Patent Publication Number: US-7214240-B2

Title: Split-bridge stent design

Description:
BACKGROUND OF THE INVENTION 
   1. Field of the Invention 
   The present invention relates to stents having a modified bridge design and more particularly to stents having a split-bridge design. In addition the present invention relates to intraluminal devices, and more particularly to intraluminal devices, such as stents, incorporating integral markers for increasing the radiopacity thereof. 
   2. Discussion of Related Art 
   Percutaneous transluminal angioplasty (PTA) is a therapeutic medical procedure used to increase blood flow through an artery. In this procedure, the angioplasty balloon is inflated within the stenosed vessel, or body passageway, in order to shear and disrupt the wall components of the vessel to obtain an enlarged lumen. With respect to arterial stenosed lesions, the relatively incompressible plaque remains unaltered, while the more elastic medial and adventitial layers of the body passageway stretch around the plaque. This process produces dissection, or a splitting and tearing, of the body passageway wall layers, wherein the intima, or internal surface of the artery or body passageway, suffers fissuring. This dissection forms a “flap” of underlying tissue, which may reduce the blood flow through the lumen, or completely block the lumen. Typically, the distending intraluminal pressure within the body passageway can hold the disrupted layer, or flap, in place. If the intimal flap created by the balloon dilation procedure is not maintained in place against the expanded intima, the intimal flap can fold down into the lumen and close off the lumen, or may even become detached and enter the body passageway. When the intimal flap closes off the body passageway, immediate surgery is necessary to correct the problem. 
   Recently, transluminal prostheses have been widely used in the medical arts for implantation in blood vessels, biliary ducts, or other similar organs of the living body. These prostheses are commonly referred to as stents and are used to maintain, open, or dilate tubular structures. An example of a commonly used stent is given in U.S. Pat. No. 4,733,665 to Palmaz. Such stents are often referred to as balloon expandable stents. Typically the stent is made from a solid tube of stainless steel. Thereafter, a series of cuts are made in the wall of the stent. The stent has a first smaller diameter, which permits the stent to be delivered through the human vasculature by being crimped onto a balloon catheter. The stent also has a second, expanded diameter, upon application of a radially, outwardly directed force, by the balloon catheter, from the interior of the tubular shaped member. 
   However, one concern with such stents is that they are often impractical for use in some vessels such as the carotid artery. The carotid artery is easily accessible from the exterior of the human body, and is close to the surface of the skin. A patient having a balloon expandable stent made from stainless steel or the like, placed in their carotid artery, and might be susceptible to severe injury through day-to-day activity. A sufficient force placed on the patient&#39;s neck could cause the stent to collapse, resulting in injury to the patient. In order to prevent this, self-expanding stents have been proposed for use in such vessels. Self-expanding stents act like springs and will recover to their expanded or implanted configuration after being crushed. 
   One type of self-expanding stent is disclosed in U.S. Pat. No. 4,655,771. The stent disclosed in U.S. Pat. No. 4,655,771 has a radially and axially flexible, elastic tubular body with a predetermined diameter that is variable under axial movement of the ends of the body relative to each other and which is composed of a plurality of individually rigid but flexible and elastic thread elements defining a radially self-expanding helix. This type of stent is known in the art as a “braided stent” and is so designated herein. Placement of such stents in a body vessel can be achieved by a device, which comprises an outer catheter for holding the stent at its distal end, and an inner piston, which pushes the stent forward once it is in position. 
   However, braided stents have many disadvantages. They typically do not have the necessary radial strength to effectively hold open a diseased vessel. In addition, the plurality of wires or fibers used to make such stents could become dangerous if separated from the body of the stent, where they could pierce through the vessel. Therefore, there has been a desire to have a self-expanding stent, which is cut from a tube of metal, which is the common manufacturing method for many commercially available balloon-expandable stents. In order to manufacture a self-expanding stent cut from a tube, the alloy used would preferably exhibit superelastic or psuedoelastic characteristics at body temperature, so that it is crush recoverable. 
   The prior art makes reference to the use of alloys such as Nitinol (Ni—Ti alloy), which have shape memory and/or superelastic characteristics, in medical devices, which are designed to be inserted into a patient&#39;s body. The shape memory characteristics allow the devices to be deformed to facilitate their insertion into a body lumen or cavity and then be heated within the body so that the device returns to its original shape. Superelastic characteristics, on the other hand, generally allow the metal to be deformed and restrained in the deformed condition to facilitate the insertion of the medical device containing the metal into a patient&#39;s body, with such deformation causing the phase transformation. Once within the body lumen, the restraint on the superelastic member can be removed, thereby reducing the stress therein so that the superelastic member can return to its original un-deformed shape by the transformation back to the original phase. 
   Alloys having shape memory/superelastic characteristics generally have at least two phases. These phases are a martensite phase, which has a relatively low tensile strength and which is stable at relatively low temperatures, and an austenite phase, which has a relatively high tensile strength and which is stable at temperatures higher than the martensite phase. 
   Shape memory characteristics are imparted to the alloy by heating the metal at a temperature above which the transformation from the martensite phase to the austenite phase is complete, i.e. a temperature above which the austenite phase is stable (the Af temperature). The shape of the metal during this heat treatment is the shape “remembered.” The heat-treated metal is cooled to a temperature at which the martensite phase is stable, causing the austenite phase to transform to the martensite phase. The metal in the martensite phase is then plastically deformed, e.g. to facilitate the entry thereof into a patient&#39;s body. Subsequent heating of the deformed martensite phase to a temperature above the martensite to austenite transformation temperature causes the deformed martensite phase to transform to the austenite phase, and during this phase transformation the metal reverts back to its original shape if unrestrained. If restrained, the metal will remain martensitic until the restraint is removed. 
   Methods of using the shape memory characteristics of these alloys in medical devices intended to be placed within a patient&#39;s body present operational difficulties. For example, with shape memory alloys having a stable martensite temperature below body temperature, it is frequently difficult to maintain the temperature of the medical device containing such an alloy sufficiently below body temperature to prevent the transformation of the martensite phase to the austenite phase when the device was being inserted into a patient&#39;s body. With intravascular devices formed of shape memory alloys having martensite-to-austenite transformation temperatures well above body temperature, the devices can be introduced into a patient&#39;s body with little or no problem, but they must be heated to the martensite-to-austenite transformation temperature that is frequently high enough to cause tissue damage. 
