Bare-Stent Technology and Its Utilization in the Treatment of Atherosclerotic Obstructive Disease

Abstract

Following the introduction of percutaneous angioplasty (PTA), arterial stents emerged as a tool for preventing the acute arterial wall recoil and chronic restenosis associated with PTA. The use of stents has grown considerably in the last decade. There is a wide spectrum of clinical applications for the use of stents in the arterial vascular system, but every vascular territory demands specific stent adaptations such as resistance to kink and fracture in the femoropopliteal segment or high radial force in aortic ostial lesions. This article discusses the different types of bare-metal stent models and how their characteristics influence endothelialization as well as their potential impact on clinical practice.

Key words: stent, atherosclerosis, endothelium, restenosis

VASCULAR DISEASE MANAGEMENT 2010;7:E95–E102

Introduction

In 1964 Charles Dotter introduced the concept of arterial remodeling and in 1974 Andreas Gruentzig performed the first peripheral human balloon angioplasty. Later, physicians started to recognize the limitations of angioplasty and stents were introduced with the objective of tacking down dissection flaps and providing mechanical support of recoiling lesions. In 1991, the U.S. Food and Drug Administration approved for the first time the use of the Palmaz stent for the treatment of atherosclerotic obstructive disease. Since then, the use of stents in the treatment of atherosclerotic obstructive lesions has widely spread, although significant limitations still remain. Restenosis is one of the most prevalent complications of stenting. Restenosis is the process of luminal narrowing in an atherosclerotic artery after endovascular interventions and has been found to occur in 20–50% of stented vessels.¹ Restenosis is the arterial wall healing response to mechanical injury and has long been attributed to neointimal proliferation,² thrombosis³ and negative remodeling.⁴ However, there is now increasing evidence for the role of inflammation in vascular healing and in the development of restenosis. Vascular injury, de-endothelialization, platelet and leukocyte interactions and expression of autocrine and paracrine inflammatory mediators (cytokines such as interleukin-1 and tumor necrosis factor), which are triggered following stent implantation, orchestrate an inflammatory response that leads to smooth-cell proliferation, potential neointimal growth and subsequent restenosis. Stents may be implanted at sites of advanced macrophage-rich atheromatous lesions. Stent implantation induces a vascular inflammatory response that also contributes to restenosis. Recently, the use of immunosuppressive drug-eluting stents (DES) has provided very promising results in the treatment of restenosis, although they inhibit endothelialization, arresting the healing process of the vessel walls after vascular injury caused by stent implantation, which may translate into adverse clinical events.⁵ Different stent models have distinct rates of in-stent restenosis; the vascular injury caused by stent placement determines the risk and degree of restenosis.⁶ The degree of vascular injury and subsequent restenosis appears to depend on the depth of strut penetration during stent implantation.⁷ This article discusses several bare-metal stent (BMS) models and how stent characteristics affect endothelialization and in-stent restenosis and their implications in clinical practice.

Types of Stents

Stents are implantable devices that can be classified according to design or geometry (mesh structure, coil, slotted tube, modular or custom design), delivery system and mechanism of expansion (self-expanding or balloon-mounted), and composition (stainless steel, cobalt-based alloy, tantalum, nitinol, etc). Additional stent characteristics include strut thickness and shape, metal-to-artery ratio, method of stent cleaning and polishing, corrosion resistance, durability, open area-to-metal surface ratio, sharpness of the end of the stent, fracture resistance, kinkability, biocompatibility and long-term clinical outcomes.

