fig2 shows one embodiment of an expandable tissue supporting device 20 in accordance with the present invention . the tissue supporting device 20 includes a series of axial slots 22 formed in a cylindrical tube 24 . each axial slot 22 is displaced axially from the slots in adjacent rows of slots by approximately half the slot length resulting in a staggered slot arrangement . the offset between adjacent rows of slots results in alternate rows of slots which extend to the ends of the cylindrical tube 24 . at each interior end of each of the axial slots 22 a circumferential slot 26 is formed . the material between the slots 22 forms a network of axial struts 28 extending substantially parallel to an axis of the tube 24 . the axial struts 28 are joined by short circumferential links 30 . the circumferential links 30 are positioned at both the interior of the cylindrical tube and at the ends of the cylindrical tube . the cross section ( and rectangular moment of inertia ) of each of the struts 28 is not constant along the length of the strut . rather , the strut cross section changes abruptly at both ends of each strut 28 at the location of the circumferential slots 26 . the struts 28 are thus not prismatic . each individual strut 28 is linked to the rest of the structure through a pair of reduced sections 32 , one at each end , which act as stress / strain concentration features . the reduced sections 32 of the struts function as hinges in the cylindrical structure . since the stress / strain concentration features 32 are designed to operate into the plastic deformation range of generally ductile materials , they are referred to as ductile hinges . such features are also commonly referred to as “ notch hinges ” or “ notch springs ” in ultra - precision mechanism design , where they are used exclusively in the elastic range . with reference to the drawings and the discussion , the width of any feature is defined as its dimension in the circumferential direction of the cylinder . the length of any feature is defined as its dimension in the axial direction of the cylinder . the thickness of any feature is defined as the wall thickness of the cylinder . the presence of the ductile hinges 32 allows all of the remaining features in the tissue supporting device to be increased in width or the circumferentially oriented component of their respective rectangular moments of inertia — thus greatly increasing the strength and rigidity of these features . the net result is that elastic , and then plastic deformation commence and propagate in the ductile hinges 32 before other structural elements of the device undergo any significant elastic deformation . the force required to expand the tissue supporting device 20 becomes a function of the geometry of the ductile hinges 32 , rather than the device structure as a whole , and arbitrarily small expansion forces can be specified by changing hinge geometry for virtually any material wall thickness . in particular , wall thicknesses great enough to be visible on a fluoroscope can be chosen for any material of interest . in order to get minimum recoil , the ductile hinges 32 should be designed to operate well into the plastic range of the material , and relatively high local strain - curvatures are developed . when these conditions apply , elastic curvature is a very small fraction of plastic or total curvature , and thus when expansion forces are relaxed , the percent change in hinge curvature is very small . when incorporated into a strut network designed to take maximum advantage of this effect , the elastic springback , or “ recoil ,” of the overall stent structure is minimized . in the embodiment of fig2 it is desirable to increase the width of the individual struts 28 between the ductile hinges 32 to the maximum width that is geometrically possible for a given diameter and a given number of struts arrayed around that diameter . the only geometric limitation on strut width is the minimum practical width of the slots 22 which is about 0 . 002 inches ( 0 . 0508 mm ) for laser machining . lateral stiffness of the struts 28 increases as the cube of strut width , so that relatively small increases in strut width significantly increase strut stiffness . the net result of inserting ductile hinges 32 and increasing strut width is that the struts 28 no longer act as flexible leaf springs , but act as essentially rigid beams between the ductile hinges . all radial expansion or compression of the cylindrical tissue supporting device 20 is accommodated by mechanical strain in the hinge features 32 , and yield in the hinge commences at very small overall radial expansion or compression . yield in ductile hinges at very low gross radial deflections also provides the superior crimping properties displayed by the ductile