   When stress is applied to a specimen of a metal such as Nitinol exhibiting superelastic characteristics at a temperature above which the austenite is stable (i.e. the temperature at which the transformation of martensite phase to the austenite phase is complete), the specimen deforms elastically until it reaches a particular stress level where the alloy then undergoes a stress-induced phase transformation from the austenite phase to the martensite phase. As the phase transformation proceeds, the alloy undergoes significant increases in strain but with little or no corresponding increases in stress. The strain increases while the stress remains essentially constant until the transformation of the austenite phase to the martensite phase is complete. Thereafter, further increases in stress are necessary to cause further deformation. The martensitic metal first deforms elastically upon the application of additional stress and then plastically with permanent residual deformation. 
   If the load on the specimen is removed before any permanent deformation has occurred, the martensitic specimen will elastically recover and transform back to the austenite phase. The reduction in stress first causes a decrease in strain. As stress reduction reaches the level at which the martensite phase transforms back into the austenite phase, the stress level in the specimen will remain essentially constant (but substantially less than the constant stress level at which the austenite transforms to the martensite) until the transformation back to the austenite phase is complete, i.e. there is significant recovery in strain with only negligible corresponding stress reduction. After the transformation back to austenite is complete, further stress reduction results in elastic strain reduction. This ability to incur significant strain at relatively constant stress upon the application of a load, and to recover from the deformation upon the removal of the load, is commonly referred to as superelasticity or pseudoelasticity. It is this property of the material that makes it useful in manufacturing tube cut self-expanding stents. 
   A concern associated with self-expanding stents is that of the compressive forces associated with stent loading and stent deployment. In stent designs having periodically positioned bridges, the resulting gaps between unconnected loops may be disadvantageous, especially during loading into a stent delivery system and subsequent deployment from a stent delivery system. In both the loading and deployment situations, the stent is constrained to a small diameter and subjected to high compressive axial forces. These forces are transmitted axially through the stent by the connecting bridges and may cause undesirable buckling or compression of the adjacent hoops in the areas where the loops are not connected by bridges. 
   One additional concern with stents and with other medical devices formed from superelastic materials, is that they may exhibit reduced radiopacity under X-ray fluoroscopy. To overcome this problem, it is common practice to attach markers, made from highly radiopaque materials, to the stent, or to use radiopaque materials in plating or coating processes. Those materials typically include gold, platinum, or tantalum. The prior art makes reference to these markers or processes in U.S. Pat. No. 5,632,771 to Boatman et al., U.S. Pat. No. 6,022,374 to lmran, U.S. Pat. No. 5,741,327 to Frantzen, U.S. Pat. No. 5,725,572 to Lam et al., and U.S. Pat. No. 5,800,526 to Anderson et al. However, due to the size of the markers and the relative position of the materials forming the markers in the galvanic series versus the position of the base metal of the stent in the galvanic series, there is a certain challenge to overcome; namely, that of galvanic corrosion. Also, the size of the markers increases the overall profile of the stent. In addition, typical markers are not integral to the stent and thus may interfere with the overall performance of the stent as well as become dislodged from the stent. Also, typical markers are used to indicate relative position within the lumen and not whether the device is in the deployed or undepolyed position. 
   SUMMARY OF THE INVENTION 
   The present invention overcomes the disadvantages associated with undesirable loading effects during stent loading and stent deployment as briefly discussed above. The present invention also overcomes many of the disadvantages associated with reduced radiopacity exhibited by self-expanding stents, balloon-expandable stents, and other medical devices as briefly discussed above. 
   In accordance with one aspect, the present invention is directed to an intraluminal medical device. The intraluminal medical device comprises a plurality of hoops forming a substantially tubular member having front and back open ends, one or more bridges interconnecting the plurality of hoops at predetermined positions to form the substantially tubular member, and at least one split-bridge positioned between each of the plurality of hoops, the at least one split-bridge including first and second independent sections which make abutting contact when the intraluminal medical device is constrained and under compressive axial loading. Each of the plurality of hoops comprises a plurality of struts and a plurality of loops connecting adjacent struts. 
   Stent structures are often constructed of radially expanding members or hoops connected by bridge elements. In certain stent designs, the bridge elements may connect every tip or loop of the radially expanding members or hoops to a corresponding tip or loop of an adjacent radially expanding member or hoop. This type of design provides for a less flexible stent. In other stent designs, the bridge elements do not connect every set of tips or loops, but rather, the bridges are placed periodically. When bridges are periodically spaced, open gaps may exist between unconnected tips or loops. This design affords increased flexibility, however, potential deformation of the unconnected tips or loops may occur when the stent is subject to compressive axial loading, for example, during loading of the stent into the stent delivery system or during deployment of the stent. The split-bridge design of the present invention may be utilized to effectively fill the gap between adjacent unconnected tips or loops without serving as a structural connection point between such tips or loops. Accordingly, there is no sacrifice in terms of flexibility. 
   In addition, the split-bridge design serves to increase the surface area of the stent. This increased surface area may be utilized to modify a drug release profile by increasing the amount of drug available for drug delivery. Essentially, increased surface area on the stent allows for more drugs coating thereon. 
   The intraluminal medical device of the present invention may utilize high radiopacity markers to ensure proper positioning of the device within a lumen. The markers comprise a housing that is integral to the device itself, thereby ensuring minimal interference with deployment and operation of the device. The housings are also shaped to minimally impact the overall profile of the stent. For example, a properly shaped housing allows a stent to maintain a radiopaque stent marker size utilized in a seven French delivery system to fit into a six French delivery system. The markers also comprise a properly sized marker insert having a higher radiopacity than the material forming the device itself. The marker insert is sized to match the curvature of the housing thereby ensuring a tight and unobtrusive fit. The marker inserts are made from a material close in the galvanic series to the device material and sized to substantially minimize the effect of galvanic corrosion. 
