WO2021181116A1

WO2021181116A1 - Method of manufacturing a medical device

Info

Publication number: WO2021181116A1
Application number: PCT/GB2021/050631
Authority: WO
Inventors: Moataz ATTALLAH; Parastoo JAMSHIDI; Hugh Hamilton
Original assignee: Johnson Matthey Public Limited Company; The University Of Birmingham
Priority date: 2020-03-12
Filing date: 2021-03-12
Publication date: 2021-09-16
Also published as: GB202003581D0

Abstract

The present disclosure relates to a method of manufacturing a tubular medical device which includes additive layer manufacture and laser cutting. Example embodiments include a method of manufacturing a tubular medical device, said method comprising: i) forming a metal alloy tube by additive layer manufacture (141); then ii) cutting a pattern into the tube by laser (144). The method may also include steps of tube drawing (142) and/or chemical etching (143).

Description

METHOD OF MANUFACTURING A MEDICAL DEVICE Field of the Invention

The present invention provides a method of manufacturing a tubular medical device which includes additive layer manufacture and laser cutting.

Background of the Invention

Stents are employed in a number of medical procedures and their structure and function are well known. Stents are typically tubular devices which can be introduced into bodily vessels by catheter, including coronary arteries, renal arteries, peripheral and other arteries as well as veins. Stents may also be used in the treatment of digestive tract strictures or in the support of collapsed abdominal cavities. Stents are typically radially expandable such that they can be introduced into a bodily vessel in an unexpanded state, before being expanded to the diameter of the vessel. When expanded, stents support or reinforce sections of lumen walls which have collapsed, are partially occluded, blocked, weakened or dilated.

There are many stent geometries and designs including, for example, those based on coils and helical spring designs, woven stents and those composed of individual and sequential rings. A review by Stoeckel et al. (Min. Inv. Ther. & Allied Tech. 2002, 11(4), 137- 147) provides illustrative examples. Particular stent designs vary with application to take account of in-service stress, delivery capabilities, etc. Of the sequential ring geometries there are closed cell designs, such as the Palmaz type stent, and open cell designs to allow the material to flexibly expand through having a negative Poisson’s ratio behaviour, also known as an auxetic behaviour. Open cell designs generally comprise a plurality of rings which comprise crests and valleys and are connected in an axial direction by struts or connecting links. Rings can be arranged and connected, for example, in crest to crest, crest to valley and offset crest to valley orientations. Fewer crests in each ring generally results in less scaffolding effect and reduced expansion range and vice versa, and fewer struts generally means more flexibility and less scaffolding effect and vice versa.

Popular materials for making stents include stainless steel 316L, tantalum, cobalt- chromium alloys, precious metal-containing alloys and shape memory materials such as nitinol. Nitinol is a nickel-titanium alloy generally containing about 50 to 52 at% nickel and 48 to 50 at% titanium. A benefit of nitinol stents is that they are chemically and biogenically inert without requiring special coatings. Nitinol is also an ideal material for use in stents because it is superelastic under certain conditions, allowing it to withstand extensive deformation and still resume its original shape. Also, the shape memory effect of nitinol means that it retains a set shape fixed during a particular heat treatment and, if deformed from that shape, can return to that original shape under proper conditions. The temperature at which the nitinol will return to its set shape is the austenite finish temperature A_F and marks the point at which the nitinol completes the change from being martensitic to austenitic upon heating of the alloy. Alloys of nickel and titanium can be prepared such that the A_F temperature is below normal body temperature but above room temperature. Stents manufactured from such alloys are desirably deformable prior to application to the body, and will resume their set shape naturally at normal body temperature. The A_F temperature is generally controlled by thermal aging of the alloy, the chemical ratio of nickel to titanium in the alloy, or by addition of other alloying elements that are typically added to stents (e.g. Cr, Pt, Pd) and tramp elements (mainly oxygen, nitrogen and carbon).