hinge - based designs . when a tissue supporting device is crimped onto a folded catheter balloon , very little radial compression of the device is possible since the initial fit between balloon and device is already snug . most stents simply rebound elastically after such compression , resulting in very low clamping forces and the attendant tendency for the stent to slip on the balloon . ductile hinges , however , sustain significant plastic deformation even at the low deflections occurring during crimping onto the balloon , and therefore a device employing ductile hinges displays much higher clamping forces . the ductile hinge designs according to the present invention may be securely crimped onto a balloon of a delivery catheter by hand or by machine without the need for auxiliary retaining devices commonly used to hold known stents in place . the geometric details of the stress / strain concentration features or ductile hinges 32 can be varied greatly to tailor the exact mechanical expansion properties to those required in a specific application . the most obvious and straightforward ductile hinges are formed by slots or notches with rounded roots , as in fig3 a and 3 c . since the laser beams often used to fabricate these features are themselves round , slots or notches with circular roots are also among the easiest to fabricate . fig3 a shows a ductile hinge 36 formed by a pair of opposed circular grooves 38 , 40 . according to this embodiment the circumferential slot 26 has semicircular ends 38 having a radius of curvature . outer semicircular grooves 40 oppose the semicircular ends 38 and also have a radius of curvature which is the same as that of the grooves 38 . fig3 c shows another ductile hinge 54 formed by a parabolic groove 56 . generally , the ductile hinges 36 of the embodiment of fig3 a formed between pairs of concave curves 38 , 40 have a minimum width along a line connecting their respective centers of curvature . when the struts connected by the ductile hinge are moved apart or together , plastic deformation is highly concentrated in a region immediately adjacent to the plane that bisects the hinge at this narrow point . for smaller deflection , this very high strain concentration at the bisecting plane is acceptable , and in some cases , useful . for stent crimping purposes , for example , it is desirable to generate relatively large plastic deformations at very small deflection angles . as a practical matter , however , strut deflection angles for device expansion are often in the 25 ° to 45 ° range . at these angles , strain at the root or bisecting plane of concave ductile hinge features can easily exceed the 50 to 60 % elongation - to - failure of 316l stainless steel , one of the most ductile stent materials . deflection limiting features which will be described further below limit the geometric deflection of struts , but these features do not in themselves affect the propagation pattern of plastic deformation in a given ductile hinge design . for concave ductile hinges at large bend angles , very high strain concentrations remain . scanning electron micrographs have confirmed this analysis . in many engineering applications , it is desirable to limit the amount of strain , or “ cold - work ,” in a material to a specified level in order to optimize material properties and to assure safe operation . for example , in medical applications it is desirable to limit the amount of cold - work in 316l stainless steel to about 30 %. at this level , the strength of the material is increased , while the material strain is still well below the failure range . ideally , therefore , a safe and effective ductile hinge should not simply limit gross deflection , but reliably limit material strain to a specified level . fig3 b shows a simple ductile hinge design that allows material strain to be limited to some specified level . the ductile hinge of fig3 b is formed by a rectangular circumferential groove 46 with filleted corners 48 on one side of a strut , the opposite side 50 of the strut remaining straight . the ductile hinges 44 are substantially rectangular sections between the ends of the groove 46 and the side walls 50 . one of the key concepts in fig3 b is that the ductile hinge 44 has a constant or substantially constant width along at least a portion of its total length . in this configuration , there is no local minimum width along the ductile hinge axis , as there is with pairs of concave roots . there is therefore no point concentration of stresses and strains along the length of the ductile hinge beam during stent expansion . in particular , maximum tensile and compressive strains will be distributed evenly along the upper and lower surfaces of the hinge 44 during stent expansion . with the gross bend angle limited