   The improved radiopacity intraluminal medical device of the present invention provides for more precise placement and post-procedural visualization in a lumen by increasing the radiopacity of the device under X-ray fluoroscopy. Given that the marker housings are integral to the device, they are simpler and less expensive to manufacture than markers that have to be attached in a separate process. 
   The improved radiopacity intraluminal medical device of the present invention is manufactured utilizing a process, which ensures that the marker insert is securely positioned within the marker housing. The marker housing is laser cut from the same tube and is integral to the device. As a result of the laser cutting process, the hole in the marker housing is conical in the radial direction with the outer surface diameter being larger than the inner surface diameter. The conical tapering effect in the marker housing is beneficial in providing an interference fit between the marker insert and the marker housing to prevent the marker insert from being dislodged once the device is deployed. The marker inserts are loaded into a crimped device by punching a disk from annealed ribbon stock-and shaping it to have the same radius of curvature as the marker housing. Once the disk is loaded into the marker housing, a coining process is used to properly seat the marker below the surface of the housing. The coining punch is also shaped to maintain the same radius of curvature as the marker housing. The coining process deforms the marker housing material to form a protrusion, thereby locking in the insert or disk. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
     The foregoing and other aspects of the present invention will best be appreciated with reference to the detailed description of the invention in conjunction with the accompanying drawings, wherein: 
       FIG. 1  is a perspective view of an exemplary stent in its compressed state, which may be utilized in conjunction with the present invention. 
       FIG. 2  is a sectional, flat view of the stent shown in  FIG. 1 . 
       FIG. 3  is a perspective view of the stent shown in  FIG. 1  but showing it in its expanded state. 
       FIG. 4  is an enlarged sectional view of the stent shown in  FIG. 3 . 
       FIG. 5  is an enlarged view of section of the stent shown in  FIG. 2 . 
       FIG. 6  is a view similar to that of  FIG. 2  but showing an alternate embodiment of the stent. 
       FIG. 7  is a perspective view of the stent of  FIG. 1  having a plurality of markers attached to the ends thereof in accordance with the present invention. 
       FIG. 8  is a cross-sectional view of a marker in accordance with the present invention. 
       FIG. 9  is an enlarged perspective view of an end of the stent with the markers forming a substantially straight line in accordance with the present invention. 
       FIG. 10  is a simplified partial cross-sectional view of a stent delivery apparatus having a stent loaded therein, which can be used with a stent made in accordance with the present invention. 
       FIG. 11  is a view similar to that of  FIG. 10  but showing an enlarged view of the distal end of the apparatus. 
       FIG. 12  is a perspective view of an end of the stent with the markers in a partially expanded form as it emerges from the delivery apparatus in accordance with the present invention. 
       FIG. 13  is an enlarged perspective view of an end of the stent with modified markers in accordance with an alternate exemplary embodiment of the present invention. 
       FIG. 14  is an enlarged perspective view of an end of the stent with modified markers in accordance with another alternate exemplary embodiment of the present invention. 
       FIG. 15  is a sectional, flat view of an exemplary embodiment of a split-bridge stent in accordance with the present invention. 
       FIG. 16  is a perspective view of the stent illustrated in  FIG. 15 , but showing the stent in the expanded state. 
   

   DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
   While the present invention may be used on or in connection with any number of medical devices, including stents, for ease of explanation, one exemplary embodiment of the invention with respect to self-expanding Nitinol stents will be described in detail. There is illustrated in  FIGS. 1 and 2 , a stent  100 , which may be utilized in connection with the present invention.  FIGS. 1 and 2  illustrate the exemplary stent  100  in its unexpanded or compressed state. The stent  100  is preferably made from a superelastic alloy such as Nitinol. Most preferably, the stent  100  is made from an alloy comprising from about 50.0 percent (as used herein these percentages refer to weight percentages) Ni to about 60 percent Ni, and more preferably about 55.8 percent Ni, with the remainder of the alloy being Ti. Preferably, the stent  100  is designed such that it is superelastic at body temperature, and preferably has an Af in the range from about twenty-four degrees C to about thirty-seven degrees C. The superelastic design of the stent  100  makes it crush recoverable, which, as discussed above, makes it useful as a stent or frame for any number of vascular devices in different applications. 
   Stent  100  is a tubular member having front and back open ends  102  and  104  and a longitudinal axis  106  extending therebetween. The tubular member has a first smaller diameter,  FIGS. 1 and 2 , for insertion into a patient and navigation through the vessels, and a second larger diameter,  FIGS. 3 and 4 , for deployment into the target area of a vessel. The tubular member is made from a plurality of adjacent hoops  108 ,  FIG. 1  showing hoops  108 ( a )– 108 ( d ), extending between the front and back ends  102  and  104 . The hoops  108  include a plurality of longitudinal struts  110  and a plurality of loops  112  connecting adjacent struts, wherein adjacent struts are connected at opposite ends so as to form a substantially S or Z shape pattern. The loops  112  are curved, substantially semi-circular with symmetrical sections about their centers  114 . 
   Stent  100  further includes a plurality of bridges  116  which connect adjacent hoops  108  and which can best be described in detail by referring to  FIG. 5 . Each bridge  116  has two ends  118  and  120 . The bridges  116  have one end attached to one strut and/or loop, and another end attached to a strut and/or loop on an adjacent hoop. The bridges  116  connect adjacent struts together at bridge to loop connection points  122  and  124 . For example, bridge end  118  is connected to loop  114 ( a ) at bridge to loop connection point  122 , and bridge end  120  is connected to loop  114 ( b ) at bridge to loop connection point  124 . Each bridge to loop connection point has a center  126 . The bridge to loop connection points are separated angularly with respect to the longitudinal axis. That is, the connection points are not immediately opposite each other. Essentially, one could not draw a straight line between the connection points, wherein such line would be parallel to the longitudinal axis of the stent. 