There are various methods for making nitinol stents, each posing some difficulties. A widely used method begins with production of sized nitinol ingot, from which the centre is removed to produce an initial tubular geometry. Subsequently, nitinol tubes of required diameter, length and wall thickness can be produced by drawing processes, for example, over a series of mandrels of progressively reduced diameter, before being cut into the required stent design pattern, for example, by laser cutting. The associated costs, time and complexity of nitinol tube drawing are greatly increased by the need for elevated temperature annealing stages interspersed between the mechanical drawing stages to maintain alloy workability. It will be recognised that uniform control of the process factors, including process cleanliness, annealing atmosphere, mandrel and die geometries and types of lubricants, is both challenging and key to the production of high quality tube. Furthermore, the wastage of material, especially during the production of the initial tube geometry from the ingot, results in significant added cost since the removed material is not easily recyclable.

An alternative to the drawing of tube for stent feedstock material is the drawing of wire, which is subsequently manipulated into the desired geometry using conventional forming techniques such as knitting or braiding. Although less technically demanding than forming tubular feedstock, for example, ensuring concentricity of wall thickness is not applicable to wire products, drawing of nitinol wires still requires a series of drawing and annealing processes and maintenance of similarly high standards of process control.

Welding steps in production of some wire formed stents can induce weaknesses in the structures that reduce mechanical robustness and stent lifetime. Poor selection of weld materials and/or poor finishing of weld joints can lead to corrosion occurring at regions of inhomogeneous composition.

Additionally, nitinol and other alloy stent designs can be manufactured directly by additive layer manufacture techniques, for example, using powder feedstocks in laser powder bed equipment. However, this procedure is complicated by the fact that extra support, beyond the struts that will remain in the final stent, is required between rings as each layer is built up. Such extra support can be difficult to remove cleanly from the final stent, leaving small irregular metallic spurs. Furthermore, stents manufactured directly via laser powder bed additive manufacturing have been found to possess poor as-made surface finish and inconsistency in dimensions of the struts/cross-pieces/wires making up the structure, resulting in poor mechanical properties. In extreme cases, struts can be incomplete, resulting in sharp spurs that would lead to tissue damage if implanted.

The challenges outlined above clearly illustrate that there is a need in the art for a timely, cost effective method of manufacturing stents, particularly nitinol stents.

Summary of the Invention

Accordingly, the present invention provides a method of manufacturing a tubular medical device, said method comprising: i) forming a metal alloy tube by additive layer manufacture; then ii) cutting a pattern into the tube by laser.

This method allows the avoidance of costly and time consuming tube drawing or wire manipulation processes. Moreover, this method allows for simple and reliable control over properties of the device such as wall thickness. Additionally, this method avoids the complications of preparing a medical device entirely by additive layer manufacture.

Step i) of the method may be carried out by powder bed deposition, involving: a) forming a layer of a metal alloy precursor powder; then b) binding or fusing the powder.

Step b) may be carried out by laser melting.

Forming the tube by powder bed deposition allows formation of uniform wall thickness tubes by forming the tube vertically, i.e. with the longitudinal axis oriented vertically during the forming process.

Step i) of the method may alternatively be carried out by a direct deposition process involving: g) flowing a stream of metal alloy precursor powder at a target area; and h) binding or fusing the powder at the target area simultaneously to step g).

Step h) may be carried out by laser melting.

The method may also comprise a step iii) of polishing, which may be carried out after step i), after step ii), or after both step i) and step ii). Polishing may be carried out by chemical etching. The step of polishing may reduce a wall thickness of the tube by between around 15% and 85%.

The method may also comprise a step iv) after step i) and before step ii) which includes one or more tube drawing steps. Said one or more tube drawing steps may include hot drawing, annealing or cold working or a mixture of one or more of these steps. A wall thickness of the tube may be reduced by between around 30% and 50% by the one or more tube drawing steps.

The metal alloy may be nitinol. The metal alloy precursor powder may for example comprise a mixture of Ni and Ti in which there is greater than 50 at% Ni. The metal allow precursor powder may have an increased proportion of Ni compared to the proportion of Ni in the tube to compensate for loss of Ni by evaporation during laser melting. The proportion of Ni may for example be increased by between around 1% and 3%.