by mechanical stops , which are described below in detail , the maximum material strain ( at the hinge surfaces ) can therefore be reliably limited by adjusting the initial length of the ductile hinge over which the total elongation is distributed . fig3 d shows a ductile hinge 60 in a cylindrical wire 62 for incorporating into a wire - form tissue - supporting device . the ductile hinge 60 is formed by a reduced diameter portion of the wire 62 . again , it is important that the ductile hinge have a substantially constant width over a portion of its length in order to provide strain control . preferably , the ductile hinge is prismatic over a portion of its length . maximum material strain can be varied by adjusting the hinge length . the ductile hinges of the present invention have a constant or substantially constant width over at least ⅓ of the ductile hinge length , and preferably over at least ½ of the ductile hinge length . fig3 e shows an asymmetric ductile hinge 64 that produces different strain versus deflection - angle functions in expansion and compression . each of the ductile hinges 64 is formed between a convex surface 68 and a concave surface 69 . the ductile hinge 64 according to a preferred embodiment essentially takes the form of a small , prismatic curved beam having a substantially constant cross section . however , a thickness of the curved ductile hinge 64 may vary somewhat as long as the ductile hinge width remains constant along a portion of the hinge length . the width of the curved beam is measure along the radius of curvature of the beam . this small curved beam is oriented such that the smaller concave surface 69 is placed in tension in the device crimping direction , while the larger convex surface 68 of the ductile hinges is placed in tension in the device expansion direction . again , there is no local minimum width of the ductile hinge 64 along the ( curved ) ductile hinge axis , and no concentration of material strain . during device expansion tensile strain will be distributed along the convex surface 68 of the hinge 64 and maximum expansion will be limited by the angle of the walls of the concave notch 69 which provide a geometric deflection limiting feature . maximum tensile strain can therefore be reliably limited by adjusting the initial length of the convex arc shaped ductile hinge 64 over which the total elongation is distributed . the ductile hinges illustrated in fig3 a - e are examples of different structures that will function as a stress / strain concentrator . many other stress / strain concentrator configurations may also be used as the ductile hinges in the present invention . the ductile hinges according to the present invention generally include an abrupt change in width of a strut that functions to concentrate stresses and strains in the narrower section of the strut . these ductile hinges also generally include features to limit mechanical deflection of attached struts and features to control material strain during large strut deflections . although the ductile hinges have been illustrated in fig2 as positioned at the ends of each of the axial slots 22 , they may also be positioned at other locations in other designs without departing from the present invention . an alternative embodiment of a tissue supporting device 80 is illustrated in fig4 a and in the enlarged side view of fig4 b . the tissue supporting device 80 includes a plurality of cylindrical tubes 82 connected by s - shaped bridging elements 84 . the bridging elements 84 allow the tissue supporting device to bend axially when passing through the tortuous path of the vasculature to the deployment site and allow the device to bend when necessary to match the curvature of a lumen to be supported . the s - shaped bridging elements 84 provide improved axial flexibility over prior art devices due to the thickness of the elements in the radial direction which allows the width of the elements to be relatively small without sacrificing radial strength . for example , the width of the bridging elements 84 may be about 0 . 0012 - 0 . 0013 inches ( 0 . 0305 - 0 . 