Patent Publication Number: US-2011066222-A1

Title: Polymeric Stent and Method of Making Same

Description:
FIELD OF THE INVENTION 
     This invention relates generally to fabrication of implantable prostheses, more particularly, to fabrication of stents from blow molded polymeric tubes. 
     BACKGROUND OF THE INVENTION 
     Radially expandable endoprostheses are artificial devices adapted to be implanted in an anatomical lumen. An “anatomical lumen” refers to a cavity, duct, of a tubular organ such as a blood vessel, urinary tract, and bile duct. Stents are examples of endoprostheses that are generally cylindrical in shape and function to hold open and sometimes expand a segment of an anatomical lumen. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments, stents reinforce the walls of the blood vessel and prevent restenosis following angioplasty in the vascular system. “Restenosis” refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success. 
     The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through an anatomical lumen to a desired treatment site, such as a lesion. “Deployment” corresponds to expansion of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into an anatomical lumen, advancing the catheter in the anatomical lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen. 
     In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter. Mounting the stent typically involves compressing or crimping the stent onto the balloon prior to insertion in an anatomical lumen. At the treatment site within the lumen, the stent is expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn from the stent and the lumen, leaving the stent at the treatment site. In the case of a self-expanding stent, the stent may be secured to the catheter via a retractable sheath. When the stent is at the treatment site, the sheath may be withdrawn which allows the stent to self-expand. 
     The stent must be able to satisfy a number of functional requirements. The stent must be capable of withstanding the structural loads, namely radial compressive forces, imposed on the stent as it supports the walls of a vessel after deployment. Therefore, a stent must possess adequate radial strength. Radial strength, which is the ability of a stent to resist radial compressive forces, is due to strength and rigidity around a circumferential direction of the stent. After deployment, the stent must also adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. 
     In addition to high radial strength, the stent must also possess sufficient toughness so that the stent exhibits sufficient flexibility to allow for crimping on the a delivery device, flexure during delivery through an anatomical lumen, and expansion at the treatment site. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous vascular path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure. A stent should have sufficient toughness so that it is resistant to crack formation, particularly, in high strain regions. 
     Furthermore, it may be desirable for a stent to be made of a biodegradable or bioerodable polymer. In many treatment applications, the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Also, it is believed that biodegradable stents allow for improved healing of the anatomical lumen as compared to metal stents, which may lead to a reduced incidence of late stage thrombosis. 
     However, a potential shortcoming of polymer stents compared to metal stents of the same dimensions, is that polymer stents typically have less radial strength and rigidity. Relatively low radial strength potentially contributes to relatively high recoil of polymer stents after implantation into an anatomical lumen. “Recoil” refers to the undesired retraction of a stent radially inward from its deployed diameter due to radially compressive forces that bear upon it after deployment. Furthermore, another potential problem with polymer stents is that struts can crack or fracture during crimping, delivery and deployment, especially for brittle polymers. Some crystalline or semi-crystalline polymers that may be suitable for use in implantable medical devices generally have potential shortcomings with respect to some mechanical characteristics, in particular, fracture toughness, when used in stents. 
     Some polymers, such as poly(L-lactide) (“PLLA”), poly(L-lactide-co-glycolide) (“PLGA”), poly(L-lactide-co-D-lactide) (“PLLA-co-PDLA”) with less than 10% D-lactide, and PLLD/PDLA stereocomplex, are stiff and strong but can exhibit a brittle fracture mechanism at physiological conditions in which there is little or no plastic deformation prior to failure. Physiological conditions include, but are limited to, human body temperature, approximately 37° C. A stent fabricated from such polymers can have insufficient toughness for the range of use of a stent. As a result, cracks, particularly in high strain regions, can be induced which can result in mechanical failure of the stent. 
     Accordingly, there is a need for manufacturing methods for fabricating polymeric stents with sufficient radial strength, fracture toughness, low recoil, and sufficient shape stability. 
     SUMMARY OF THE INVENTION 
     Briefly and in general terms, the present invention is directed to a stent and a method of forming a stent. 