   The above-described geometry helps to better distribute strain throughout the stent, prevents metal-to-metal contact when the stent is bent, and minimizes the opening size between the struts, loops and bridges. The number of and nature of the design of the struts, loops and bridges are important factors when determining the working properties and fatigue life properties of the stent. It was previously thought that in order to improve the rigidity of the stent that struts should be large, and therefore there should be fewer struts per hoop. However, it has now been discovered that stents having smaller struts and more struts per hoop actually improve the construction of the stent and provide greater rigidity. Preferably, each hoop has between twenty-four to thirty-six or more struts. It has been determined that a stent having a ratio of number of struts per hoop to strut length L (in inches) which is greater than four hundred has increased rigidity over prior art stents, which typically have a ratio of under two hundred. The length of a strut is measured in its compressed state parallel to the longitudinal axis  106  of the stent  100  as illustrated in  FIG. 1 . 
   As seen from a comparison of  FIGS. 2 and 3 , the geometry of the stent  100  changes quite significantly as the stent  100  is deployed from its unexpanded state to its expanded state. As a stent undergoes diametric change, the strut angle and strain levels in the loops and bridges are affected. Preferably, all of the stent features will strain in a predictable manner so that the stent is reliable and uniform in strength. In addition, it is preferable to minimize the maximum strain experienced by struts loops and bridges, since Nitinol properties are more generally limited by strain rather than by stress. As will be discussed in greater detail below, the stent sits in the delivery system in its un-expanded state as shown in  FIGS. 10 and 11 . As the stent is deployed, it is allowed to expand towards its expanded state, as shown in  FIG. 3 , which preferably has a diameter which is the same or larger than the diameter of the target vessel. Nitinol stents made from wire deploy in much the same manner, and are dependent upon the same design constraints, as laser cut stents. Stainless steel stents deploy similarly in terms of geometric changes as they are assisted by forces from balloons or other devices. 
   In trying to minimize the maximum strain experienced by features of the stent, the present invention utilizes structural geometries, which distribute strain to areas of the stent, which are less susceptible to failure than others. For example, one of the most vulnerable areas of the stent is the inside radius of the connecting loops. The connecting loops undergo the most deformation of all the stent features. The inside radius of the loop would normally be the area with the highest level of strain on the stent. This area is also critical in that it is usually the smallest radius on the stent. Stress concentrations are generally controlled or minimized by maintaining the largest radii possible. Similarly, we want to minimize local strain concentrations on the bridge and bridge connection points. One way to accomplish this is to utilize the largest possible radii while maintaining feature widths that are consistent with applied forces. Another consideration is to minimize the maximum open area of the stent. Efficient utilization of the original tube from which the stent is cut increases stent strength and its ability to trap embolic material. 
   Many of these design objectives have been accomplished by an exemplary embodiment of the present invention, illustrated in  FIGS. 1 ,  2  and  5 . As seen from these figures, the most compact designs that maintain the largest radii at the loop to bridge connections are non-symmetric with respect to the centerline of the strut connecting loop. That is, loop to bridge connection point centers  126  are offset from the center  114  of the loops  112  to which they are attached. This feature is particularly advantageous for stents having large expansion ratios, which in turn requires them to have extreme bending requirements where large elastic strains are required. Nitinol can withstand extremely large amounts of elastic strain deformation, so the above features are well suited to stents made from this alloy. This feature allows for maximum utilization of Ni—Ti or other material properties to enhance radial strength, to improve stent strength uniformity, to improve fatigue life by minimizing local strain levels, to allow for smaller open areas, which enhance entrapment of embolic material, and to improve stent apposition in irregular vessel wall shapes and curves. 
   As seen in  FIG. 5 , stent  100  comprises strut-connecting loops  112  having a width W 1 , as measured at the center  114  parallel to axis  106 , which are greater than the strut widths W 2 , as measured perpendicular to axis  106  itself. In fact, it is preferable that the thickness of the loops vary so that they are thickest near their centers. This increases strain deformation at the strut and reduces the maximum strain levels at the extreme radii of the loop. This reduces the risk of stent failure and allows one to maximize radial strength properties. This feature is particularly advantageous for stents having large expansion ratios, which in turn requires them to have extreme bending requirements where large elastic strains are required. Nitinol can withstand extremely large amounts of elastic strain deformation, so the above features are well suited to stents made from this alloy. As stated above, this feature allows for maximum utilization of Ni—Ti or other material properties to enhance radial strength, to improve stent strength uniformity, to improve fatigue life by minimizing local strain levels, to allow for smaller open areas, which enhance entrapment of embolic material, and to improve stent apposition in irregular vessel wall shapes and curves. 
   As mentioned above, bridge geometry changes as a stent is deployed from its compressed state to its expanded state and vise-versa. As a stent undergoes diametric change, strut angle and loop strain is affected. Since the bridges are connected to either the loops, struts or both, they are affected. Twisting of one end of the stent with respect to the other, while loaded in the stent delivery system, should be avoided. Local torque delivered to the bridge ends displaces the bridge geometry. If the bridge design is duplicated around the stent perimeter, this displacement causes rotational shifting of the two loops being connected by the bridges. If the bridge design is duplicated throughout the stent, as in the present invention, this shift will occur down the length of the stent. This is a cumulative effect as one considers rotation of one end with respect to the other upon deployment. A stent delivery system, such as the one described below, will deploy the distal end first, and then allow the proximal end to expand. It would be undesirable to allow the distal end to anchor into the vessel wall while holding the stent fixed in rotation, then release the proximal end. This could cause the stent to twist or whip in rotation to equilibrium after it is at least partially deployed within the vessel. Such whipping action may cause damage to the vessel. 
   However, one exemplary embodiment of the present invention, as illustrated in  FIGS. 1 and 2 , reduces the chance of such events happening when deploying the stent. By mirroring the bridge geometry longitudinally down the stent, the rotational shift of the Z-sections or S-sections may be made to alternate and will minimize large rotational changes between any two points on a given stent during deployment or constraint. That is, the bridges  116  connecting loop  108 ( b ) to loop  108 ( c ) are angled upwardly from left to right, while the bridges connecting loop  108 ( c ) to loop  108 ( d ) are angled downwardly from left to right. This alternating pattern is repeated down the length of the stent  100 . This alternating pattern of bridge slopes improves the torsional characteristics of the stent so as to minimize any twisting or rotation of the stent with respect to any two hoops. This alternating bridge slope is particularly beneficial if the stent starts to twist in vivo. As the stent twists, the diameter of the stent will change. Alternating bridge slopes tend to minimize this effect. The diameter of a stent having bridges that are all sloped in the same direction will tend to grow if twisted in one direction and shrink if twisted in the other direction. With alternating bridge slopes this effect is minimized and localized. 