The medical device may be a stent.

Brief Description of the Drawings

Figures 1a-b are images of a group of nitinol tubes on an additive layer manufacturing build plate and an individual tube, respectively, manufactured by a selective laser melting process.

Figures 2a-d are scanning electron microscope (SEM) images of the cross sections of walls of nitinol tubes manufactured by selective laser melting at various laser energy densities.

Figure 3 is a SEM image of a cross section of the wall of the tube shown in Fig. 2b after chemical etching.

Figure 4 is an image of two laser cut stents manufactured according to the method of the invention.

Figures 5a-c are SEM images of a laser cut stents manufactured according to the method of the invention.

Figures 6a-c are SEM images of the laser cut stent shown in Figs. 5a-c after a step of chemical polishing.

Figure 7 is an image of two laser cut stents manufactured from a commercially available, drawn, nitinol tube.

Figures 8a-c are SEM images of a laser cut stent manufactured from a commercially available, drawn, nitinol tube. Figure 9a-c are images of a series of 316L stainless steel tubes on an additive layer manufacturing build plate shown in different views and an individual tube, respectively, manufactured by a selective laser melting process.

Figure 10 is SEM image of the cross section of walls of 316L stainless steel tubes manufactured by selective laser melting by optimised process parameters.

Figure 11a-c are images of laser cut design manufactured according to the method of the invention illustrating the feasibility of repeating the same pattern along the additively manufactured tubes.

Figure 12 is an image of Ti-Nb-Ta-Zr (TNTZ) stents manufactured by a selective laser melting process. TNTZ stent before etching (left) and Etched TNTZ stent (right).

Figure 13a-b are SEM images of additively manufactured stents before and after etching.

Figure 14 is a flow diagram illustrating an example method of manufacturing a tubular medical device.

Preferred and/or optional features of the invention will now be set out. Any aspect of the invention may be combined with any other aspect of the invention unless the context demands otherwise. Any of the preferred and/or optional features of any aspect may be combined, either singly or in combination, with any aspect of the invention unless the context demands otherwise.

The medical device produced by the method of the invention is tubular, i.e. its body comprises walls which define a hollow tube. The cross sectional area perpendicular to the longitudinal axis of the hollow body may be polygonal, or substantially circular, preferably circular. The cross sectional area perpendicular to the longitudinal axis may vary along the length of the device, e.g. the device may have a variable diameter. Step ii) of the invention of cutting a pattern into the tube will be understood to require that the laser removes portions of the wall of the tube formed in step i), and leaves voids in the wall. Accordingly, the walls of the medical device comprise voids in the form of the pattern cut in step ii). So, it will be understood that, in the particular case of, for example, a sequential ring stent, the walls of the stent are defined by the rings and struts in the stent.

The wall thickness of the tubular medical device is not particularly limited and will depend on its intended use. For example, in the case of a stent, a wall thickness may suitably be in the range of and including 25mhi to dqqmhi, preferably in the range of and including dqmhi to 400mhi, more preferably in the range of and including dqmhi to 250mhi. As will be evident to a skilled person, wall thickness corresponds to the thickness of material in the radial direction which remains after a pattern has been cut into the metal alloy tube in step i) of the invention. So, in the specific example of a sequential ring stent, wall thickness is equivalent to the thickness in the radial direction of the rings and struts. Also, the cross sectional area perpendicular to the longitudinal axis of the device (which may vary along its length), and the length of the tubular medical device manufactured by the method of the invention are not particularly limited and will depend on its intended use.