     In aspects of the present invention, a method of forming a stent comprises deforming a precursor tube of poly(L-lactide) to form a deformed tube. The deforming includes maintaining fluid pressure in the tube at a process pressure from about 110 psi to about 150 psi, heating the tube to a process temperature from about 160 deg F. to about 220 deg F., radially expanding the precursor tube according to a radial expansion ratio between about 300% and about 450% during the maintaining of fluid pressure and the heating, and axially extending the precursor tube according to an axial extension ratio from about 20% to about 100% during the maintaining of fluid pressure and the heating. The method further comprises forming a network of stent struts from the deformed tube. In detailed aspects of the present invention, heating the tube includes heating a tubular mold containing the tube, the heating including moving a heat source disposed outside the tube at a linear rate of movement parallel to the central axis of the mold, the linear rate of movement being about 0.1 mm to 0.7 mm per minute. In further aspects of the present invention, a stent comprises the network of stent struts formed from the deformed tube. 
     In aspects of the present invention, a method of making a stent comprises providing a poly(L-lactide) tube inside a tubular mold, heating a segment of the tube with a heat source, the segment of the tube being heated to a process temperature from about 160 deg F. to about 220 deg F., and moving the heat source in a process direction. The method further comprises causing deformation of the heated segment to form a deformed segment of the tube, the deformation propagating in the process direction, the deformation including radial expansion and axial extension of the tube, the radial expansion in accordance with a radial expansion ratio between about 300% and about 450%, the axial extension in accordance with an axial extension ratio between about 20% and about 100%. The method further comprises forming stent struts from the deformed segment. 
     A method for making a stent, according to aspects of the present invention, comprises deforming a precursor tube of a polymer formulation to form a deformed tube. The deforming includes maintaining fluid pressure in the tube at a process pressure from about 50 psi to about 200 psi, heating the tube to a process temperature from about 100 deg. F. to about 300 deg F., radially expanding the precursor tube according to a radial expansion ratio between about 100% and about 600% during the maintaining of fluid pressure and the heating, and axially extending the precursor tube according to an axial extension ratio from about 10% to about 200% during the maintaining of fluid pressure and the heating. The method further comprises forming a network of stent struts from the deformed tube. In further aspects, the polymer formation is a material selected from the group consisting of PLGA, PLLA-co-PDLA, PLLD/PDLA stereocomplex, and PLLA-based polyester block copolymer containing a rigid segment and a soft segment, the rigid segment being PLLA or PLGA, the soft segment being PCL or PTMC. 
     A method of making a stent, according to aspects of the present invention, comprises providing a polymer tube inside a tubular mold, the polymer tube made of a polymer formulation selected from the group consisting of PLGA, PLLA-co-PDLA, PLLD/PDLA stereocomplex, and PLLA-based polyester block copolymer containing a rigid segment and a soft segment, the rigid segment being PLLA or PLGA, the soft segment being PCL or PTMC. The method further comprises heating a segment of the tube with a heat source, the segment of the tube being heated to a process temperature from about 100 deg F. to about 300 deg F. The method further comprises moving the heat source in a process direction. The method further comprises causing deformation of the heated segment to form a deformed segment of the tube, the deformation propagating in the process direction, the deformation including radial expansion and axial extension of the tube, the radial expansion in accordance with a radial expansion ratio between about 100% and about 600%, the axial extension in accordance with an axial extension ratio from about 10% to about 200%. The method further comprises forming stent struts from the deformed segment. 
     The features and advantages of the invention will be more readily understood from the following detailed description which should be read in conjunction with the accompanying drawings. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a perspective view of a stent. 
         FIG. 2  is a perspective view of a polymer tube for making a stent. 
         FIG. 3A  is an axial cross-sectional view of a blow molding system showing a blow mold and a polymer tube in the blow mold. 
         FIG. 3B  is a radial cross-sectional view of the blow molding system of  FIG. 3A , showing nozzles heating the blow mold. 
         FIG. 3C  is an axial cross-sectional view of the blow molding system of  FIG. 3A , showing the polymer tube being deformed. 
         FIG. 3D  is an axial cross-sectional view of the blow molding system of  FIG. 3A , showing further deformation of the polymer tube. 
         FIG. 4  is a schematic plot of quiescent crystal nucleation rate and the quiescent crystal growth rate, and the overall rate of quiescent crystallization. 
         FIG. 5  is a top view of a pattern of struts for a stent. 