Stent materials. Materials used for the manufacturing of stents must be biocompatible, corrosion-resistant and should have adequate radiopacity. Balloon-expandable stents are made of malleable metals, materials that can be deformed (low-yield stress) to the size of the target vessel by coaxial balloon inflation and remain expanded (high elastic modulus).⁸ 316L stainless steel, which has good radial strength, deformability and resists corrosion, was widely used to build balloon-mounted stents. Materials such as tantalum, platinum alloys, niobium alloys, cobalt-chromium alloys, nitinol and polymers such as polyester are replacing stainless steel due to better radiopacity, higher strength, improved corrosion resistance and magnetic resonance (MR) compatibility. Materials with better radiopacity and strength allow manufacturing of stents with smaller delivery profiles. Self-expandable stents are made of elastic materials (low elastic modulus and high-yield stress). Large elastic strains are achieved by super elasticity or thermal memory of stents.⁹ These stents are manufactured in the expanded shape and then compressed and constrained into the delivery system. Nitinol, a nickel-titanium thermo-elastic alloy, is widely used in these self-expandable stents. The elasticity of some self-expandable stents can also be achieved from the design of the stent mesh, such as the Cook Z stent (stainless steel) (Cook, Inc., Bloomington, Indiana) and the Boston Scientific Wallstent (elgiloy, a cobalt alloy) (Boston Scientific Corp., Natick, Massachusetts). External forces may affect the stent shape depending on the deformability and elasticity of the stent material. Balloon-expandable stents can experience permanent plastic deformation if the external force exceeds its maximum hoop strength. External compression, bending and flexion may result in permanent stainless steel stent collapse and compression. On the other hand, nickel titanium (NiTi) alloy stents are more flexible and resist the permanent deformation associated with bending and flexion due to their shape memory capabilities. Materials used in stent construction have demonstrated different interaction with fibrinogen, platelets and endothelial cells (Figures 1, 2 and 3).^10–13 The activation of platelets and fibrinogen deposition on stent surfaces make them thrombogenic. Thrombogenic stent materials result in less tissue colonization by endothelial cells, therefore affecting endothelialization of the stented segment. Stent geometry and design. Stent geometry and design are defined as the organization or structure of the stent elements. Coil design is conformed by a continuous wire or a series of flat sheet coils. The struts of coil stents are wide, and there are gaps between struts and no connections between them, which give the coil stent design more flexibility at the cost of radial strength (Figure 4). A helical stent consists of one or more integrated helical bands that provide flexibility and enough axial rigidity to maximize lumen integrity and avoid stent kinking (Figure 5). Slotted-tube stents are made by direct cutting from a single metal tube using a laser beam or waterjet cutting. Slotted tube designs have excellent radial force at the cost of flexibility (Figure 6). Modular stent design consists of sequential rings or crowns connected with welded joins or connecting struts to provide flexibility with good conformability and scaffolding (Figure 7). The number and arrangement of bridge connectors differentiate open-cell designs from closed-cell designs. When adjacent ring or crown segments are connected at every possible junction, the design is classified as closed cell, and when some of the connecting junction points are removed, the design is classified as open cell. Open-cell stent design allows more flexion between adjacent rings because fewer connection points allow for greater flexion and conformability. But the improved flexions of open-cell designs have a cost in terms of scaffolding uniformity, while closed-cell designs are less flexible and conformable, but feature better scaffolding (Figures 8 and 9). The connections between rings are divided in non-flex and flex-bridge connectors, allowing some limited degree of flexion between adjacent rings (Figure 10). The stent strut shape and size are other important aspects of stent design. Specifically, in cross-sections, the struts can be classified as round or square and thick or thin. The use of thinner strut devices (strut thicknesses of 50–75 µm) compared to thicker stent struts (strut thicknesses up to 140 µm) has been associated with a significant reduction of angiographic restenosis from 15.0% in the thin-strut group to 25.8% in the thick-strut group after coronary artery stenting14 (Figure 11). Strut edge angles also influence endothelialization rates. Endothelialization is improved in the presence of a strut edge angle of 35 degrees in the direction of blood flow (Figure 12).¹⁵ The smoothness of a stent surface affects the performance and biocompatibility of the stent. Smooth surfaces can reduce platelet adhesion, thrombus formation and neointimal growth and improve endothelialization. The laser-cutting production process leaves slag that includes depositions and burrs on the surface of stents. Pressurizing removes surface contaminants and then acid-pickling and electrochemical polishing are used to achieve smoothness. Coating of the stent surface through a galvanization process with another alloy such as argon ion bombardment can lead to porosity and surface cracks. On the other hand, modification of the stent surface such as creating grooves on the surface can speed endothelialization (Figures 13 and 14).