   Preferably, stents are laser cut from small diameter tubing. For prior art stents, this manufacturing process led to designs with geometric features, such as struts, loops and bridges, having axial widths W 2 , W 1  and W 3 , respectively, which are larger than the tube wall thickness T (illustrated in  FIG. 3 ). When the stent is compressed, most of the bending occurs in the plane that is created if one were to cut longitudinally down the stent and flatten it out. However, for the individual bridges, loops and struts, which have widths greater than their thickness, there is a greater resistance to this in-plane bending than to out-of-plane bending. Because of this, the bridges and struts tend to twist, so that the stent as a whole may bend more easily. This twisting is a buckling condition that is unpredictable and can cause potentially high strain. 
   However, this problem has been solved in an exemplary embodiment of the present invention, as illustrated in  FIGS. 1–5 . As seen from these figures, the widths of the struts, hoops and bridges are equal to or less than the wall thickness of the tube. Therefore, substantially all bending and, therefore, all strains are “out-of-plane.” This minimizes twisting of the stent, which minimizes or eliminates buckling and unpredictable strain conditions. This feature is particularly advantageous for stents having large expansion ratios, which in turn requires them to have extreme bending requirements where large elastic strains are required. Nitinol, as stated above, can withstand extremely large amounts of elastic strain deformation, so the above features are well suited to stents made from this alloy. This feature allows for maximum utilization of Ni—Ti or other material properties to enhance radial strength, to improve stent strength uniformity, to improve fatigue life by minimizing local strain levels, to allow for smaller open areas that enhance entrapment of embolic material, and to improve stent apposition in irregular vessel wall shapes and curves. 
   An alternate exemplary embodiment of a stent that may be utilized in conjunction with the present invention is illustrated in  FIG. 6 .  FIG. 6  shows stent  200 , which is similar to stent  100  illustrated in  FIGS. 1–5 . Stent  200  is made from a plurality of adjacent hoops  202 ,  FIG. 6  showing hoops  202 ( a )– 202 ( d ). The hoops  202  include a plurality of longitudinal struts  204  and a plurality of loops  206  connecting adjacent struts, wherein adjacent struts are connected at opposite ends so as to form a substantially S or Z shape pattern. Stent  200  further includes a plurality of bridges  208  which connect adjacent hoops  202 . As seen from the figure, bridges  208  are non-linear and curve between adjacent hoops. Having curved bridges allows the bridges to curve around the loops and struts so that the hoops can be placed closer together which in turn, minimizes the maximum open area of the stent and increases its radial strength as well. This can best be explained by referring to  FIG. 4 . The above described stent geometry attempts to minimize the largest circle which could be inscribed between the bridges, loops and struts, when the stent is expanded. Minimizing the size of this theoretical circle greatly improves the stent because it is then better suited to provide consistent scaffolding support to the vessel and trap embolic material once it is inserted into the patient. 
   As mentioned above, it is preferred that the stent of the present invention be made from a superelastic alloy and most preferably made of an alloy material having greater than 50.5 atomic percentage Nickel and the balance Titanium. Greater than 50.5 atomic percentage Nickel allows for an alloy in which the temperature at which the martensite phase transforms completely to the austenite phase (the Af temperature) is below human body temperature, and preferably is about twenty-four degrees C to about thirty-seven degrees C, so that austenite is the only stable phase at body temperature. 
   In manufacturing the Nitinol stent, the material is first in the form of a tube. Nitinol tubing is commercially available from a number of suppliers including Nitinol Devices and Components, Fremont Calif. The tubular member is then loaded into a machine that will cut the predetermined pattern of the stent into the tube, as discussed above and as shown in the figures. Machines for cutting patterns in tubular devices to make stents or the like are well known to those of ordinary skill in the art and are commercially available. Such machines typically hold the metal tube between the open ends while a cutting laser, preferably under microprocessor control, cuts the pattern. The pattern dimensions and styles, laser positioning requirements, and other information are programmed into a microprocessor, which controls all aspects of the process. After the stent pattern is cut, the stent is treated and polished using any number of methods or combination of methods well known to those skilled in the art. Lastly, the stent is then cooled until it is completely martensitic, crimped down to its un-expanded diameter and then loaded into the sheath of the delivery apparatus. 
     FIG. 15  illustrates an alternate exemplary embodiment of a self-expanding stent  1500  formed from Nitinol. In this exemplary embodiment, a plurality of split-bridges may be utilized to fill the gap between unbridged loops without serving as a structural connection point between these loops. In stent designs featuring periodically placed bridges, as is described herein, the resulting gaps between unconnected loops may be disadvantageous, especially during loading of the stent into a stent delivery system and subsequent deployment from a stent delivery system. In both the loading and deployment situations, the stent is constrained to a small diameter and subjected to high compressive axial forces. These forces are transmitted axially through the stent by the connecting bridges and may cause undesirable buckling or compression of the adjacent hoops in the areas where the loops are not connected by bridges. A split-bridge may be utilized to substantially minimize this undesirable deformation under conditions of constrained axial compression. Essentially, when a stent, having split-bridges is constrained and subjected to axial compression, the wide flat surfaces of adjacent ends of the split-bridge, as is explained in detail subsequently, quickly come into contact and transmit compressive axial loads without allowing undesirable deformation of the stent structure. The split-bridge design is particularly advantageous in that it allows the transmission of the compressive axial loads during stent loading and deployment without the loss of flexibility caused by standard bridges once the stent is deployed. 