         FIG. 6  is a perspective view of a portion of a stent having the pattern of  FIG. 5 . 
     
    
    
     DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS 
     The various embodiments of the present invention relate to methods of fabricating a polymeric stent that has good or optimal toughness and selected mechanical properties along the axial, radial and circumferential directions. The present invention can be applied to devices including, but is not limited to, self-expandable stents, balloon-expandable stents, stent-grafts, grafts (e.g., aortic grafts), and generally to tubular implantable medical devices. 
     For the purposes of the present invention, the following terms and definitions apply: 
     The glass transition temperature (referred to herein as “Tg”) is the temperature at which the amorphous domains of a polymer change from a brittle vitreous state to a solid deformable or ductile state at atmospheric pressure. In other words, Tg corresponds to the temperature where the onset of segmental motion in the chains of the polymer occurs. Tg of a given polymer can be dependent on the heating rate and can be influenced by the thermal history of the polymer. Furthermore, the chemical structure of the polymer heavily influences the glass transition by affecting mobility of polymer chains. 
     “Stress” refers to force per unit area, as in the force acting through a small area within a plane within a subject material. Stress can be divided into components, normal and parallel to the plane, called normal stress and shear stress, respectively. Tensile stress, for example, is a normal component of stress that leads to expansion (increase in length) of the subject material. In addition, compressive stress is a normal component of stress resulting in compaction (decrease in length) of the subject material. 
     “Strain” refers to the amount of expansion or compression that occurs in a material at a given stress or load. Strain may be expressed as a fraction or percentage of the original length, i.e., the change in length divided by the original length. Strain, therefore, is positive for expansion and negative for compression. 
     “Modulus” may be defined as the ratio of a component of stress or force per unit area applied to a material divided by the strain along an axis of applied force that results from the applied force. For example, a material has both a tensile and a compressive modulus. 
     “Toughness” is the amount of energy absorbed prior to fracture, or equivalently, the amount of work required to fracture a material. One measure of toughness is the area under a stress-strain curve from zero strain to the strain at fracture. The stress is proportional to the tensile force on the material and the strain is proportional to its length. The area under the curve then is proportional to the integral of the force over the distance the polymer stretches before breaking. This integral is the work (energy) required to break the sample. The toughness is a measure of the energy a sample can absorb before it breaks. There is a difference between toughness and strength. A material that is strong, but not tough is said to be brittle. Brittle materials are strong, but cannot deform very much before breaking. 
     As used herein, the terms “axial” and “longitudinal” are used interchangeably and refer to a direction, orientation, or line that is parallel or substantially parallel to the central axis of a stent or the central axis of a tubular construct. The term “circumferential” refers to the direction along a circumference of the stent or tubular construct. The term “radial” refers to a direction, orientation, or line that is perpendicular or substantially perpendicular to the central axis of the stent or the central axis of a tubular construct. 
     Mechanical properties of a polymer may be modified by processes that alter the molecular structure or morphology of the polymer. Polymers in the solid state may be completely amorphous, partially crystalline, or almost completely crystalline. Crystalline regions are where polymer molecules are geometrically arranged in a regular order or pattern. Crystalline regions may be clusters of polymer crystals. Each crystal may have polymer molecules arranged geometrically around a nucleus. Amorphous regions in a polymer matrix are where polymer molecules have no regular order or arrangement. Amorphous regions may be located between ordered polymer chains, between polymer crystals, and between clusters of polymer crystals. 
     Polymer molecule chains in crystalline regions may radiate outwardly from many nuclei without a preferred orientation or alignment. In other instances, polymer molecules in crystalline regions may have a preferred orientation or long range order with respect to a particular direction, as may occur with strain induced crystallization. 
     As indicated above, molecular orientation in a polymer may be induced, and hence modify mechanical properties, by applying a stress to the polymer which deforms the polymer in the direction of the applied stress. The degree of molecular orientation induced with applied stress may depend upon the temperature of the polymer. For example, below the glass transition temperature, Tg, of a polymer, polymer segments do not have sufficient energy to move past one another. In general, molecular orientation may not be induced without sufficient segmental mobility. Above Tg, molecular orientation may be induced with applied stress since rotation of polymer chains, and hence segmental mobility is possible. Between Tg and the melting temperature of the polymer (referred to herein as “Tm”), rotational barriers exist, however, the barriers are not great enough to substantially prevent segmental mobility. As the temperature of a polymer is increased above Tg, the energy barriers to rotation decrease and segmental mobility of polymer chains tend to increase. As a result, as the temperature increases, molecular orientation is more easily induced with applied stress. A polymer with a high level of polymer chain alignment would have enhanced strength and toughness in the direction of alignment of the polymer chains. 