The tubular medical device should preferably conform to guidelines laid out in industry accepted standards for such devices, depending upon their specific application, and the materials used for their construction. Such guidelines may be found, for example, in ASTM Standard F2063 which specifies the chemical, physical, mechanical and metallurgical requirements for wrought nitinol bar, flat-rolled products and tubing; ASTM Standard F2633 which specifies variables which differentiate drawn medical grade tube from the forms covered in ASTM Standard F2063; ASTM Standard F3306 which provides test method for assessment of metal or other ions released from single use implantable devices, or ASTM Standard F2004 which defines procedures for determining transformation temperatures of nitinol alloys. A skilled person will understand that these standards are listed by way of illustration only, the tubular medical device not being limited to the teachings of these standards.

Suitably, the medical device prepared by the method of the invention has a porosity of no more than 1% (volume percent), preferably no more than 0.5% (volume percent), more preferably, no more than 0.3% (volume percent), most preferably no more than 0.1% (volume percent), as measured by micro-computed tomography. Micro computed tomography (“micro-CT”) is an X-ray imaging technique whereby a focused X-ray source is used to irradiate an object; magnified projection images are collected on a planar X-ray detector. Images are collected at each step as objects are rotated stepwise through 180°. The series of images is reconstructed into 2-dimensional cross sectional slices, which can be further computed into a 3-dimensional reconstruction. With micro-CT, samples can typically be imaged with pixel sizes down to 100 nanometres.

Preferably, the tubular medical device is a stent. There are many stent geometries and designs including, for example, those based on coils and the helical spring design, woven stents and those composed of individual and sequential rings. A skilled person is aware of such stent designs. Suitably, a stent manufactured by the method of the invention is of the sequential ring design, for example having a closed cell ring geometry such as the Palmaz type stent, or an open cell geometry. Open cell designs generally comprise a plurality of rings which comprise crests and valleys and are connected in an axial direction by struts or connecting links. Rings can be arranged and connected, for example, in crest to crest, crest to valley and offset crest to valley orientations. Fewer crests in each ring generally means less scaffolding effect and less expansion range and vice versa, and fewer struts generally means more flexibility and less scaffolding effect and vice versa. As will be understood by a skilled person, the rings and struts in such a stent will correspond to the wall material left after the pattern has been cut in step ii) of the method of the invention.

The step of additive layer manufacture, which is also known, for example, as layer manufacturing, constructive manufacturing, generative manufacturing, direct digital manufacturing, freeform fabrication or solid freeform fabrication may be applied to metal alloy tube manufacture using known techniques. In all cases, the additive layer manufacture processes are enabled by conventional 3-dimensional design computer packages that allow design of the shaped unit as a so-called “STL file”, which is simply a mesh depiction of the 3- dimensional shape. The STL file is dissected using the design software into multiple 2- dimensional layers, which are the basis for the fabrication process. The fabrication equipment, reading the 2-dimensional pattern, then sequentially deposits layer upon layer of powder material corresponding to 2-dimensional slices. In order that the shaped unit has structural integrity, the powder material is bound or fused together as the layers are deposited. The process of layer deposition and binding or fusion is repeated until a robust shaped unit is generated. Any unbound or unfused powder is readily separated from the shaped unit, e.g. by gravity, or blowing.

A number of additive layer manufacture binding and fusion fabrication techniques are available, notably selective laser melting techniques. In the present method, selective laser melting is preferred, which can be performed in systems such as the Concept Laser 400 W M2 Laser Cusing Powder-bed machine. Selective laser melting may also be known as powder bed fusion, direct laser metal sintering or melting and is a technique whereby 2- dimensional layers (typically of depth 5 to 100 mhi) of powdered materials (with a typical powder size in the range of 15 m to 60 mhi as determined by the sieve method as stipulated in ASTM Standard B214) are sequentially laid down and fused or bound together to form 3- dimensional solid objects using a laser or an electron beam.

In selective laser melting, the process comprises three main steps. In a first step, a thin layer of powder material is initially deposited onto a base plate using, for example, a blade, roller, or moving hopper. The thickness, which corresponds to thickness in the axial direction of the tube that is being manufactured, of the layer is controlled. Laser radiation is then applied to melt and fuse the layer. The laser position is controlled, e.g. using galvanometer mirrors, according to the desired pattern. After the layer is fused, the plate on which the layer rests may move downwards by the thickness of the layer and a fresh layer of powder is deposited over the fused layer. The procedure is repeated thus producing the shaped unit in three dimensions. The procedure is suitably performed in a controlled atmosphere of inert gas.