0330 mm ). each of the cylindrical tubes 82 has a plurality of axial slots 86 extending from an end surface of the cylindrical tube toward an opposite end surface . a plurality of axial struts 88 having ductile hinges 90 are formed between the axial slots 86 . the ductile hinges 90 are formed by circumferential slots 92 formed at the interior ends of the axial slots 86 and opposed notches 94 . the notches 94 each have two opposed angled walls 96 which function as a stop to limit geometric deflection of the ductile hinge , and thus limit maximum device expansion . as the cylindrical tubes 82 are expanded and bending occurs at the ductile hinges 90 , the angled side walls 96 of the notches 94 move toward each other . once the opposite side walls 96 of a notch come into contact with each other , they resist further expansion of the particular ductile hinge causing further expansion to occur at other sections of the tissue supporting device . this geometric deflection limiting feature is particularly useful where uneven expansion is caused by either variations in the tissue supporting device 80 due to manufacturing tolerances or uneven balloon expansion . the tissue supporting device 20 , 80 according to the present invention may be formed of any ductile material , such as steel , gold , silver , tantalum , titanium , nitinol , other shape memory alloys , other metals , or even some plastics . one preferred method for making the tissue supporting device 20 , 80 involves forming a cylindrical tube and then laser cutting the slots 22 , 26 , 86 , 92 and notches 94 into the tube . alternatively , the tissue supporting device may be formed by electromachining , chemical etching followed by rolling and welding , or any other known method . the design and analysis of stress / strain concentration for ductile hinges , and stress / strain concentration features in general , is complex . for example , the stress concentration factor for the simplified ductile hinge geometry of fig3 a can be calculated and is given by the following expression where d is the width of the struts 28 , h is the height of the circular grooves 38 , 40 , and r is the radius of curvature of the grooves . for purposes of this example the ratio of h / r is taken to be 4 . however , other ratios of h / r can also be implemented successfully . k = 4 . 935 - 7 . 586   ( 2  h d ) + 0 . 515   ( 2  h d ) 2 + 0 . 432   ( 2  h d ) 3 the stress concentration factors are generally useful only in the linear elastic range . stress concentration patterns for a number of other geometries can be determined through photoelastic measurements and other experimental methods . stent designs based on the use of stress / strain concentration features , or ductile hinges , generally involve more complex hinge geometries and operate in the non - linear elastic and plastic deformation regimes . the general nature of the relationship among applied forces , material properties , and ductile hinge geometry can be more easily understood through analysis of an idealized hinge 66 as shown in fig5 a - 5 c . the hinge 66 is a simple beam of rectangular cross section having a width h , length l and thickness b . the idealized hinge 66 has elastic - ideally - plastic material properties which are characterized by the ideal stress / strain curve of fig5 d . it can be shown that the “ plastic ” or “ ultimate bending moment ” for such a beam is given by the expression : m p ≡ m ult = δ  yp  bh 2 4 where b corresponds to the cylindrical tube wall thickness , h is the circumferential width of the ductile hinge , and δ yp is the yield stress of the hinge material . assuming only that expansion pressure is proportional to the plastic moment , it can be seen that the required expansion pressure to expand the tissue supporting device increases linearly with wall thickness b and as the square of ductile hinge width h . it is thus possible to compensate for relatively large changes in wall thickness b with relatively small changes in hinge width h . while the above idealized case is only approximate , empirical measurements of expansion forces for different hinge widths in several different ductile hinge geometries have confirmed the general form of this relationship . accordingly , for different ductile hinge geometries it is possible to increase the thickness of the tissue supporting device to achieve radiopacity while compensating for the increased thickness with a much smaller decrease in hinge width . ideally , the stent wall thickness b should be as thin as possible while still providing good visibility on a fluoroscope . for most stent materials , including stainless steel , this would suggest a thickness of about 0 . 005 - 0 . 007 inches ( 0 . 127 - 0 . 178 mm ) or greater . the inclusion of ductile hinges in a stent design can lower expansion forces / pressures to very low levels for any material thickness of interest . thus ductile hinges allow the construction of optimal wall thickness tissue supporting devices at expansion force levels significantly lower than current non - visible designs . the expansion forces required to expand the tissue supporting device 20 according to the present invention from an initial condition illustrated in fig2 to an expanded condition is between 1 and 5 atmospheres , preferably between 2 and 3 atmospheres . the expansion may be performed in a known manner , such as by inflation of a balloon or by a mandrel . the tissue supporting device 20 in the expanded condition has a diameter which is preferably up to three times the diameter of the device in the initial unexpanded condition . many tissue supporting devices fashioned from cylindrical tubes comprise networks of long , narrow , prismatic beams of essentially rectangular cross section as shown in fig6 . these beams which make up the known tissue supporting devices may be straight or curved , depending on the particular design . known expandable tissue supporting devices have a typical wall thickness b of 0 . 