   A simple example may be utilized to illustrate the usefulness of a stent comprising split-bridges. A constrained stent which comprises three bridges, typically spaced one hundred twenty degrees apart, must transmit the entire compressive load associated with stent loading and deployment through these three bridges. Unconnected loops within the one hundred twenty degree arc or span between bridges may be undesirably deformed, potentially out of plane, thereby allowing compression of the entire stent structure and potentially adversely impacting loading or deployment characteristics. However, at stent with three standard bridges and three split-bridges would better distribute the axial compressive load, with half the load at or on each bridge, now spaced apart by sixty degrees. In this scenario, there are fewer unconnected loops within the sixty-degree arc or span and these loops would be less likely to become undesirably deformed when the structure is subject to compressive axial loads. Essentially, by allowing efficient transmission of compressive axial loads, the split-bridge helps to prevent undesirable compression or deformation of the constrained stent and loading or deployment difficulties, which may result from such compression or deformation. This may facilitate loading and delivery of stent designs that might otherwise be impractical. 
   It is important to note that symmetric loading and hence symmetric placement of the bridges is preferable but not necessary. 
   Although the split-bridge design may be utilized in any number of stent designs, for ease of explanation, the split-bridge design is described with respect to the exemplary stent illustrated in  FIG. 15 . As illustrated, the stent  1500  comprises a plurality of adjacent hoops  1502 . The hoops  1502  include a plurality of longitudinal struts  1504  and a plurality of loops  1506  connecting adjacent struts  1504 , wherein adjacent struts  1504  are connected at opposite ends so as to form a substantially S or Z shape pattern. The loops  1506  are curved, substantially semi-circular with symmetrical sections about their centers  1508 . The stent  1500  further comprises a plurality of bridges  1510  which connect adjacent hoops  1502 . The bridges  1510  are equivalent to the bridges illustrated in  FIG. 5  and described above. Also as described above, the bridge orientation is changed from hoop to hoop so as to minimize rotational changes between any two points on a given stent during stent deployment or constraint. The number of and nature of the design of the struts, loops and bridges are important factors when determining the working properties and fatigue life properties of the stent as is discussed above. 
   The stent  1500  also comprises a plurality of split-bridges  1512 . The split-bridges  1512  may comprise any suitable configuration and may be positioned in any suitable pattern between bridges  1510 . In the exemplary embodiment illustrated in  FIG. 15 , the split-bridges  1512  are orientated in a direction opposite from that of the bridges  1510  such that a symmetric configuration of bridges  1510  and split-bridges  1512  results. As stated above, a symmetric configuration is not required, but it is preferred. Unlike the bridge  1510  design, the split-bridges  1512  are designed to maximize the surface area for contact between adjacent sections of the split-bridges  1512 . This split-bridge design allows for contact between adjacent sections and thus force transmission even if the adjacent hoops  1502  become somewhat misaligned. The width or thickness of the split-bridges  1512  is preferably larger than the width of the standard bridges  1510  to provide additional surface area for abutting contact. Similarly to bridges  1510 , the split-bridges  1512  have one end of one independent section attached to the one loop  1506  and another end of one independent section attached to a loop  1506  on an adjacent hoop  1502 . Essentially, each split-bridge  1512  comprises first and second independent sections that come into contact when the stent  1500  is under compressive axial loading and make no contact when the stent  1500  is deployed. 
   The geometry of the split-bridge may take any number of forms which serves the purpose of filling the gaps which may be unoccupied by standard bridges. In addition, the number and arrangement of split-bridges is virtually unlimited. 
   By its nature, the split-bridge advantageously allows transmission of compressive loads when constrained because the sections of each split-bridge abut at least partially. However, unlike a traditional bridge, it does not transmit tensile or compressive strains when the expanded structure is stretched, compressed or bent. As illustrated in  FIG. 16 , the split-bridges are not aligned once the structure is expanded. As such, the split-bridge may prove advantageous over a traditional bridge in contourability and fatigue durability. 
   Various drugs, agents or compounds may be locally delivered via medical devices such as stents. For example, rapamycin and/or heparin may be delivered by a stent to reduce restenosis, inflammation and coagulation. One potential limiting factor in these stents is the surface area available on the stent for the drugs, agents and/or compounds. Accordingly, in addition to the advantages discussed above, the split-bridge offers additional surface area onto which various drugs, agents and/or compounds may be affixed. 
   As stated in previous sections of this application, markers having a radiopacity greater than that of the superelastic alloys may be utilized to facilitate more precise placement of the stent within the vasculature. In addition, markers may be utilized to determine when and if a stent is fully deployed. For example, by determining the spacing between the markers, one can determine if the deployed stent has achieved its maximum diameter and adjusted accordingly utilizing a tacking process.  FIG. 7  illustrates an exemplary embodiment of the stent  100  illustrated in  FIGS. 1–5  having at least one marker on each end thereof. In a preferred embodiment, a stent having thirty-six struts per hoop can accommodate six markers  800 . Each marker  800  comprises a marker housing  802  and a marker insert  804 . The marker insert  804  may be formed from any suitable biocompatible material having a high radiopacity under X-ray fluoroscopy. In other words, the marker inserts  804  should preferably have a radiopacity higher than that of the material comprising the stent  100 . The addition of the marker housings  802  to the stent necessitates that the lengths of the struts in the last two hoops at each end of the stent  100  be longer than the strut lengths in the body of the stent to increase the fatigue life at the stent ends. The marker housings  802  are preferably cut from the same tube as the stent as briefly described above. Accordingly, the housings  802  are integral to the stent  100 . Having the housings  802  integral to the stent  100  serves to ensure that the markers  800  do not interfere with the operation of the stent 
     FIG. 8  is a cross-sectional view of a marker housing  802 . The housing  802  may be elliptical when observed from the outer surface as illustrated in  FIG. 7 . As a result of the laser cutting process, the hole  806  in the marker housing  802  is conical in the radial direction with the outer surface  808  having a diameter larger than the diameter of the inner surface  810 , as illustrated in  FIG. 8 . The conical tapering in the marker housing  802  is beneficial in providing an interference fit between the marker insert  804  and the marker housing  802  to prevent the marker insert  804  from being dislodged once the stent  100  is deployed. A detailed description of the process of locking the marker insert  804  into the marker housing  802  is given below. 
