MANUFACTURING OF STENTS: OPTIMIZE THE STENT WITH NEW MANUFACTURING TECHNOLOGIES

Transcription

1 12 MANUFACTURING OF STENTS: OPTIMIZE THE STENT WITH NEW MANUFACTURING TECHNOLOGIES Andreas SCHUESSLER, Ullrich BAYER, Gerd SIEKMEYER, Rainer STEEGMUELLER, Martin STROBEL, Admedes SCHUESSLER We proposed a review of the different driving forces for further development of stent manufacturing technologies. Influences from the market, the clinics and the customers lead to manifold research and engineering activities. We focused on developments in the area of materials development, including biodegradable stent materials, and on developments and optimization of manufacturing and surface modification technologies. 1

2 INTRODUCTION Since the invention of the balloon expandable stent by Palmaz in 1988, and FDA approval for peripheral vascular in 1990 and for coronary vascular procedures in 1994, more than 10 million people in the US have undergone a coronary or peripheral stenting procedure. Such a unique and rapid development in this field of medicine in only 12 years was based on the significantly better results obtained for stenting compared to standard balloon angioplasty. Angioplasty was successful in reopening clogged arteries, however in many cases, immediate reclosure or arterial recoil caused failure of the angioplasty procedure. Stents provide a mechanical means of resisting recoil of the artery after balloon angioplasty, but do not prevent (and may to a certain extent even cause) another failure mode known as restenosis. Restenosis is a reclosure of the treatment site due to cell proliferation caused by local injury and a subsequent tissue response. With increasing clinical application of and experiences in stenting and increasing production quantities, requirements for stents and their manufacturing technologies have changed, too. FACTORS IN OPTIMIZING STENT TECHNOLOGIES We divided the driving forces for technology development in the area of stents into three basic groups: design requirements, material related properties, and market aspects as demonstrated in Figure 1. These aspects are affecting and will continue to affect development in manufacturing technologies in different ways and at different levels. The details will be discussed in the following chapters Market Radiopacity Price Visualisation and diagnostic techniques Customized surfaces Material Clinical Requirements Material combinations Fatigue resistance Biodegradability MRI visibility Miniaturization Manufacturing Technologies Design Tolerances Drug Adhesion Figure 1: Influences on manufacturing technologies for stents. MARKET INFLUENCES CELL CULTURE TRIALS The primary market driver for stents has been the relative ease that most vessels in the body accept a foreign body such as a stent implanted on a permanent basis. Ongoing optimizations of the procedure continue however: a) prevention of acute and/or late restenosis within the stent via drug coatings on the stent surface; b) reduction of the profile of the device to the smallest possible size in order to minimize the invasiveness of the procedure and allow access to even the most distal locations; c) increase in stent durability in harsh environmental anatomic locations such as the superficial femoral artery (SFA) where torsion, radial and longitudinal tension and compression combine to create material fatigue challenges; d) creation of a device that can be seen easily under x-ray visualization techniques (also known as fluoroscopy); 2