¹⁶ Finally, the extent of metal coverage at the stented lesion site varies widely depending on the stent design. For example, stents with small cells, closed cell devsign and thick struts offer more surface coverage and less lesion prolapse. In terms of lesion coverage, a closed-cell design appears to be less important than the actual cell size. Large free-cell areas are more prone to plaque material prolapse and embolization. Important aspects of delivery systems and modes of stent expansion to consider include trackability, pushability, crossing profile, amount of friction of the delivery system, flexibility of the stent and the delivery balloon, and precise stent delivery and positioning. Trackability and pushability refer to the ease of a stent on its delivery system to negotiate vessel angles and tortuosity. This feature of delivery system depends on the flexibility of the stent, its delivery catheter and balloon-catheter in balloon-expandable stents. Self-expandable stents are mounted on the distal end of the catheter and positioned under a retractable sheath. Once the catheter has been advanced intraluminally to the site where the stent is to be implanted, the sheath is withdrawn, thereby allowing the self-expanding stent to expand radially and outwardly into contact with the arterial wall. In the case of balloon-mounted stents, the stent is crimped on a balloon and then deployed by balloon inflation. Stent retention on the balloon to avoid stent loss, overhang of the delivery balloon from the stent edges to limit vessel wall trauma and the associated risk of peri-stent dissection, and low-compliance balloons to assure homogeneous stent deployment are qualities to consider when balloon-mounted stents are chosen. Balloon-mounted stent delivery systems provide more precise positioning and deployment of the stent than self-expandable delivery systems but tend to track less effectively. Mechanical properties of stents include scaffolding, radial force, hoop strength, flexibility, lesion coverage, conformability, vessel wall apposition, radiopacity and MR imaging compatibility. Scaffolding refers to the support that the stent provides to the vessel wall. Radial force of a stent is the force directed outwardly against the arterial wall, while hoop strength is the ability of a stent to resist radial compression by external forces including the arterial wall’s elasticity. Radial force is a primary factor governing the performance of self-expanding stents. Lesion coverage is similar to metal coverage and defined as the surface area of the vessel wall covered by the summation of the stent’s elements. Flexibility (axial and longitudinal) is the ability of stents to bend. A rigid stent may produce a hinge effect at the end of the stented segment. Fracture resistance is the ability to oppose external forces that result in breaking of one or multiple stent struts, which may result in complete transverse stent fracture with or without stent displacement. The surface finish is the interface between the material of the stent and the local biologic environment where the stent is implanted. Poorly polished stent surfaces with cracks, fissures and retained slag activate an inflammatory response that may result in weakness points. In addition, biomechanical forces within the artery, such as shortening, torque and bending may result in metal fatigue and strut failure and eventual strut fractures. Conformability refers to the stent’s ability to accommodate to the vessel shape. Greater conformability of the stent to the inner lumen geometry of the vessel cause less vascular injury and in-stent restenosis.6 Vessel wall apposition is the ability of the stent to come in contact with the vessel wall and is the end result of stent conformability. Conformability and good wall apposition preserve the original shear stress pattern of the arterial flow, improving endothelialization and reducing the amount of tissue hyperplasia.¹⁴ Fifteen dynes/cm² is the wall shear stress of a non-diseased arterial segment, and modification in shear stress patterns of blood flow affect endothelialization rates (Figures 15 and 16). Thus, a low or disturbed flow pattern at a stented site greatly diminishes the re-endothelialization rate. Radiopacity involves the property of obstructing the passage of X-rays. Nitinol stents are less radiopaque than stainless steel stents. When less radiopaque stents are used, X-ray visibility is improved with radiopaque markers or coating (gold, tantalum and platinum) may be added to the stent. Foreshortening, common in balloon-mounted stents, occurs when the stent length shortens from the unexpanded to the expanded form (Figure 6). On the other hand, nitinol self-expandable stents exhibit minimal foreshortening upon expansion. Finally, MR imaging compatibility is an important issue influencing the choice of stent material. MR imaging and MR angiography are widely use imaging modalities that are often required in patients with stents. Inert non-ferromagnetic materials (nitinol and other non-stainless steel stents) do not experience significant torque when placed in strong magnetic fields and cause significant less artifacts on MR than ferromagnetic materials such as stainless steel. On the other hand, stents made of ferromagnetic materials not only result in greater susceptibility and radiofrequency artifacts, but also increase the risk of stent translation and or torque if imaging is performed soon after placement and before tissue incorporation occurs.¹⁷ Endothelialization of stents prevents platelet adhesion and activation as well as fibrinogen deposit on the stent surface. Thus, the presence of a restored, intact endothelium may reduce stent thrombosis and platelet activation which, by way of growth factor and other cytokine release, may drive a proliferate response. Avoiding excessive collateral intimal injury at the implantation site is important for preservation of the adjacent endothelium needed to ensure device endothelialization. Conformability and wall apposition, which promote high shear flow, are important for endothelialization as well. Stent kinks and poor wall apposition may result in low shear flow, delaying endothelialization and facilitating thrombosis. Moreover, the duration of antiplatelet therapy after stent implantation should be considered depending on the endothelialization time.