Accordingly, step i) of the invention is preferably carried out by: a) forming a layer of a metal alloy precursor powder; then b) binding or fusing the powder. Step b) is preferably carried out by laser melting.

A metal alloy precursor powder in the context of the present invention is a powder which forms an alloy during the step of additive layer manufacture. For example, it is a mixture of different metals in powder form that will form the desired alloy on binding or fusing.

In the present invention, the depth of each powder layer deposited is suitably in the range of and including 5 to 100 pm, preferably 10 to 30 pm, more preferably 15 to 25 pm, for example about 20 pm. Also, when laser melting is used to bind or fuse a metal precursor powder, the laser power is suitably in the range of an including 10 to 500 W, preferably 50 to 350 W, more preferably 80 to 200 W. Additionally, the scan speed of the laser is suitably in the range of and including 20 to 750 mm/s, preferably 100 to 500 mm/s.

In the preferred embodiment of the invention in which selective laser melting is used it is a benefit of this invention that the wall thickness of the metal alloy tubes manufactured in step i) can be controlled by controlling the energy density of the laser in step b). The desired wall thickness will depend on the intended design and application of the tubular medical device being manufactured. For example, the linear energy density of the laser may suitably be no less than 0.3 J/mm and no more than 1.5 J/mm, preferably no more than 1.3 J/mm.

For example, such a linear density will provide a nitinol tube having a wall thickness in the range of and including 300 to 400 pm, an optimal range for preparing nitinol stents.

An alternative additive layer manufacturing technique is known as directed energy deposition. Directed energy deposition may be defined as an additive manufacturing process in which focused thermal energy is used to fuse materials by melting as they are being deposited. By way of example, when the process feedstock is in powder form, the flow of the stream of powder is directed into a targeted area and melted by a focused energy source, such as a laser or an electron beam, so forming a metallurgical bond with the underlying surface.

Accordingly, step i) may alternatively be carried out by: g) flowing a stream of metal alloy precursor powder at a target area; and h) binding or fusing the powder at the target area simultaneously to step g). Step h) is preferably carried out by laser melting.

This technique may also be known as blown-powder deposition, laser metal deposition, direct energy deposition, direct metal deposition, laser engineered net shaping, direct laser deposition or laser cladding.

When directed energy deposition is used, the wall thickness and surface finish of the metal alloy tubes manufactured in step (i) can also be controlled by controlling the energy density of the power source in step h) and can also be controlled by the flow rate of the powder stream. Faster material deposition rates may be achieved with the directed energy deposition process than with the selective laser melting process.

An alternative additive manufacturing technique involves the replacement of the powder feedstock with a wire feedstock. Wire-feed additive manufacturing can be classified into three groups depending on the energy source used for deposition; wire and arc additive manufacturing, electron beam freeform fabrication and wire and laser additive manufacturing.

The metal alloy tube formed in step i) of the invention can be formed of materials suitable for the medical device application. Such materials include stainless steel alloys, tantalum, cobalt-chrome alloys, alloys comprising suitable metals from the platinum group metals, i.e. Pt, Pd, Rh, Ru, Ir and Os or gold or silver, and nitinol alloys. The metal alloy tube is preferably a nitinol tube, and the medical device is preferably a stent. In order to prepare a nitinol tube, the metal alloy precursor powder used is suitably a mixture of Ni and Ti in which there is greater than 50 at% Ni. Preferably, for selective laser melting the powder contains no less than 52 at% Ni. The amount of Ni in the precursor powder is important for the desired properties of the stent manufactured by the method. In order for a nitinol stent to have an A_f temperature in the region of about 30 to 35°C such that the stent is deformable prior to application to the body, and will resume its set shape naturally at normal body temperature (i.e. about 38 °C), the stent should preferably have a Ni content of no less than 51 at%. Commercially available nitinol powders tend to contain the at% Ni required for the nitinol to meet these requirements. However, Ni can be lost by selective vaporisation during the additive layer manufacture process and so, when a commercial nitinol powder is used it is preferable to add Ni such that the Ni content of the powder becomes no less than 52 at% Ni, and no more than 55 at% Ni. For example, 2 at% Ni powder may be added to an as received powder comprising less 52 at%.

