt %, about 54 wt % to about 56.4 wt %, about 54 wt % to about 56.3 wt %, about 54 wt % to about 56.2 wt %, about 54 wt % to about 56.1 wt %, about 54 wt % to about 56 wt %, about 54.1 wt % to about 56 wt %, about 54.2 wt % to about 56 wt %, about 54.3 wt % to about 56 wt %, about 54.4 wt % to about 56 wt %, about 54.5 wt % to about 56 wt %, about 54.6 wt % to about 56 wt %, about 54.7 wt % to about 56 wt %, about 54.8 wt % to about 56 wt %, about 54.9 wt % to about 56 wt %, about 55 wt % to about 56 wt %, about 55 wt % to about 55.9 wt %, about 55 wt % to about 55.8 wt %, about 55 wt % to about 55.7 wt %, about 55 wt % to about 55.6 wt %, or about 55 wt % to about 55.5 wt % of the composition. In other embodiments, nickel is about 54.5 wt % to about 55.8 wt % the composition. In some embodiments, the nickel is about 55.2 wt % the composition.
In some embodiments, the titanium is about 43 wt % to about 45.9 wt %, about 43 wt % to about 45.8 wt %, about 43 wt % to about 45.7 wt %, about 43 wt % to about 45.6 wt %, about 43 wt % to about 45.5 wt %, about 43 wt % to about 45.4 wt %, about 43 wt % to about 45.3 wt %, about 43 wt % to about 45.2 wt %, about 43 wt % to about 45.1 wt %, about 43 wt % to about 45 wt %, about 43.1 wt % to about 45 wt %, about 43.2 wt % to about 45 wt %, about 43.3 wt % to about 45 wt %, about 43.4 wt % to about 45 wt %, about 43.5 wt % to about 45 wt %, about 43.6 wt % to about 45 wt %, about 43.7 wt % to about 45 wt %, about 43.8 wt % to about 45 wt %, about 43.9 wt % to about 45 wt %, about 44 wt % to about 45 wt %, about 44.1 wt % to about 45 wt %, about 44.2 wt % to about 45 wt %, about 44.3 wt % to about 45 wt %, about 44.4 wt % to about 45 wt %, or about 44.5 wt % to about 45 wt % of the composition. In other embodiments, the titanium is about 44.5 wt % to about 45.2 wt % of the composition. In other embodiments, the titanium is about 44.8 wt % of the composition.
In some embodiments, wires of the 3D printed stent have a partially flat cross sectional shape. In some embodiments, wires of the 3D printed stent have an elliptical, tear drop, partially flattened tear drop or circular cross section shape.
In some embodiments, the 3D printed stent has a curvature along its longitudinal dimension when in the expanded state. In some embodiments, the 3D printed stent has a curvature of about 1° to about 160°. In some embodiments, the 3D printed stent has a radius of curvature of about 1 mm to about 200 cm.
The 3D printed stent can have rough surfaces. To further improve the surface morphology, a post processing step can be added.
In some embodiments, the method further comprises a step of heat treating the stent.
In some embodiments, the heat treating step comprises heating the stent from about 200° C. to about 800° C. In other embodiments, the heating is from about 200° C. to about 750° C., about 200° C. to about 700° C., about 200° C. to about 650° C., about 200° C. to about 600° C., about 250° C. to about 600° C., about 300° C. to about 600° C., about 350° C. to about 600° C., about 400° C. to about 600° C., or about 450° C. to about 600° C. The heating step can be performed for about 10 min, about 20 min, about 30 min, about 40 min, about 60 min, or for more than 60 min.
In some embodiments, the method further comprises a step of heat treating the stent when the stent is printed using condition ii, iii or iv. In other embodiments, the method further comprises a step of heat treating the stent when the stent is printed outside condition i.