     Referring now in more detail to the exemplary drawings for purposes of illustrating embodiments of the invention, wherein like reference numerals designate corresponding or like elements among the several views, there is shown in  FIG. 1  a stent  100  in an uncrimped state or a deployed state. The stent  110  has a scaffolding composed of a pattern of interconnected structural elements or struts  110 . The struts  110  form a hollow body having cylindrical shape or tubular shape. The struts  110  have straight or relatively straight portions  120 . The struts also have bending elements  130 ,  140 , and  150 , which are configured to bend during stent crimping and deployment to allow the straight portions  120  to collapse next to each other and expand apart from each other. The tubular body has two opposite open ends, a central passageway that runs from one end to the opposite end, and a central axis  160  that extends longitudinally through the center of the central passageway. Surfaces of the struts  110  that face radially inward toward the central axis  160  form the luminal or inner surface of the stent. Surfaces of the struts  110  that face radially outward away from the central axis  160  form the abluminal or outer surface of the stent. When deployed in a blood vessel, the luminal surface faces blood flowing through the central passageway of the stent and the abluminal surface faces and supports the walls of the blood vessel. 
     The stresses involved during compression and expansion are generally distributed throughout the various structural elements of the stent pattern. The present invention is not limited to the stent pattern depicted in  FIG. 1 . The variation in stent patterns is virtually unlimited. 
     The struts  110 , which may serve as the underlying structure or substrate of a stent, is completely or at least in part made from a biodegradable polymer or combination of biodegradable polymers, a biostable polymer or combination of biostable polymers, or a combination of biodegradable and biostable polymers. Suitable examples of polymers include without limitation, poly(L-lactide) (“PLLA”) and poly(lactic-co-glycolic acid) (“PLGA”). PLLA and PLGA are semi-crystalline polymers in that their morphology includes crystalline and amorphous regions, though the amount of crystallinity can be altered. For example, the maximum crystallinity of pure PLLA is about 70%, while that of PLGA with 20% GA is below 10%. Additionally, a polymer-based coating on the stent substrate can be a biodegradable polymer or combination of biodegradable polymers, a biostable polymer or combination of biostable polymers, or a combination of biodegradable and biostable polymers. 
     The stent  100  is fabricated from a polymeric tube  200  shown in  FIG. 2 . The tube  200  may serves as a stent precursor construct in the sense that further processing may be performed on the tube before the pattern of stent struts is cut formed from the tube. The tube  200  is cylindrically-shaped with an outside diameter  205 , an inside diameter  210 , an outside surface  215 , and a central axis  220 . The tube  200  may be formed by various types of methods, including, but not limited to extrusion, injection molding, and rolling a flat sheet of material to form a tube. A pattern of struts may be formed on the tube  200  by chemical etching, mechanical cutting, and laser cutting material away from the tube. Representative examples of lasers that may be used include, but are not limited to, excimer, carbon dioxide, and YAG. 
     In some embodiments, the polymer tube  200  can have a outer diameter of 1-4 mm. The present invention is also applicable to polymer tubes less than 1 mm or greater than 4 mm in diameter. The wall thickness of the polymer tube can be between 0.1 mm to 0.3 mm. The present invention is also applicable to wall thicknesses below 0. 1 mm and above 0.3 mm. 
     As indicated above, the tube  200  may be formed by an extrusion process. During extrusion, a polymer melt is conveyed through an extruder which is then formed into a tube. Extrusion tends to impart large forces on the polymer molecules in the longitudinal direction of the tube due to shear forces on the polymer melt. The shear forces arise from forcing the polymer melt through an opening of a die at the end of an extruder. Additional shear forces may arise from any pulling and forming of the polymer melt upon exiting the die, such as may be performed in order to bring the extruded material to the desired dimensions of a finished tube. As a result, polymer tubes formed by some extrusion methods tend to possess a significant degree of molecular or crystal orientation in the direction that the polymer is extruded with a relatively low degree of orientation in the circumferential direction. 