3 e) creation of a device that can be visualized via non-invasive techniques such as magnetic resonance imaging (MRI). Regulatory bodies continue to try to en-sure safer and more efficacious devices by requesting larger clinical studies with longer term follow-up. Due to the large amount of devices recalled either voluntarily or by regulatory injunction from the market in the last five years 1. The Federal Drug Administration (FDA) is announcing fundamental changes in the way the agency will track medical devices. Recently, adverse effects (late thrombosis with deadly outcomes) were reported for long-term clinical studies for drug eluting stents (CYPHER stent of Cordis and TAXUS stent of Boston Scientific). Also, broken struts and broken stents in peripheral arteries, especially in the SFA are a concern of the FDA. The FDA has undertaken different means to tighten the regulatory requirements related to stents. Among those are the decrease of off-label uses and potentially the removal of 510 k approvals (a notification of intent to market) in favour of a PMA process (a request to market) and the application of 21 CFR part 820 QSR for contract manufacturers. Finally, cost containment has been an ongoing struggle in health care for the last decade and with the aging population, this will continue to be a major economic drain on society. MATERIAL CHOICE AND FUNCTION Conventional stent materials Stainless Steel, Tantalum, Niobium, CoCr alloys and Nitinol have been introduced as materials for stents in the past 15 years. However, unmet medical needs and general technological advancements have led to additional efforts to improve materials properties and to develop new materials for vascular stents, especially biodegradable materials such as magnesium and degradable polymers. Biodegradable materials Polymer materials like poly (l-lactid acid) PLLA seem to be generally applicable for drug eluting stents and if they finally make their way into a commercial application may offer cost advantages in stent manufacturing compared to the conventional manufacturing processes of today s stents. As an alternative to polymers, the use of degradable metallic materials (Figure 2), such as magnesium alloys and iron, have been investigated during the past 5 years. In general, degradable materials can be used for indications which do not require a permanent mechanical enlargement of the artery. However, the main advantage of biodegradable materials is assumed to be that any long term risk associated with a foreign object implanted in the body is obviated and no long term mechanical adverse effects of conventional stents, such as stent crushes or strut fractures 2 to what extent the degradation process and degradation products might effect long term patency is unclear and subject to investigations As such, results on reduction of restenosis rate of biodegradable stents (magnesium) are inconsistent, ranging from a significant reduction 3 to no difference compared with balloon angioplasty 4. However, one of the main challenges in the development of Mg stents is the control of the degradation mechanism and the lack of x-ray visibility 5,6 (Figure 3) and is therefore subject of intensive investigation 7,8,9,10. 3

4 J&J Palmaz: SS Wiktor: Tantalum Inflow Dynamics Starflex: Niobium Medtronic Driver: MP35N (Cobalt Alloy) Biotronik AMS: Mg 1990 s 2005 Schneider Wallstent: Elgiloy* Angiomed Memotherm: Nitinol Angiodynamics AngioStent: Pt Ir Guidant Vision: L605* Abbott TriMaxx: SS/Ta Composite Figure 2: Stent Material Introduction Figure 3: Magnesium stent (top) versus stainless steel stent (bottom) approximately 30 days after implantation in porcine coronaries 5. Biopolymers Biodegradable polymers possess a variety of challenges, such as radial force, recoil, large profile, degradation rate and products, biocompatibility and radiopacity. However, there are an increasing number of development programs with polymer stents 11 : a. The Igaki-Tamai stent is made from PLLA and features a zig-zag design which is deployed using a heated balloon b. The Reva Medical stent uses tyrosinederived poly (DTE carbonate). This stent uses ratchet links in order to achieve a higher radial force and is loaded with iodine for inherent radiopacity c. The BVS stent (Bioabsorbable Vascular Solutions, Guidant) is made from PLLA loaded with the drug everolimus and features both a self expandable and a balloon expandable design Table I shows a comparison of different stent material groups with regard to technical and market aspects. 4

5 Table I. Properties and trends of different stent materials Remarks Polymer stents Metallic stents Degradable (non degradable not Degradable (Mg-, Fe-alloys) Non-degradable (Nitinol, SS, CoCr, Ta) considered herein) Current market value n/a n/a ++++ Expected growth rate Clinical experience Trends in clinical acceptance Manufacturing by Casting, injection moulding Laser cutting Laser cutting, Wire technologies Price trends None marketed yet Mg SS, CoCr, Nitinol Main advantages Degradability Drug release Degradability No inflammatory effects compared to polymers Proven biocompatibility Radiopacity High radial force (SS, CoCr) Superelasticity and excellent wall apposition (Nitinol) Main disadvantages Mechanical properties Need for further research Recoil Local inflammation Poor vizualisation Poor mechanical properties Radiopacity and MRI visibility For exact positioning and release at the treatment site, good visibility of the stent under fluoroscopy is required. Due to materials of high atomic number conventional metal stents have significantly better fluoroscopic visibility compared to magnesium or biodegradable polymer stents (Figure 4). Nevertheless, some indications, especially for peripheral stents require an improvement of the radiopacity by the addition of markers from tantalum or gold at the end of stents by riveting or welding 12. So far, new concepts Low radiopacity Degradation uncontrolled Fatigue issues No potential as drug carrier to improve the radiopacity of the less visible degradable materials are lacking. The radiopacity can also be improved by direct alloying elements or ion implantation with a higher atomic number into the material. For Nitinol the addition of ternary elements like platinum or palladium which still retains the mechanical properties of the base material are described in literature 13,14. Figure 5 shows another new technological solution to enhance radiopacity. 5