Clinical Aspects of Stenting

Following the introduction of the first rigid balloon-mounted stents, a primary concern of stent development was the need to increase flexibility in order to improve conformability, wall apposition and facilitate safe stent delivery. Manufacturers have improved flexibility at the cost of compromising radial support and lesion coverage. Thus, the flexion benefits of an open-cell design have a tradeoff in terms of scaffolding uniformity, just as the scaffolding benefits of a closed-cell design have offer less flexion and conformability. In clinical practice, the implantation of stents is an integral part of most interventional procedures for percutaneous revascularization, and the operator must decide which stent is most appropriate for the patient being treated. Features of stent models make a specific stent more or less suitable for a particular type of lesion or anatomy. In general, the affected artery, lesion characteristics (calcified, soft, ulcerated, eccentric, concentric, fibrous, focal, diffuse, etc.), lesion location, access site, sheath size, vessel angle and tortuosity influence stent selection. For calcified, resistant, fibrous or eccentric lesions prone to elastic recoil, a balloon-expandable stent would be more appropriate. In tortuous and elongated vessels, placement of a rigid stent results in vessel straightening and angulations as well as a kink of the vessel at the ends of the stent. To avoid this, self-expandable stents with good flexibility and vessel conformability should be used. Self-expandable stents are also preferred in arteries that taper or transition abruptly in size because they accommodate better to the shape of the vessel and come into contact with the vessel wall. In carotid stenting, complication rates vary according to stent type and free-cell area, Bosiers reported that transient ischemic attack, stroke and death rates varied from 1.2–3.4% for free-cell areas 7.5 mm, respectively (p 12 These results suggest that surface area coverage is clinically significant in preventing plaque prolapse and embolization. For aorto-ostial lesions, slotted-tube stents are the preferred choice due to their strong radial support, low recoil, precise positioning and deployment and radiologic visibility. The strong elastic recoil inherent to the wall of the aorta favors the use of thicker struts to provide greater resistance when dealing with ostial lesions (Figure 17). The femoropopliteal artery undergoes shortening and bending during leg flexion and extension. Stenting of this mobile artery induces axial rigidity, thereby reducing the artery’s ability to accommodate foreshortening. This lack of accommodation adds stress to the stent and the adjacent unstented artery, which may ultimately contribute to stent kinking or fracturing. This mechanical stress may also delay vessel healing and stent incorporation. Depending on their severity, stent fractures may promote intimal hyperplasia, aneurysms and vascular occlusions. Various studies, including SIROCCO and FESTO suggested that stent fracture rates increase with higher numbers of stents placed and when long segments of contiguous stents overlap. These trials have reported fracture rates as high as 37%.^17,18 On the other hand, the RESILENT trial demonstrated that a novel helical stent design provided patency with no stent fractures at 12-month follow-up (Figure 18).