     The degree of pulling that is applied to the polymer melt as it exits a die of an extruder and, thus, the degree of longitudinal orientation induced in the finished tube  200  can be partially characterized by what is referred to as a “draw down ratio.” Typically, the polymer melt is in the form of an annular film as it is extruded through and exits an annular opening of the die. The annular film has an initial outer diameter upon exiting the annular opening. The annular film is drawn or pulled, which causes a reduction of the annular film cross-sectional size to the final outer diameter. The drawn down portion of the tube may be cooled to ensure that it maintains its shape and diameter. The final outer diameter corresponds to the outer diameter of the finished, solidified polymeric tube  200 . The draw down ratio is defined as the ratio of the final outer diameter to the initial outer diameter. 
     As indicated above, the finished, solidified polymeric tube  200  may serve as a precursor construct in that further processing of the tube is performed. Further processing includes heating combined with deformation of the tube in radial and axial directions, such as may be performed by blow molding. After blow molding, pieces of the blow molded tube are cut away to form stent struts. 
     The degree of radial expansion that the polymer tube undergoes can partially characterize the degree of induced circumferential molecular or crystal orientation as well as strength of the deformed tube in a circumferential direction. The degree of radial expansion is quantified by a radial expansion (“RE”) ratio, defined as RE Ratio=(Inside Diameter of Expanded Tube)/(Original Inside Diameter of the tube). The RE ratio can also be expressed as a percentage, defined as RE %=(RE ratio−1)×100%. 
     The degree of axial extension that the polymer tube undergoes can partially characterize induced axial molecular or crystal orientation as well as strength of the deformed tube in an axial direction. The degree of axial extension is quantified by an axial extension (“AE”) ratio, defined as AE Ratio=(Length of Extended Tube)/(Original Length of the Tube). The AE ratio can also be expressed as a percentage, defined as AE %=(AE ratio−1)×100%. 
     Blow molding includes first positioning the tube  200  in a hollow cylindrical member or mold. The mold controls the degree of radial deformation of the polymer tube by limiting the deformation of the outside diameter or surface of the polymer tube to the inside diameter of the mold. The inside diameter of the mold may correspond to a diameter less than or equal to a desired diameter of the finished polymer tube. 
     While in the mold, the temperature of the polymer tube  200  is heated to a temperature above Tg of the polymer to facilitate deformation. The temperature to which the tube  200  is heated during blow molding is a processing parameter referred to as the “expansion temperature” or “process temperature.” The heating of the polymer tube  200  to the expansion temperature can be achieved by heating a gas to the expansion temperature and discharging the heated gas onto an exterior surface of the mold containing the polymer tube. 
     While in the mold, one end of the polymer tube  200  is sealed or blocked. Thus, introduction of gas into the opposite end of the polymer tube will increase internal fluid pressure relative to ambient pressure in a region between the outer surface of the polymer tube and the inner surface of the mold. The internal fluid pressure is a processing parameter referred to as the “expansion pressure” or “process pressure.” Examples of gas that may be used to create the expansion pressure include without limitation ambient air, substantially pure oxygen, substantially pure nitrogen, and other substantially pure inert gases. In combination with other blow molding process parameters, the expansion pressure affects the rate at which the tube deforms radially and axially. 
     Blow molding may include pulling one end of the polymer tube  200 . A tensile force, which is another processing parameter, is applied to one end of the polymer tube  200  while holding the other end of the polymer tube stationary. Alternatively, the two opposite ends of the polymer tube may be pulled apart. In combination with other blow molding process parameters, the tensile force affects the rate at which the tube deforms radially and axially. 
     The radially and axially deformed polymer tube may then be cooled from above Tg to below Tg, either before or after decreasing the pressure and/or decreasing tension. Cooling the deformed tube helps insure that the tube maintains the proper shape, size, and length following radial expansion and axial extension. The rate at which the deformed tube is cooled is yet another processing parameter. Slow cooling through a temperature range between Tm and Tg might result in a loss of amorphous chain orientation and cause a decrease in fracture toughness of the finished stent. Preferably, though not necessarily, the tube can be cooled quickly or quenched in relatively cold gas or liquid to a temperature below Tg to maintain chain orientation that was formed during tubing expansion. 