Nitinol powders preferably have particle size in the range of and including 15 to 60 mGh, a D10 particle size of no more than 15 mhi, and a D90 particle size of no more than 60 mGP. Powder particle size is determined by the sieve method as stipulated in ASTM Standard B214, and the particle size distribution is measured using the laser diffraction technique in accordance with ASTM Standard B822.

The medical device may also contain additives to modify properties, such as platinum, palladium and gold to increase radiopacity; cobalt, iron, aluminium and chromium to suppress martensite; copper and niobium to modify hysteresis behaviour; hafnium and tantalum to stabilise martensite; copper or silver to induce antibacterial properties, which can be incorporated in the precursor powder used for additive layer manufacture. It is recognised that modifications to the ratio of Ni to Ti in such compositions may be required to accommodate the added additives.

A step iii) of polishing is preferably carried out after step i) or after step ii), or after both step i) and step ii). Preferably it is carried out after step ii). It is important, especially in the context of a stent, for the medical device to have a smooth, defect free surface. As such, surface polishing steps can include use of abrasive media, chemical etching or electrochemical polishing processes. Preferably, the polishing is carried out by chemical etching. Chemical etching is suitably performed by immersion in an acidic media, for example an acidic solution e.g. HFiHClihhO or HF-HNO3 solution. Immersion may, for example, be carried out for a period of no more than 1 hour, suitably no more than 30 minutes, for example no less than 5 minutes. Moreover, the polishing step, in particular chemical etching, can be used to control the wall thickness of the medical device so that it is in the required range for a particular application. For example, the thickness of a nitinol tube can be reduced by an amount in the range of and including 50 mhi to 300 mhi by the chemical etching step of the invention. As will be understood by a skilled person, in the case of a stent, chemical etching or electrochemical polishing can be used to modify, for example, strut size and ring size. For example, both the thickness in the radial direction and thickness in the circumferential direction of the struts and rings can be modified. In a general aspect therefore, the step of chemical etching may result in a reduction in wall thickness of the tube by between around 50 mhi and 300 mhi and may apply to any metal alloy. Alternatively stated, the step of chemical etching may reduce the wall thickness of the tube by between around 15% and 85%.

Preferably, there are no additional steps of manipulating the metal alloy tube formed in step i) before step ii) is carried out, other than the step of polishing mentioned above. Put another way, it is preferable that the as-made tube is used directly in step ii) to cut the medical device. However, it may be desirable to manipulate the metal alloy tube further after step i) and before step ii) of the method. Accordingly, the method of the invention may also comprise a step iv) after step i) and before step ii) which includes one or more tube drawing steps. Suitably, said one or more tube drawing steps include hot drawing, annealing or cold working or a mixture of one or more of these steps. For example, a limited number of tube drawing and optional annealing stages to further reduce or unify the wall thickness and/or homogenise the microstructure may be employed. The number of drawing processes would be kept to a minimum but may typically be in the range of and including four and six, with annealing processes included as necessary. Such drawing processes may include cold working to confer desired mechanical and functional properties upon the final device. A skilled person will recognise that the necessary extent of such cold work will vary but typical amounts can be in the range of and including 30 to 50%. A step iii) of polishing is carried out carried out after step iv) and may also be carried out before step iv). Preferably, a step iii) of polishing is carried out before and after step iv). In a general aspect therefore, the step of tube drawing may reduce the wall thickness of the tube between around 30% and 50%. The reduction in wall thickness through tube drawing may be in addition or instead of the wall thickness resulting from a chemical etching step. An advantage of reducing the wall thickness, either by tube drawing or chemical etching, is that an initial wall thickness achievable by powder bed deposition may be too large for a practical tubular stent. Tube drawing can both improve the surface finish of the tube as well as achieving the desired final wall thickness.