6 Figure 4: Similar stent designs manufactured from different materials. By welding or laminating a sandwich material (e.g. Trimaxx manufactured by Abott Vascular, stainless steel / tantalum / stainless steel) radiopacity of a small stent can be improved appreciably. Figure 6 : Stent images during regular MRI investigations (Source: Biophan Technologies Inc.). In contrast, new surface coatings implementing resonator technology, which uses tuned circuits to increase the radio frequency (RF) signal, make it possible to image within and around a stent (Figure 7). Figure 5 : Trimaxx Tube concept of Abott Vascular Perclose, Inc. Material interaction with diagnostic clinical visualization tools is important during MRI. The magnetic field might cause heat and forces on the stent component due to its ferromagnetic material behaviour. As the majority of today s stents are made out of nonmagnetic materials, most stents are MRI safe. However, a stent acts as a Faraday cage (it shields radio-frequency signals) and creates a large image artefact. (Figure 6) shows a standard vascular stent imaged under MRI. Figure 7 : Improved stent MRI visibility by coating the metallic stent component (Source: Biophan Technologies Inc.). Customized surfaces Surface micro patterning Standard drug eluting stent (DES) technology places active agents which are antagonists of restenosis into a polymer matrix. Drugs are released into the surrounding tissue by a controlled degradation process (Figure 8). 6

7 Figure 8 : Drug release through polymer drug pads which are embedded e.g. by laser cutting cavities (Source: Conor Medsystems Inc.). Today, most DES are directly drug coated, however, recent clinical studies have shown an increased risk of late thrombosis which is attributed to this solution. Newer technology routes use small cavities as reservoirs within stent struts (Figure 8) or in cavities along the strut as shown in Figure 9 below. Figure 10 : (a) Stent structure with micro structure, (b) porous surface from oxides, (c) porous surface from metal Another way to get functionalized surfaces with a higher medical performance is described in the US Patent 6,190,404. It refers to a special surface topography of an intravascular stent. The morphology is shown in (Figure 11). The grooves are suggested to promote the migration of endothelial cells onto the stent surface. Different methods to produce surface morphologies for improved endothelialization such as chemical etching or etching by laser treatment have been described (US Patent 2002/ ). Figure 9: Cavities to hold drugs by surface structuring (Source: Sorin Biomedical) In contrast to the above discussed routes new technology to manufacture porous surfaces (as displayed in Figure 10) are possible, too. Porous surfaces can be built up from oxides or metallic materials. Figure 11 : Cross section of a stent with a grooved inner surface (US Patent 6,190,404). Fracture and fatigue resistance Fracture and fatigue of vascular stents, such as coronary stents, AAA stent grafts and lately SFA stents has been a concern for a while While failures in coronary stents and AAA stent grafts seem to be highly attributed to high cycle fatigue, most 7