Based on the DSC results, 24 wire samples with Ms and Af closer or below the room temperature was selected for mechanical tests, attempting to examine the superelasticity of those wires printed. Wire samples having diameters less than 1.0 mm and length of 15 mm, which were not suitable for the application of extensometer in the tensile tests using Instron equipment. A camera was used with digital image correlation and TEMA software to capture the accurate deformation strain with respect to load. The maximum stress (ultimate tensile stress) and fracture strain was summarized in the table below.
As shown in Table 6, under Regime A1-low power-low speed-high energy, the fracture strain has reached to 10%, indicating good ductility of the wire samples fabricated from SLM. However the fracture stress was low compared to other samples, likely due to high internal porosity since powders were not fully melted under Regime A1. Under Regime
B, the overall fracture strain was lower than that of regime A, however several samples under Regime B2-high power-low speed-high energy and Regime B3-high power-varies speed-varies energy showed acceptable fracture strain of about 8%-9%, most importantly high fracture stress was obtained under this regime.
It was found that the structural geometry can be particularly advantageous for improving surface quality of fine Nitinol structures without changing process parameter.
Further, it was also found that the printing direction of the stent can be particularly advantageous for improving surface quality of fine Nitinol structures. When thin structures of diameter 0.3 mm were built at various inclination angles at 30°, 35°, 40°, 45° 50°, 55°, it was found that particles attachment was reduced in comparison to vertically printed structures. At the upskin region, at 30° inclination, it could be clearly observed the upskin section of the structure is free from particles. As the inclination angle increases, more particles attachment occurs at the upskin section but this increase was within a standard deviation.
In some embodiments, the method of 3D printing a stent comprises printing the stent such that the longitudinal dimension is at an angle to a horizontal plane of about 30° to about 60°. This can for example be done by providing the template of the stent with the longitudinal dimension at an angle, such that when read by the 3D printing machine, the stent is printable at an inclination angle to a horizontal plane of about 30° to about 60°. This advantageously reduces or eliminates balling effect on the stent on the upskin section.
It was found that a partially curved cross sectional morphology such as an arc provides for about a ten times reduction of downskin surface area roughness compared to a circular cross section.
The sharp edge of the partially curved cross-section (such as an arc-shape design) reduces the contact surfaces between the particles and the structure at the downskin section. With a reduction of contact surfaces between particles and structure, adherence of partially melted particles to the structure downskin will be weaker thus effort to remove the particles during post processing will be reduced.
Other partially curved geometry can also reduce particle bound to the stent structures. For example, arc, elliptical, tear drop or partially flattened tear drop (such as aerofoil) cross section shape can be used to further improve the surface finishing of the stent.
Reducing the structure thickness to the desired dimension can be achieved by surface treatment methods such as electropolishing. With the reduction of balling effect, surface finish after electropolishing can be enhanced.
The present invention also provides a method of 3D printing a stent, comprising:
The present invention also provides a stent delivery device, comprising:
As the crimped stent has shape memory properties, the stent is revertible back to its original uncrimped state when at least exposed to a temperature above its As temperature.
The tube has a lumen for containing the crimped stent. The tube can be a catheter. In this regard, the tube can be a flexible tube of a suitable length and width.
The stent disposed within the tube is in a crimped state. In this regard, the stent is pressed or pinched into a small, compressed state. The stent holds itself in this state when the temperature is less than Ms temperature. The crimped stent has a smaller dimension compared to the lumen of the tube, and is thus slidable within the lumen of the tube. When the crimped stent is exposed to a temperature above As temperature, it reverts back to its original uncrimped state for deployment at the target site in the channel or vessel.
In some embodiments, the stent in its original uncrimped state is adapted to revert back to its original configuration after release of an external force and when exposed to a temperature of about 25° C. to about 50° C. This relies on the superelastic property of Nitinol processed according to the methods as disclosed herein.
In some embodiments, the stent delivery device further comprises ejecting means for ejecting the crimped stent out from the tube. The ejecting means can be a rod or wire insertable at one end of the tube for sliding the crimped stent out from the other end. The rod can be a flexible rod.