Examples

The method described herein, which combines additive manufacturing with laser cutting, may be used for not only manufacturing of a tubular precursor medical device but also can be applied to manufacture of pre-form design, (e.g. stents with thicker struts) from various metallic biomaterials alloys. The device can be subsequently tailored and adjusted to a final design. Several examples of this invention are provided below and illustrated in the drawings.

Nitinol tubes

A series of nitinol tubes was manufactured by selective laser melting using a Concept 400W M2 Laser Cusing Powder bed machine. The laser power was 80W, the scan speed was 250 mm/s, and the powder layer thickness was 20 pm. A simple laser scan strategy, i.e. single axis laser scans, was used. The powder used was a commercial nitinol powder having an atomic composition of TUe_.gNisi_.i blended with ~2at% Ni. The additional percentage of Nickel was added to compensate for evaporation of some Ni during laser melting to result in a correct final composition. An image of the series of nitinol tubes is provided in Fig. 1a, with a single tube shown in Fig. 1b. Additional series of tubes were manufactured using the same method and nitinol powder, but using laser energy densities of 1.2 J/mm, 0.8 J/mm, 0.6 J/mm, 0.4 J/mm. The tubes had wall thicknesses of 393, 363, 384 and 310 pm respectively. SEM images of cross sections of the walls of these tubes are shown in Fig. 2, obtained using A JEOL6060 field scanner operated at 10-20 kV (used for all SEM imaging herein). Accordingly, it is possible to optimise the thickness of the tube for a particular stent design and application by varying the laser energy density during selective laser melting.

A tube manufactured using a laser energy density of 0.8 J/mm and having a wall thickness of 363 pm was chemically etched using HFiHCkhhO solution, immersion for 10 minutes, to provide a tube having a wall thickness of 150 pm. An SEM image of a cross section of the wall of the tube is shown in Fig. 3. Thus, there is further possibility to optimise wall thickness of the tube for a particular stent design and application. Chemical etching also has the function of polishing the tube surface and reducing the porosity of the surface to an acceptable level.

The chemically etched tube was laser cut to provide stents with two different designs, zig-zag and Palmaz, as shown in Fig. 4. Laser cutting was performed using a reconfigurable multi-axis processing system with sub-pico and nanosecond laser sources. Laser cutting is usually conducted in the power range of 5 to 50 W. The system can import the STL file for the required geometry, and the laser is passed in prescribed tool paths to create the holes according to the required dimensions. SEM images of the zig-zag stent are provided in Figs. 5a-c. The surface is similar in quality to stents laser cut from commercial, drawn, nitinol tubes, shown in Figs. 7 and 8a-c. A further chemical etching step using HFiHCkhhO solution, immersion for 30 seconds was performed further polish the stent and SEM images of the polished stent are shown in Figs.6a-c, Fig. 6c being a magnification of the surface. As can be seen in the figures, partially melted particles are removed from the stent, and the surface is similar to that of the commercial stents shown in Figs. 7 and 8a-c.

316L stainless steel tubes

A series of 316L stainless steel (SS) thin-walled tubes was manufactured by selective laser melting using a Concept 400W M2 Laser Cusing Powder bed machine. The laser power was 250W, the scan speed was 1300 mm/s, and the powder layer thickness was 20 pm. A simple laser scan strategy, i.e. single axis laser scans, was used. The powder used was a gas-atomised powder with particle size distributions in the range 15-45pm . An image of the series of 316L stainless steel tubes is provided in Fig. 9a-c.

One of the SS tubes was manufactured using the optimised process parameters of 250w laser power, the scan speed of 1300 mm/s and layer thickness of 20 pm was checked under scanning electron microscope. An SEM image of a cross section of the wall of the tube is shown in Fig. 10. The tube had wall thicknesses of -500 pm. In the case of 316 L SS tubes also accordingly, there is further possibility to optimise wall thickness of the tube for a particular stent design and application by varying the laser energy density during selective laser melting and also by etching technique.