8 stent fractures in superficial femoral arteries occur in relatively short time periods and are due to the high strains/stresses applied to the implants due to the pronounced mechanical loading (extension, compression, torsion) in the limb. The fracture and fatigue behaviour has been mainly addressed by design improvements. However, it is well known and accepted that stent manufacturing processes do also affect the fatigue behaviour. Excessive heat input during laser cutting can cause micro cracking and needs to be avoided. Consequently, working with small laser spot sizes and short pulse duration which lead to smaller heat affected zones is a preferred method. It is also state-of-the-art that the heat affected zone should be entirely removed during the subsequent processing steps. It has also been recognized during the past years that electro polishing does not just produce corrosion resistant biocompatible surfaces, moreover edge rounding and the removal of micro cracks and surface imperfections improves the fracture and fatigue behaviour even further. With the use of appropriate laser cutting and electro polishing techniques and the careful adjustment of all process steps the inducement of flaws leading to early fracture and fatigue can be avoided. Stent material and raw material processing is another parameter affecting the fatigue behaviour and is unfortunately not very well understood. This is especially true for Nitinol due to its complex microstructure and material processing cycle. However, there is a number of results suggesting that at least at high alternating strains (low cycle fatigue) - which is the predominant load for self expanding stents in the SFA - is associated with TiC and/or TiNiO inclusions in the raw material which act as a crack initiator This is supported by our own investigations showing that different tube material (different ingot suppliers and tube manufacturers) result in different lifetime of test samples (Figure 12). Strain [%] Strain vs. Number of Cycles Diagram Cycles to Failure [10E3] Figure 12 : Comparison of the fatigue behaviour of different Nitinol tubing materials determined on MicroBone (Tensile Test) samples from tube OD: 2.3mm, wall:0.24 tubing at 20 C and a cycling frequency of 12Hz MANUFACTURING TECHNOLOGIES Laser machining Solid state lasers (Nd:YAG) using flash lamps is the standard beam source for stent cutting. These lasers have been optimised throughout the past 15 years for precision cutting and for approximately the past 7 years even special stent cutting laser machines have been made commercially available. For stent cutting of SS, CoCr or Nitinol stents aspect ratios (Cutting curve width / wall thickness) of about 12:1 can be achieved. The major disadvantage of those lasers is the inherent formation of a temperature gradient within the laser rod during operation. This causes lensing effects and mechanical stresses in the rod leading to a disturbed beam shape with a lower beam quality. The trend towards device miniaturization and tighter tolerances for stents requires smaller cuts and less heat input and con- A B C D E 8

9 sequently improved beam quality and pulse durations. In recent years, new promising laser resonator architectures have been developed. In fact, diode-pumped solid-state lasers (DPSSL) provide a sharper power distribution i.e. an improved beam profile quality. Nevertheless, mechanical resonator instability effects still can affect the cutting process. Fiber lasers appear to present the solution for this issue. Housed within a glass fiber OD <0.5mm, no misalignment can occur; an optimized beam profile is ensured. Meanwhile, serviceable laser diodes ready for direct fiber pumping with short powerup rise times provide high peak powers for good cutting quality. However, the industriial capability and reliability needs to be demonstrated. Although optimized melt cutting lasers are available, the achievable cutting width for the final product is still limited downwards. Since laser melt cutting is a thermal process, the formation of a heat affected zone (HAZ) is not avoidable. The enlargement of the cutting width related to the removal of the HAZ leads for a given wall thickness ~300µm to a reduced achievable aspect ratio for the final product of 12:1. In the case of inert gases used for the process (which is preferable for fine cutting of Nitinol), burrs and residual parts remain within the tubing and the cut structure and need to be removed semi-automatically. As a result of mechanical impact, a deformation of the part can occur. Also, melt removal by a gas stream momentum is strongly restricted to small cutting widths since an appropriate steam channel is not established below ~8µm. The fineness of possible structures is therefore limited (Figure 13). Figure 13 : SEM image of laser cut Admedes logo in Nitinol tubing (top) and the 10µm cutting width of the H letter (bottom). Ultrashort-pulse lasers offer another way of cutting materials. No melt is produced as the laser transfers the material directly from the solid to the vapour phase. Therefore, no burrs or slag are produced leading to a clean surface immediately. Also, due to the very short energy impact, the whole irradiated material is removed by vaporisation and HAZ is not formed. The lateral resolution is therefore diffraction-limited leading to the ultimate miniaturization. However, beam cauterization reduces the possible miniaturization, as the depth of focus leads to an enlargement of the cutting width at the upper and lower slot end (Figure 14). 9