   As set forth above, the marker inserts  804  may be made from any suitable material having a radiopacity higher than the superelastic material forming the stent or other medical device. For example, the marker insert  804  may be formed from niobium, tungsten, gold, platinum or tantalum. In the preferred embodiment, tantalum is utilized because of its closeness to nickel-titanium in the galvanic series and thus would minimize galvanic corrosion. In addition, the surface area ratio of the tantalum marker inserts  804  to the nickel-titanium is optimized to provide the largest possible tantalum marker insert, easy to see, while minimizing the galvanic corrosion potential. For example, it has been determined that up to nine marker inserts  804  having a diameter of 0.010 inches could be placed at the end of the stent  100 ; however, these marker inserts  804  would be less visible under X-ray fluoroscopy. On the other hand, three to four marker inserts  804  having a diameter of 0.025 inches could be accommodated on the stent  100 ; however, galvanic corrosion resistance would be compromised. Accordingly, in the preferred embodiment, six tantalum markers having a diameter of 0.020 inches are utilized on each end of the stent  100  for a total of twelve markers  800 . 
   The tantalum markers  804  may be manufactured and loaded into the housing utilizing a variety of known techniques. In the exemplary embodiment, the tantalum markers  804  are punched out from an annealed ribbon stock and are shaped to have the same curvature. as the radius of the marker housing  802  as illustrated in  FIG. 8 . Once the tantalum marker insert  804  is loaded into the marker housing  802 , a coining process is used to properly seat the marker insert  804  below the surface of the housing  802 . The coining punch is also shaped to maintain the same radius of curvature as the marker housing  802 . As illustrated in  FIG. 8 , the coining process deforms the marker housing  802  material to lock in the marker insert  804 . 
   As stated above, the hole  806  in the marker housing  802  is conical in the radial direction with the outer surface  808  having a diameter larger than the diameter of the inner surface  810  as illustrated in  FIG. 8 . The inside and outside diameters vary depending on the radius of the tube from which the stent is cut. The marker inserts  804 , as stated above, are formed by punching a tantalum disk from annealed ribbon stock and shaping it to have the same radius of curvature as the marker housing  802 . It is important to note that the marker inserts  804 , prior to positioning in the marker housing  804 , have straight edges. In other words, they are not angled to match the hole  806 . The diameter of the marker insert  804  is between the inside and outside diameter of the marker housing  802 . Once the marker insert  804  is loaded into the marker housing  802 , a coining process is used to properly seat the marker insert  804  below the surface of the marker housing  802 . In the preferred embodiment, the thickness of the marker insert  804  is less than or equal to the thickness of the tubing and thus the thickness or height of the hole  806 . Accordingly, by applying the proper pressure during the coining process and using a coining tool that is larger than the marker insert  804 , the marker insert  804  may be seated in the marker housing  802  in such a way that it is locked into position by a radially oriented protrusion  812 . Essentially, the applied pressure, and size and shape of the housing tool forces the marker insert  804  to form the protrusion  812  in the marker housing  802 . The coining tool is also shaped to maintain the same radius of curvature as the marker housing  802 . As illustrated in  FIG. 8 , the protrusion  812  prevents the marker insert  804  from becoming dislodged from the marker housing  802 . 
   It is important to note that the marker inserts  804  are positioned in and locked into the marker housing  802  when the stent  100  is in its unexpanded state. This is due to the fact that it is desirable that the tube&#39;s natural curvature be utilized. If the stent were in its expanded state, the coining process would change the curvature due to the pressure or force exerted by the coining tool. 
   As illustrated in  FIG. 9 , the marker inserts  804  form a substantially solid line that clearly defines the ends of the stent in the stent delivery system when seen under fluoroscopic equipment. As the stent  100  is deployed from the stent delivery system, the markers  800  move away from each other and flower open as the stent  100  expands as illustrated in  FIG. 7 . The change in the marker grouping provides the physician or other health care provider with the ability to determine when the stent  100  has been fully deployed from the stent delivery system. 
   It is important to note that the markers  800  may be positioned at other locations on the stent  100 . 
     FIG. 13  illustrates an alternate exemplary embodiment of a radiopaque marker  900 . In this exemplary embodiment, the marker housing  902  comprises flat sides  914  and  916 . The flat sides  914  and  916  serve a number of functions. Firstly, the flat sides  914  and  916  minimize the overall profile of the stent  100  without reducing the radiopacity of the stent  100  under x-ray fluoroscopy. Essentially, the flat sides  914  and  916  allow the marker housings  902  to fit more closely together when the stent  100  is crimped for delivery. Accordingly, the flat sides  914  and  916  of the marker housing  902  allow for larger stents to utilize high radiopacity markers while also allowing the stent to fit into smaller delivery systems. For example, the flat sides  914  and  916  on radiopaque markers  900  of the size described above (i.e. having appropriately sized markers) allow a stent to maintain a radiopaque stent marker size utilized in a seven French delivery system to fit into a six French delivery system. Secondly, the flat sides  914  and  916  also maximize the nitinol tab to radiopaque marker material ratio, thereby further reducing the effects of any galvanic corrosion as described above. The marker insert  904  and the marker hole  906  are formed of the same materials and have the same shape as described above with respect to  FIGS. 1–12 . The markers  900  are also constructed utilizing the same coining process as described above. 
     FIG. 14  illustrates yet another alternate exemplary embodiment of a radiopaque marker  1000 . This exemplary embodiment offers the same advantages as the above-described embodiment; namely, reduced profile without reduction in radiopacity and a reduction in the effects of galvanic corrosion. In this exemplary embodiment, the radiopaque marker  1000  has substantially the same total area as that of the markers  900 ,  800  described above, but with an oval shape rather than a circular shape or circular shape with flat sides. As illustrated, the marker  1000  comprises a substantially oval shaped marker housing  1002  and a substantially oval shaped marker insert  1004 . Essentially, in this exemplary embodiment, the marker  1000  is longer in the axial direction and shorter in the radial direction to allow a larger stent to fit into a smaller delivery system as described above. Also as in the above-described exemplary embodiment, the nitinol tab to radiopaque marker material ratio is improved. In addition, the substantially oval shape provides for a more constant marker housing  1002  thickness around the marker insert  1004 . Once again, the markers  1000  are constructed from the same materials and is constructed utilizing the same coining process as described above. 
   Any of the markers described herein may be utilized or any of the stent designs illustrated as well as any other stent requiring improved radiopacity. 