The additively manufactured SS tube was then laser cut to provide specific cross like design as shown in figure 11 a and b. Figure 11c illustrates the feasibility of repeating the same pattern along the additively manufactured tubes by Laser cutting. Laser cutting was performed using the same procedure that was used for nitinol tubes.

Ti-Nb-Ta-Zr (TNTZ) stents

Stents from a b-Ti alloy Ti-34Nb-13Ta-5Zr (TNTZ) alloy, as promising biocompatible implant materials, were manufactured by selective laser melting using a Concept 400W M2 Laser Cusing Powder bed machine. The laser power was 320 W, the scan speed was 500 mm/s, and the powder layer thickness was 20 pm. A contour scan was used. An image of additively manufactured TNTZ stents is provided in Fig. 12. As -fabricated TNTZ stent (left) and chemical etched stent (right).

The additively manufactured stents were characterized using scanning electron microscope. Figure 13a and b shows the surface quality of the stents in both condition of before after chemical etching. Also SEM images illustrates the effect of etching on reduction of strut size. The present invention for this case can be used to tailor and adjust the pre-form stent design to its final optimum design.

Figure 14 shows a simplified flow diagram of a method of manufacturing a tubular medical device. In a first step 141 the tubular shape is formed by additive layer manufacturing. As described above, this may be done by powder layer deposition in which layers of powder are selectively fused by a scanned laser. Following step 141 the tube may be subjected to a tube drawing process (step 142), which reduces the wall thickness and smooths the internal and external surfaces of the tube. The wall thickness in the drawing process may for example be reduced by over around 30% and up to around 50% compared to the thickness of the tube as formed. Following this, a chemical etching or polishing process is applied (step 143), which cleans the surfaces of the tube, removing any burrs and debris. A laser cutting process (step 144) forms the required pattern in the tube wall. The laser cutting process at step 144 may alternatively be carried out before the chemical etching step 143 or an additional etching or polishing step may be carried out after laser cutting to remove any burrs or debris resulting from laser cutting.

Claims

1. A method of manufacturing a tubular medical device, said method comprising: i) forming a metal alloy tube by additive layer manufacture; then ii) cutting a pattern into the tube by laser.

2. A method according to claim 1 , wherein step i) is carried out by: a) forming a layer of a metal alloy precursor powder; then b) binding or fusing the powder.

3. A method according to claim 2, wherein step b) is carried out by laser melting.

4. A method according to claim 3, wherein a longitudinal axis of the tube is oriented vertically during step i).

5. A method according to claim 1 , wherein step i) is carried out by: g) flowing a stream of metal alloy precursor powder at a target area; and h) binding or fusing the powder at the target area simultaneously to step g).

6. A method according to claim 5, wherein step h) is carried out by laser melting.

7. A method according to any preceding claim, wherein the method also comprises a step iii) of polishing, which may be carried out after step i), after step ii), or after both step i) and step ii).

8. A method according to claim 7, wherein said polishing is carried out by chemical etching.

9. A method according to claim 8, wherein the step of polishing reduces a wall thickness of the tube by between around 15% and 85%.

10. A method according to any preceding claim, wherein the method also comprises a step iv) after step i) and before step ii) which includes one or more tube drawing steps.

11. A method according to claim 10, wherein said one or more tube drawing steps include hot drawing, annealing or cold working or a mixture of one or more of these steps.

12. A method according to claim 10 or claim 11, wherein a wall thickness of the tube is reduced by between around 30% and 50% by the one or more tube drawing steps.

13. A method according to any preceding claim, wherein the metal alloy is nitinol.

14. A method according to claim 13, wherein the metal alloy precursor powder comprises a mixture of Ni and Ti in which there is greater than 50 at% Ni.

15. A method according to any preceding claim, wherein the medical device is a stent.