10 Figure 14 : SEM images of a resorbable polymer stent machined with a femto laser system (top) and magnification of the cutting edge (bottom). Photo-chemical etching Photo-chemical etching or photo-chemical machining has been used to manufacture stainless steel stents in combination with forming and welding (NIR stent of Medinol) or Nitinol stents in combination with shape setting (aspire stent of Vascular Architects). Although photo-chemical etching possesses a number of advantages, such as stress- and burr-free machining, suitability for all conventional stent materials and broad use of the process has not seen (Figure 15). One of the main disadvantages is the low aspect ratio depth/width of about 1:1 witch prevents to apply the technique to stent designs of standard coronary and peripheral stents (aspect ratio ~ 5:1 to 12:1). Figure 15: Spire stent of Vascular Architects. The frame is manufactured by photo chemical etching of Nitinol sheet (left) covered with eptfe material (right). Thin film technology Manufacturing of three dimensional structures of Nitinol for the use as stents, stent grafts and related devices by thin film (sputter) technology followed by photoetching have been described for example by Gupka et al. 23 and Banas et al. 24. The proposed manufacturing techniques include planar sputter deposition methods as well as sputter deposition on a rotating 3D substrate. Manufacturing of NiTi thin films, first investigated by the Japanese in the late 80s (JP A) has to be done in an ultra high vacuum (UHV) atmosphere and is not a simple technique. Consequently, there have been only a few research groups worldwide which were able to achieve promising results on the functional properties of thin film nitinol structures. The sputter deposition method for producing medical devices such as stents has the advantage of making very thin devices (in the range of a few micrometers) with small features and tight tolerances (Figure 16). Sputter deposited NiTi films are extremely pure containing no contaminants and are claimed to be less throm- 10

11 bogenic compared to conventionally manufactured material 25. Although sputter deposition seems to have distinct advantages for niche applications (e.g. devices for very small vessels), high manufacturing costs and an unfavourable patent situation seem to hamper a wide penetration of this technology (Figure 17). Figure 16: Thin film cylinder manufactured by sputter deposition 15 Figure 17 : Stent Graft consisting out of a fine NiTi mesh produced by sputter deposition attached to a NiTi stent conventionally manufactured 16 Electroforming Manufacturing of high aspect ratio micro metal parts using electroforming involves photolithography to form the structural information (image) in a photoresist and electroplating of the metal at the openings in the resist. This process, for example in the form of the LIGA-technique, is a standard process in MEMS technology. The company ESI - Electroformed Stents Inc., has developed an electroforming process to produce stents from Gold, Silver or Nickel. The ESI process coats a conductive mandrel with a photoresist. The resist is exposed to a UV light source to form the stent image on the cylinder. Developing and stripping the resist forms the openings in which the metal is deposited. The photoresist and mandrel are dissolved, leaving the electroformed stent. (Source: Electroformed stents can be manufactured very precisely and, if done in a batch process, very cost effectively. The disadvantage, however is that it is has not been achieved applicability of components out of stainless steel, cobalt-chromium alloys or nitinol. A commercial use of electroformed stents is not known. CONCLUSIONS At a first glance, requirements from the market to improve treatment options with stents will be more likely achieved by new and improved stent materials rather then by the optimisation and development of new manufacturing processes. In case degradable materials will start to penetrate the market new manufacturing technologies will be required and developed. Concurrently, the potential of the existing materials, manufacturing processes and surface modification technologies is not yet fully optimized. In particular advanced laser technology and surface patterning and modification technologies (nano technologies) are still in development. We expect an incremental improvement of stents by manufacturing technologies rather than drastic changes in the next few years. To achieve this, improved interaction between process engineers, materials and surface scientist and clinical experts is required. 11

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