   It is believed that many of the advantages of the present invention can be better understood through a brief description of a delivery apparatus for the stent, as shown in  FIGS. 10 and 11 .  FIGS. 10 and 11  show a self-expanding stent delivery apparatus  10  for a stent made in accordance with the present invention. Apparatus  10  comprises inner and outer coaxial tubes. The inner tube is called the shaft  12  and the outer tube is called the sheath  14 . Shaft  12  has proximal and distal ends. The proximal end of the shaft  12  terminates at a luer lock hub  16 . Preferably, shaft  12  has a proximal portion  18 , which is made from a relatively stiff material such as stainless steel, Nitinol, or any other suitable material, and a distal portion  20  which may be made from a polyethylene, polyimide, Pellethane, Pebax, Vestamid, Cristamid, Grillamid or any other suitable material known to those of ordinary skill in the art. The two portions are joined together by any number of means known to those of ordinary skill in the art. The stainless steel proximal end gives the shaft the necessary rigidity or stiffness it needs to effectively push out the stent, while the polymeric distal portion provides the necessary flexibility to navigate tortuous vessels. 
   The distal portion  20  of the shaft  12  has a distal tip  22  attached thereto. The distal tip  22  has a proximal end  24  whose diameter is substantially the same as the outer diameter of the sheath  14 . The distal tip  22  tapers to a smaller diameter from its proximal end to its distal end, wherein the distal end  26  of the distal tip  22  has a diameter smaller than the inner diameter of the sheath  14 . Also attached to the distal portion  20  of shaft  12  is a stop  28 , which is proximal to the distal tip  22 . Stop  28  may be made from any number of materials known in the art, including stainless steel, and is even more preferably made from a highly radiopaque material such as platinum, gold or tantalum. The diameter of stop  28  is substantially the same as the inner diameter of sheath  14 , and would actually make frictional contact with the inner surface of the sheath. Stop  28  helps to push the stent out of the sheath during deployment, and helps keep the stent from migrating proximally into the sheath  14 . 
   A stent bed  30  is defined as being that portion of the shaft between the distal tip  22  and the stop  28 . The stent bed  30  and the stent  100  are coaxial so that the distal portion  20  of shaft  12  comprising the stent bed  30  is located within the lumen of the stent  100 . However, the stent bed  30  does not make any contact with stent  100  itself. Lastly, shaft  12  has a guidewire lumen  32  extending along its length from its proximal end and exiting through its distal tip  22 . This allows the shaft  12  to receive a guidewire much in the same way that an ordinary balloon angioplasty catheter receives a guidewire. Such guidewires are well known in art and help guide catheters and other medical devices through the vasculature of the body. 
   Sheath  14  is preferably a polymeric catheter and has a proximal end terminating at a sheath hub  40 . Sheath  14  also has a distal end, which terminates at the proximal end  24  of distal tip  22  of the shaft  12 , when the stent is in its fully un-deployed position as shown in the figures. The distal end of sheath  14  includes a radiopaque marker band  34  disposed along its outer surface. As will be explained below, the stent is fully deployed from the delivery apparatus when the marker band  34  is lined up with radiopaque stop  28 , thus indicating to the physician that it is now safe to remove the apparatus  10  from the body. Sheath  14  preferably comprises an outer polymeric layer and an inner polymeric layer. Positioned between outer and inner layers is a braided reinforcing layer. Braided reinforcing layer is preferably made from stainless steel. The use of braided reinforcing layers in other types of medical devices can be found in U.S. Pat. No. 3,585,707 issued to Stevens on Jun. 22, 1971, U.S. Pat. No. 5,045,072 issued to Castillo et al. on Sep. 3, 1991, and U.S. Pat. No. 5,254,107 issued to Soltesz on Oct. 19, 1993. 
     FIGS. 10 and 11  illustrate the stent  100  as being in its fully undeployed position. This is the position the stent is in when the apparatus  10  is inserted into the vasculature and its distal end is navigated to a target site. Stent  100  is disposed around stent bed  30  and at the distal end of sheath  14 . The distal tip  22  of the shaft  12  is distal to the distal end of the sheath  14 , and the proximal end of the shaft  12  is proximal to the proximal end of the sheath  14 . The stent  100  is in a compressed state and makes frictional contact with the inner surface  36  of the sheath  14 . 
   When being inserted into a patient, sheath  14  and shaft  12  are locked together at their proximal ends by a Tuohy Borst valve  38 . This prevents any sliding movement between the shaft and sheath, which could result in a premature deployment or partial deployment of the stent  100 . When the stent  100  reaches its target site and is ready for deployment, the Tuohy Borst valve  38  is opened so that that the sheath  14  and shaft  12  are no longer locked together. 
   The method under which the apparatus  10  deploys the stent.  100  is readily apparent. The apparatus  10  is first inserted into the vessel until the radiopaque stent markers  800  (leading  102  and trailing  104  ends, see  FIG. 7 ) are proximal and distal to the target lesion. Once this has occurred the physician would open the Tuohy Borst valve  38 . The physician would then grasp hub  16  of shaft  12  so as to hold it in place. Thereafter, the physician would grasp the proximal end of the sheath  14  and slide it proximal, relative to the shaft  12 . Stop  28  prevents the stent  100  from sliding back with the sheath  14 , so that as the sheath  14  is moved back, the stent  100  is pushed out of the distal end of the sheath  14 . As stent  100  is being deployed the radiopaque stent markers  800  move apart once they come out of the distal end of sheath  14 . Stent deployment is complete when the marker  34  on the outer sheath  14  passes the stop  28  on the inner shaft  12 . The apparatus  10  can now be withdrawn through the stent  100  and removed from the patient. 
     FIG. 12  illustrates the stent  100  in a partially deployed state. As illustrated, as the stent  100  expands from the delivery device  10 , the markers  800  move apart from one another and expand in a flower like manner. 
   Although shown and described is what is believed to be the most practical and preferred embodiments, it is apparent that departures from specific designs and methods described and shown will suggest themselves to those skilled in the art and may be used without departing from the spirit and scope of the invention. The present invention is not restricted to the particular constructions described and illustrated, but should be constructed to cohere with all modifications that may fall within the scope of the appended claims.