0025 inches ( 0 . 0635 mm ), and a typical strut width h of 0 . 005 to 0 . 006 inches ( 0 . 127 - 0 . 1524 mm ). the ratio of b : h for most known designs is 1 : 2 or lower . as b decreases and as the beam length l increases , the beam is increasingly likely to respond to an applied bending moment m by buckling , and many designs of the prior art have displayed this behavior . this can be seen in the following expression for the “ critical buckling moment ” for the beam of fig6 . m crit = π   b 3  h  eg  ( 1 - 0 . 63   b / h ) 6  l by contrast , in a ductile hinge based design according to the present invention , only the hinge itself deforms during expansion . the typical ductile hinge 32 is not a long narrow beam as are the struts in the known stents . wall thickness of the present invention may be increased to 0 . 005 inches ( 0 . 127 mm ) or greater , while hinge width is typically 0 . 002 - 0 . 003 inches ( 0 . 0508 - 0 . 0762 mm ), preferably 0 . 0025 inches ( 0 . 0635 mm ) or less . typical hinge length , at 0 . 002 to 0 . 005 inches ( 0 . 0508 - 0 . 0127 mm ), is more than an order of magnitude less than typical strut length . thus , the ratio of b : h in a typical ductile hinge 32 is 2 : 1 or greater . this is an inherently stable ratio , meaning that the plastic moment for such a ductile hinge beam is much lower than the critical buckling moment m crit , and the ductile hinge beam deforms through normal strain - curvature . ductile hinges 32 are thus not vulnerable to buckling when subjected to bending moments during expansion of the tissue supporting device 20 . to provide optimal recoil and crush - strength properties , it is desirable to design the ductile hinges so that relatively large strains , and thus large curvatures , are imparted to the hinge during expansion of the tissue supporting device . curvature is defined as the reciprocal of the radius of curvature of the neutral axis of a beam in pure bending . a larger curvature during expansion results in the elastic curvature of the hinge being a small fraction of the total hinge curvature . thus , the gross elastic recoil of the tissue supporting device is a small fraction of the total change in circumference . it is generally possible to do this because common stent materials , such as 316l stainless steel have very large elongations - to - failure ( i . e ., they are very ductile ). it is not practical to derive exact expressions for residual curvatures for complex hinge geometries and real materials ( i . e ., materials with non - idealized stress / strain curves ). the general nature of residual curvatures and recoil of a ductile hinge may be understood by examining the moment - curvature relationship for the elastic - ideally - plastic rectangular hinge 66 shown in fig5 a - c . it may be shown that the relationship between the applied moment and the resulting beam curvature is : m = m p  [ 1 - 1 3  ( y o h / 2 ) 2 ] = 3 / 2  m yp  [ 1 - 1 3  ( κ yp κ ) 2 ] this function is plotted in fig7 . it may be seen in this plot that the applied moment m asymptotically approaches a limiting value m p , called the plastic or ultimate moment . beyond { fraction ( 11 / 12 )} m p large plastic deformations occur with little additional increase in applied moment . when the applied moment is removed , the beam rebounds elastically along a line such as a - b . thus , the elastic portion of the total curvature approaches a limit of 3 / 2 the curvature at the yield point . these relations may be expressed as follows : m p = 3 2  m yp ⇒ κ rebound = 3 2  κ yo imparting additional curvature in the plastic zone cannot further increase the elastic curvature , but will decrease the ratio of elastic to plastic curvature . thus , additional curvature or larger expansion of the tissue supporting device will reduce the percentage recoil of the overall stent structure . as shown in fig8 when a rigid strut 28 is linked to the ductile hinge 66 described above , the strut 28 forms an angle θ with the horizontal that is a function of hinge curvature . a change in hinge curvature results in a corresponding change in this angle θ . the angular elastic rebound of the hinge is the change in angle δθ that results from the rebound in elastic curvature described above , and thus angular rebound also