       FIGS. 3A-D  schematically depicts a molding system  300  for simultaneous radial and axial deformation of a polymer tube.  FIG. 3A  depicts an axial cross-section of a polymer tube  301  with an undeformed outside diameter  305  positioned within a mold  310 . The mold  310  limits the radial deformation of the polymer tube  301  to a diameter  315  corresponding to the inside diameter of the mold  310 . The polymer tube  301  is closed at a distal end  320 . A gas is conveyed, as indicated by an arrow  325 , into an open end  321  of the polymer tube  301  to increase internal fluid pressure within tube  301 . 
     A tensile force  322  is applied to the distal end  320  in an axial direction. In other embodiments, a tensile force is applied at the proximal end  321  and the distal end  320 . 
     A circular band or segment of the polymer tube  300  is heated by a nozzle  330 . The nozzle has fluid ports that direct a heated fluid, such as hot air, at two circumferential locations of the mold  310 , as shown by arrows  335  and  340 .  FIG. 3B  depicts a radial cross-section showing the tube  301  within the mold  310 , and the nozzle  330  supported by structural members  360 . Additional fluid ports can be positioned at other circumferential locations of the mold  310  to facilitate uniform heating around a circumference of the mold  310  and the tube  301 . The heated fluid flows around the mold  310 , as shown by arrows  355 , to heat the mold  310  and the tube  301  to a predetermined temperature above ambient temperature. 
     The nozzle  330  translates along the longitudinal axis  373  of the mold  310  as shown by arrows  365  and  367 . That is, the nozzle  330  moves linearly in a direction parallel to the longitudinal axis  373  of the mold  310 . As the nozzle  330  translates along the axis of the mold  310 , the tube  301  radially deforms. The combination of elevated temperature of the tube  301 , the applied axial tension, and the applied internal pressure cause simultaneous axial and radial deformation of the tube  301 , as depicted in  FIGS. 3C and 3D . 
       FIG. 3C  depicts the system  300  with an undeformed section  371 , a deforming section  372 , and a deformed section  370  of the polymer tube  301 . Each section  370 ,  371 ,  372  is circular in the sense that each section extends completely around the central axis  373 . The deforming section  372  is in the process of deforming in a radial direction, as shown by arrow  380 , and in an axial direction, as shown by arrow  382 . The deformed section  370  has already been deformed and has an outside diameter that is the same as the inside diameter of the mold  310 . 
       FIG. 3D  depicts the system  300  at some time period after  FIG. 3C . The deforming section  372  in  FIG. 3D  is located over a portion of what was an undeformed section in  FIG. 3C . Also, the deformed section  370  in  FIG. 3D  is located over what was the deforming section  372  in  FIG. 3C . Thus it will be appreciated that the deforming section  372  propagates linearly along the longitudinal axis  373  in the same general direction  365 ,  367  that the heat sources  330  are moving. 
     In  FIG. 3D , the deforming section  372  has propagated or shifted by an axial distance  374  from its former position in  FIG. 2D . The deformed section  370  has grown longer by the same axial distance  374 . Deformation of the tube  301  occurs progressively at a selected longitudinal rate along the longitudinal axis  373  of the tube. Also, the tube  301  has increased in length by a distance  323  compared to  FIG. 3C . 
     Depending on other processing parameters, the speed at which the heat sources or nozzles  330  are linearly translated over the mold  310  may correspond to the longitudinal rate of propagation (also referred to as the axial propagation rate) of the polymer tube  301 . Thus, the distance  374  that the heat sources  330  have moved is the same distance  375  that the deformed section  370  has lengthened. 
     The rate or speed at which the nozzles  330  are linearly translated over the mold  310  is a processing parameter that relates to the amount of time a segment of the polymer tube is heated at the expansion temperature and the uniformity of such heating in the polymer tube segment. 