Conclusions

Stent design, materials and surface preparation have a significant impact on long-term clinical outcomes. In-stent restenosis remains a prevalent problem affecting stent revascularization therapies. Understanding the influence of stent design and materials on the healing process of the arterial wall is critical in order to achieve long-term clinical outcomes. Therefore, novel stent therapies that promote a “controlled wound healing of the arterial wall,” including endothelialization of the stented segment, are needed to overcome the long-term clinical limitations of endovascular treatment of atherosclerosis.

References

1. Stoeckel D, Bonsignore C, Duda S. A survey of stent designs. Minim Invasive Ther Allied Technol 2002;11:137–147.

2. Kastrati A, Mehilli J, Dirschinger J, et al. Intracoronary stenting and angiographic results: Strut thickness effect on restenosis outcome (ISAR-STEREO) trial. Circulation 2001;103:2816–2821.

3. Lally CF, Dolan F, Prendergast PJ. Cardiovascular stent design and vessel stresses: A finite element analysis. J Biomech 2005;38:1574–1581.

4. Sprague EA, Luo J, Palmaz JC. Human aortic endothelial cell migration onto stent surfaces under static and flow conditions. J Vasc Interv Radiol 1997:8(1 Pt 1):83–92.

5. Grewe PH, Deneke T, Machraoui A, et al. Acute and chronic tissue response to coronary stent implantation: Pathologic findings in human specimen. J Am Coll Cardiol 2000;35:157–163.

6. Forrester JS, Fishbein M, Helfant R, Fagin J. A paradigm for restenosis based on cell biology: Clues for the development of new preventive therapies. J Am Coll Cardiol 1991;17:758–769.

7. Libby P, Schwartz D, Brogi E, et al. A cascade model for restenosis. A special case of atherosclerosis progression. Circulation 1992;86(6 Suppl):III47–III52.

8. Gaya M, Liu WK. Stent modeling using immersed finite element method Computer Methods in Applied Mechanics and Engineering 2006;195:4358–4370.

9. Schwartz RS, Topol EJ, Serruys PW, et al. Artery size, neointima, and remodeling: time for some standards. J Am Coll Cardiol 1998;32:2087–2094.

10. Liuzzo, JP, Ambrose JA, Coppola JT. Sirolimus- and taxol-eluting stents differ towards intimal hyperplasia and re-endothelialization. J Invasive Cardiol 2005;17:497–502.

11. Arakawa K, Isoda K, Sugiyabu Y, et al. Intimal proliferation after stenting reflected by increased stent-to-vessel cross-sectional area ratio: Serial intravascular ultrasound study. J Cardiol 1998;32:379–389.

12. Bosiers M, de Donato G, Deloose K, et al. Does free cell area influence the outcome in carotid artery stenting? Eur J Vasc Endovasc Surg 2007;33:135–141; Discussion 142–143.

13. Sprague EA, Palmaz JC. A model system to assess key vascular responses to biomaterials. J Endovasc Ther 2005;12:594–604.

14. Duda SH, Bosiers M, Lammer J, et al. Sirolimus-eluting versus bare nitinol stent for obstructive superficial femoral artery disease: The SIROCCO II trial. J Vasc Interv Radiol 2005;16:331–338.

15. Hamuro M, Palmaz J, Sprague E, et al. Influence of stent edge angle on endothelialization in an in vitro model. J Vasc Interv Radiol 2001;12:607–611.

16. Palmaz JC, Benson A, Sprague EA. Influence of surface topography on endothelialization of intravascular metallic material. J Vasc Interv Radiol 1999;10:439–444.

17. Scheinert D, Scheinert S, Sax J, et al. Prevalence and clinical impact of stent fractures after femoropopliteal stenting. J Am Coll Cardiol 2005;45:312–315.

18. Duda SH, Pusich B, Richter G, et al. Sirolimus-eluting stents for the treatment of obstructive superficial femoral artery disease: Six-month results. Circulation 2002;106:1505–1509.

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From the University of Texas Health Science Center of San Antonio (UTHSCSA), San Antonio, Texas. The authors report no conflicts of interest regarding the content herein. Address for correspondence: Marco Cura, MD, Radiology, Doctors Hospital, 5501 S. McColl Rd. Edinburgh, TX 78229. E-mail: marcocura@yahoo.com