approaches a limiting value as plastic deformation proceeds . the following expression gives the limiting value of angular elastic rebound for the idealized hinge of fig8 . θ rebound = 3  ε yp  l h where strain at the yield point is an independent material property ( yield stress divided by elastic modulus ); l is the length of the ductile hinge ; and h is the width of the hinge . for non - idealized ductile hinges made of real materials , the constant 3 in the above expression is replaced by a slowly rising function of total strain , but the effect of geometry would remain the same . specifically , the elastic rebound angle of a ductile hinge decreases as the hinge width h increases , and increases as the hinge length l increases . to minimize recoil , therefore , hinge width h should be increased and length l should be decreased . ductile hinge width h will generally be determined by expansion force criteria , so it is important to reduce hinge length to a practical minimum in order to minimize elastic rebound . empirical data on recoil for ductile hinges of different lengths show significantly lower recoil for shorter hinge lengths , in good agreement with the above analysis . the ductile hinges 32 of the tissue supporting device 20 provide a second important advantage in minimizing device recoil . the embodiment of fig2 shows a network of struts joined together through ductile hinges to form a cylinder . in this design , the struts 28 are initially parallel to an axis of the device . as the device is expanded , curvature is imparted to the hinges 32 , and the struts 28 assume an angle θ with respect to their original orientation , as shown in fig8 . the total circumferential expansion of the tissue supporting device structure is a function of hinge curvature ( strut angle ) and strut length . moreover , the incremental contribution to stent expansion ( or recoil ) for an individual strut depends on the instantaneous strut angle . specifically , for an incremental change in strut angle δθ , the incremental change in circumference δc will depend on the strut length r and the cosine of the strut angle θ . since elastic rebound of hinge curvature is nearly constant at any gross curvature , the net contribution to circumferential recoil δc is lower at higher strut angles θ . the final device circumference is usually specified as some fixed value , so decreasing overall strut length can increase the final strut angle θ . total stent recoil can thus be minimized with ductile hinges by using shorter struts and higher hinge curvatures when expanded . empirical measurements have shown that tissue supporting device designs based on ductile hinges , such as the embodiment of fig2 display superior resistance to compressive forces once expanded despite their very low expansion force . this asymmetry between compressive and expansion forces may be due to a combination of factors including the geometry of the ductile hinge , the increased wall thickness , and increased work hardening due to higher strain levels . according to one example of the tissue supporting device of the invention , the device can be expanded by application of an internal pressure of about 2 atmospheres or less , and once expanded to a diameter between 2 and 3 times the initial diameter can withstand a compressive force of about 16 to 20 gm / mm or greater . examples of typical compression force values for prior art devices are 3 . 8 to 4 . 0 gm / mm . while both recoil and crash strength properties of tissue supporting devices can be improved by use of ductile hinges with large curvatures in the expanded configuration , care must be taken not to exceed an acceptable maximum strain level for the material being used . for the ductile hinge 44 of fig3 b , for example , it may be shown that the maximum material strain for a given bend angle is given by the expression : ε max = h l   θ 2 where ε max is maximum strain , h is ductile hinge width , l is ductile hinge length and θ is bend angle in radians when strain , hinge width and bend angle are determined through other criteria , this expression can be evaluated to determine the correct ductile hinge length l . for example , suppose the ductile hinge 44 of fig3 b was to be fabricated of 316l stainless steel with a maximum strain of 30 %; ductile hinge width h is set at 0 . 0025 inch ( 0 . 0635 mm ) by expansion force criteria ; and the bend angle θ is mechanically limited to 0 . 5 radians (≈ 30 %) at full stent expansion . solving the above expression for l gives the required ductile hinge length of at least about 0 . 0033 inches ( 0 . 0838 mm ). similar expressions may be developed to determine required lengths for more complicated ductile hinge geometries , such as shown in fig3 e . typical values for the prismatic portions of these curved ductile hinges range from about 0 . 