     It is to be understood that the tensile force, expansion temperature, and expansion pressure are applied simultaneously to the tube  301  while the nozzle  330  moves linearly at a constant speed over the mold. Again, the “expansion pressure” is the internal fluid pressure in the polymer tube while it is blow molded inside the mold. In  FIGS. 3A-3D , the “expansion temperature” is the temperature to which a limited segment of the polymer tube is heated during blow molding. The “limited segment” is the segment of the polymer tube surrounded by the nozzle  330 . The “limited segment” may include the deforming section  372 . The heating of the polymer tube to the expansion temperature can be achieved by heating a gas to the expansion temperature and discharging the heated gas from the nozzle  330  onto the mold  310  containing the polymer tube. 
     The processing parameters of the above-described blow molding process include without limitation the tensile force, expansion temperature, the expansion pressure, and nozzle translation rate or linear movement speed. It is expected that the rate at which the tube deforms during blow molding depends at least upon these parameters. The deformation rate has both a radial component, indicated by arrow  380  in  FIGS. 3C and 3D , and an axial component, indicated by an arrow  382 . It is believed that the radial deformation rate has a greater dependence on the expansion pressure and the axial component has a greater dependence on the translation rate of the heat source along the axis of the tube. It is also expected that the deformation rate is dependant upon the pre-existing morphology of the polymer in the undeformed section  371 . Also, since deformation rate is a time dependent process, it is expected to have an effect on the resulting polymer morphology of the deformed tube after blow molding. 
     The term “morphology” refers to the microstructure of the polymer which maybe characterized, at least in part, by the percent crystallinity of the polymer, the relative size of crystals in the polymer, the degree of uniformity in spatial distribution of crystals in the polymer, and the degree of long rage order or preferred orientation of molecules and/or crystals. The crystallinity percentage refers to the proportion of crystalline regions to amorphous regions in the polymer. Polymer crystals can vary in size and are sometimes geometrically arranged around a nucleus, and such arrangement may be with or without a preferred directional orientation. A polymer crystal may grow outwardly from the nucleus as additional polymer molecules join the ordered arrangement of polymer molecule chains. Such growth may occur along a preferred directional orientation. 
     Applicant believes that all the above-described processing parameters affect the morphology of the deformed polymer tube  301 . As used herein, “deformed tube  301 ” and “blow molded tube  301 ” are used interchangeably and refer to the deformed section  370  of the polymer tube  301  of  FIGS. 3C and 3D . Without being limited to a particular theory, Applicant believes that increasing the crystallinity percentage will increase the strength of the polymer but also tends to make the polymer brittle and prone to fracture when the crystallinity percentage reaches a certain level. Without being limited to a particular theory, Applicant believes that having a polymer with relatively small crystal size has higher fracture toughness or resistance to fracture. Applicant also believes that having a deformed tube  301  with spatial uniformity in the radial direction, axial direction, and circumferential direction also improves strength and fracture toughness of the stent made from the deformed tube. 
     It should be noted that the above-described processing parameters are interdependent or coupled to each other. That is, selection of a particular level for one processing parameter affects selection of appropriate levels for the other processing parameters that would result in a combination of radial expansion, axial extension, and polymer morphology that produces a stent with improved functional characteristics such as reduced incidence of strut fractures and reduced recoil. For example, a change in expansion temperature may also change the expansion pressure and nozzle translation rate required to obtain improved stent functionality. 
     Expansion temperature affects the ability of the polymer to deform (radially and axially) while simultaneously influencing crystal nucleation rate and crystal growth rate, as shown in  FIG. 4 .  FIG. 4  depicts an exemplary schematic plot of crystallization under quiescent condition, showing crystal nucleation rate (“R N ”) and the crystal growth rate (“R CG ”) as a function of temperature. The crystal nucleation rate is the rate at which new crystals are formed and the crystal growth rate is the rate of growth of formed crystals. The exemplary curves for R N  and R CG  in  FIG. 4  have a curved bell-type shape that is similar to R N  and R CG  curves for PLLA. The overall rate of quiescent crystallization (“R CO ”) is the sum of curves R N  and R CG . 
     Quiescent crystallization can occur from a polymer melt, which is to be distinguished from crystallization that occurs solely due to polymer deformation. In general, as shown in  FIG. 4 , quiescent crystallization tends to occur in a semi-crystalline polymer at temperatures between Tg and Tm of the polymer. The rate of quiescent crystallization in this range varies with temperature. Near Tg, nucleation rate is relatively high and quiescent crystal growth rate is relatively low; thus, the polymer will tend to form small crystals at these temperatures. Near Tm, nucleation rate is relatively low and quiescent crystal growth rate is relatively high; thus, the polymer will form large crystals at these temperatures. 