The present invention also provides a method of delivering a stent in a stent delivery device into a channel, the stent delivery device comprising:
To determine the flexibility and suitability of flex unit for 3D printing, different flex units were fabricated for comparison. These designs were compared with regard to the curvature and height of flex segments (FS), and how they fare against a cantilever beam that takes up the same width. With reference to
Simulation studies of these models are performed with Autodesk Fusion 360, where each model is constrained on one end, with an arbitrary downward force of 1N placed on the other. Material selected is steel (Young's Modulus: 210000 MPa, Poisson's Ratio: 0.3, Yield Strength:207 MPa, Ultimate Tensile Strength: 345 MPa).
Simulations of designs 1 to 7 were carried out using Autodesk Fusion 360 to compare how each design performed when placed under conditions similar to the mechanical tests. The tests performed in the simulations are torsion, bending, axial tension and compression, with an equal arbitrary load of 10N for testing each different design for relative comparison. The material used for the simulations is steel, which is the same as what was used in the simulation of the flex segments. A base plate is added to the end of each model to facilitate the simulation, where the force will be placed on the base plate to simulate loading conditions.
Through the simulations, areas of the stent designs that will fail due to high stress and strain concentrations were identified. The results also revealed how different parameters such as number of circumferential units and strut unit thickness affects the performance of the stent, as well as how the different designs will deform when placed under different loading conditions.
Torsion: To simulate a torsional force on the various designs, the models are fully constrained on one end and a torsional force of 10N is applied onto the baseplate connected to the other end of the model. The baseplate is also constrained in the axis along the print direction (axis Y).
Bending: To simulate bending on the various designs, the models are fully constrained on one end and a vertical downward force of 10N is applied onto the baseplate connected to the other end of the model. The baseplate is also constrained in the axes that are not along the direction of the downward force (axes Y and Z).
Compression: To simulate compression on the various designs, the models are fully constrained on one end and a horizontal force of 10N along axis Y is applied onto the baseplate (towards the stent) connected to the other end of the model. The baseplate is also constrained in the axes that are not along the direction of the horizontal force (axes X and Z).
Tension: To simulate tension on the various designs, the models are fully constrained on one end and a horizontal force of 10N along axis Y is applied onto the baseplate (away from the stent) connected to the other end of the model. The baseplate is also constrained in the axes that are not along the direction of the horizontal force (axes X and Z).
Differential scanning calorimeter (DSC) tests were conducted at heating/cooling rate of on the as-printed samples. The test temperature was ranged from −80° C. to 100° C. Several samples having peaks below −80° C. and above 100° C. were ignored. For the powders, the Ms temperature was at −18° C. while the As temperature was at −6° C.
We have purchased this composition to achieve superelasticity of wire after SLM printing.
The stent can be fabricated using a metal printer such as EOS M 290 3D printer, which has a fibre laser focus diameter of 100 μm, estimated laser affected area of 150-200 μm and estimated total affected area of 230-280 μm, as well as a building volume of 250×250×535 mm. The commercial Ni (55.4 wt %)-Ti powder (size rage 15-45 μm) was provided by Advanced Powders and Coating (GE Additive, Canada).
The laser power was varied from 50 W to 350 W. Meanwhile, the scanning speed was varied from 50 mm/s to 1,500 mm/s, contributed to an energy density ranged from 11.1 J/mm3 to 2,333.3 J/mm3 (Table 7). Moreover, the default hatch distance was at 0.1 mm, layer thickness was at 0.03 mm, and the oxygen level was controlled below 100 ppm.
In general, Nitinol struts having a diameter of 0.3 mm and height of 15 mm were designed and then printed vertically, reached an aspect ratio of 50:1. Thereafter, nitinol stents with a strut diameter of 0.3 mm were printed using the selected parameters.
Table 7 shows exemplary printing parameters that are suitable for printing stents with a strut (wire) diameter of about 0.1 mm to about 0.4 mm.