002 to about 0 . 0035 inches ( 0 . 051 - 0 . 089 mm ) in hinge width and about 0 . 002 to about 0 . 006 inches ( 0 . 051 - 0 . 152 mm ) in hinge length . the tissue supporting device design of fig4 a and 4 b include a stop which limits the maximum geometric deflection at the ductile hinges by the design of the angled walls 96 of the notches 94 . in many designs of the prior art , circumferential expansion was accompanied by a significant contraction of the axial length of the stent which may be up to 15 % of the initial device length . excessive axial contraction can cause a number of problems in device deployment and performance including difficulty in proper placement and tissue damage . designs based on ductile hinges 32 can minimize the axial contraction , or foreshortening , of a tissue supporting device during expansion as follows . fig9 a and 9 b illustrate an exaggerated ductile hinge 32 and shortened struts 28 in initial and expanded conditions . each strut 28 is attached to two ductile hinges 32 at opposite ends . each ductile hinge 32 has an instant center of rotation c 1 , c 2 that is an effective pivot point for the attached strut 28 . initially , during expansion the pivot point c 1 is displaced vertically by a distance d until c 1 is positioned even with c 2 as shown in fig9 b . when the array is expanded vertically , the axial struts 28 move in a circular arc with respect to the pivot points , as shown in fig9 b . it can be seen that the horizontal distance e between pivot points c 1 and c 2 actually increases initially , reaching a maximum e max when the two points are on the same horizontal axis as shown in fig9 b . as the vertical expansion continues , the device compresses axially back to its original length . only when vertical expansion of the array continues beyond the point where the horizontal distance e between c 1 and c 2 is the same as the original horizontal distance e does the overall length of the array actually begin to contract . for the stent shown in fig2 for example , approximately ⅓ of the total circumferential expansion has been accomplished by the time the configuration of fig9 b is reached , and the stent exhibits very low axial contraction . this ability to control axial contraction based on hinge and strut design provides great design flexibility when using ductile hinges . for example , a stent could be designed with zero axial contraction . an alternative embodiment that illustrates the trade off between crush strength and axial contraction is shown in fig1 . fig1 shows a portion of a tissue supporting device 70 having an array of struts 72 and ductile hinges 74 in the unexpanded state . the struts 72 are positioned initially at an angle θ 1 with respect to a longitudinal axis x of the device . as the device is expanded radially from the unexpanded state illustrated in fig1 , the angle θ 1 increases . in this case the device contracts axially from the onset of vertical expansion throughout the expansion . once the device has been completely expanded the final angle θ 1 made by the strut 72 with the horizontal will be much greater than the angle θ in the device of fig8 a and 8 b . as shown previously , a higher final strut angle θ 1 , can significantly increase crush strength and decrease circumferential recoil of the stent structure . however , there is a trade off between increased crush strength and increase in axial contraction . according to one example of the present invention , the struts 72 are positioned initially at an angle of about 0 ° to 45 ° with respect to a longitudinal axis of the device . as the device is expanded radially from the unexpanded state illustrated in fig1 , the strut angle increases to about 20 ° to 80 °. according to one alternative embodiment of the present invention , the expandable tissue supporting device can also be used as a delivery device for certain beneficial agents including drugs , chemotherapy , or other agents . due to the structure of the tissue supporting device incorporating ductile hinges , the widths of the struts can be substantially larger than the struts of the prior art devices . the struts due to their large size can be used for beneficial agent delivery by providing beneficial agent on the struts or within the struts . examples of beneficial agent delivery mechanisms include coatings on the struts , such as polymer coatings containing beneficial agents , laser drilled holes in the struts containing beneficial agent , and the like . while the invention has been described in detail with reference to the preferred embodiments thereof , it will be apparent to one skilled in the art that various changes and modifications can be made and equivalents employed , without departing from the present invention .