The 3D printed samples were mounted and ground by SiC 1200 grit sandpaper, and further polished with a grinding-polishing machine. The polished samples were etched for 180 seconds with ‘H2O (82.7%), HNO3 (14.1%), and HF (3.2%) solution’. The samples were cleaned with ethanol and pure water, then dried with an air gun.
Nikon Microscope (Nikon, Japan, ECLIPSE LV150N, S/N 251814) microscope under a bright field was used to examine the exposed microstructure. An energy dispersive X-ray detector ‘X-act’ (Oxford Instruments plc, United Kingdom), attached to a JOEL JSM-6010 PLUS/LV SEM (The Japan Electron Optics Laboratory Company, Limited, Japan), was used to examine the elemental composition on the exposed plane. The phases, or crystal structures, presented in those selected samples were accessed using Micro XRD (D8 Discover, Bruker, United States). Measurements were conducted at room temperature, with step intervals of 0.2° in 2θ ranged from 20° to 100°.
XCT was used to examine the three-dimensional geometric features of SLM processed nitinol stents. The GE Nanotom M (General Electric Company, United States) was used to capture thousands of images and reconstruct them into 3D volume. In addition, the VG studio Max 3.0 (Volume Graphics GmbH, Germany) was used for surface determination and volume analysis.
Experiments were performed using a SLM printer Aconity3D MINI system with a 400 W Ytterbium Fiber Laser CW with a wavelength of 1070 nm. The laser spot is calibrated to the smallest achievable size of 55 μm. The chamber was purged with argon protection gas to keep the oxygen level below 100 ppm. A commercial Nitinol powder (NiTi) powder of 50-50 wt % (or about 55.4 wt % Ni) provided by the Advanced Powders and Coatings, GE Additive, was used as a raw material. The average powder size was about 15-53 μm.
Keyence VHX-6000 digital microscope was used to inspect the surface quality of the printed samples and measuring relative surface roughness profile of the polished samples. As printed samples for the metallographic tests were sliced, ground, polished, and then etched using a reagent (3 mL HCl, 1 mL HNO3, and 96 mL H2O) for 10 s. The microstructures were observed with an Olympus DM-4000M metallographic microscope (Wetzlar, Germany).
Printed specimens with different arc width underwent an electro-mechanical polishing process using the Dlyte machine. Particles attachment were absent with pit defects observed for all samples. Arc width of 0.1 mm has the least pit defect in the downskin section. The pit defects could be attributed to the partially fused particles attached to the downskin section. Pits were formed upon removed during electro-mechanical polishing.
It will be appreciated that many further modifications and permutations of various aspects of the described embodiments are possible. Accordingly, the described aspects are intended to embrace all such alterations, modifications, and variations that fall within the spirit and scope of the appended claims.
Throughout this specification and the claims which follow, unless the context requires otherwise, the word “comprise”, and variations such as “comprises” and “comprising”, will be understood to imply the inclusion of a stated integer or step or group of integers or steps but not the exclusion of any other integer or step or group of integers or steps.
Throughout this specification and the claims which follow, unless the context requires otherwise, the phrase “consisting essentially of”, and variations such as “consists essentially of” will be understood to indicate that the recited element(s) is/are essential i.e. necessary elements of the invention. The phrase allows for the presence of other non-recited elements which do not materially affect the characteristics of the invention but excludes additional unspecified elements which would affect the basic and novel characteristics of the method defined.
The reference in this specification to any prior publication (or information derived from it), or to any matter which is known, is not, and should not be taken as an acknowledgment or admission or any form of suggestion that that prior publication (or information derived from it) or known matter forms part of the common general knowledge in the field of endeavour to which this specification relates.
| Number | Date | Country | Kind |
|---|---|---|---|
| 10202007537P | Aug 2020 | SG | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/SG2021/050460 | 8/5/2021 | WO |