     As previously indicated, crystallization also occurs due to deformation of the polymer. Deformation stretches long polymer chains and sometimes results in fibrous crystals generally oriented in a particular direction. Deforming a polymer tube made of PLLA by blow molding at a particular expansion temperature above Tg results in a combination of deformation-induced crystallization and temperature-induce crystallization. 
     As indicated above, the ability of the polymer to deform is dependent on the blow molding temperature (“expansion temperature”) as well as being dependant on the applied internal pressure (“expansion pressure”) and tensile force. As temperature increases above Tg, molecular orientation is more easily induced with applied stress. Also, as temperature approaches Tm, quiescent crystal growth rate increases and quiescent nucleation rate decreases. Thus, it will also be appreciated that the above described blow molding process involves complex interaction of the processing parameters all of which simultaneously affect crystallinity percentage, crystal size, uniformity of crystal distribution, and preferred molecular or crystal orientation. 
     The desired mechanical properties of the stent made from the deformed tube  301  includes high radial strength, high toughness, high modulus, and low recoil upon deployment of the stent. High toughness can be demonstrated by a lower incidence of cracked and/or broken struts upon expansion of the stent to a deployment diameter. 
       FIG. 5  shows another stent pattern  400  illustrated in a planar or flattened view for ease of illustration and clarity. The stent pattern  400  was cut from a tubular precursor construct. Thus, stent pattern  400  actually forms a tubular stent structure, as partially shown in  FIG. 6 , so that line A-A is parallel to the central axis of the stent.  FIG. 6  shows the stent in a fully deployed state. 
     The stent pattern  400  includes various struts  402  oriented in different directions and gaps  403  between the struts. Each gap  403  and the struts  402  immediately surrounding the gap defines a closed cell  404 . At the proximal and distal ends of the stent, a strut  406  includes depressions, blind holes, or through holes adapted to hold a radiopaque marker that allows the position of the stent inside of a patient to be determined. One of the closed cells  404  is shown with cross-hatch lines to illustrate the shape and size of the cells. All the cells  404  have the same size and shape. 
     The pattern  400  is illustrated with a bottom edge  408  and a top edge  410 . On a stent, the bottom edge  408  meets the top edge  410  so that line B-B forms a circle around the stent central axis. In this way, the stent pattern  400  forms sinusoidal hoops or rings  412  that include a group of struts arranged circumferentially. The rings  412  include a series of crests  407  and troughs  409  that alternate with each other. The sinusoidal variation of the rings  412  occurs primarily in the axial direction, not in the radial direction. That is, all points on the outer surface of each ring  412  are at the same or substantially the same radial distance away from the central axis of the stent. 
     Still referring to  FIG. 5 , the rings  412  are connected to each other by another group of struts that have individual lengthwise axes  413  parallel or substantially parallel to line A-A. The rings  412  are capable of being collapsed to a smaller diameter during crimping and expanded to their original diameter or to a larger diameter during deployment in a vessel. 
     The present invention applies to any stent pattern, not just to the pattern shown in  FIGS. 5 and 6 . A stent may have a different number of rings  412  and cells  404  than what is shown. The number and size of rings  412  and cells  404  may vary depending on the desired axial length and the desired deployed diameter of the stent. For example, a diseased segment of a vessel may be relatively small so a stent having a fewer number of rings can be used to treat the diseased segment. 
     Applicant has unexpectedly found that stents cut from a PLLA tube that has been blow molded under certain processing parameter levels demonstrate improved fracture toughness upon deployment while maintaining sufficient flexibility for crimping and delivery and sufficient radial strength to prevent undue recoil. The PLLA tube was made entirely of PLLA. The preferred levels are given below for the blow molding process parameters for a PLLA precursor tubular construct having an initial (before blow molding) crystallinity percentage of up to about 20% and more narrowly from about 5% to about 15%. Applicant believes that the blow molding process parameter levels given blow result in a deformed PLLA tube having a crystallinity percentage below 50% and more narrowly from about 30% to about 40%. 
     In combination with other blow molding process parameters, improved performance in PLLA stents was seen with percent radial expansion (RE %) from about 200% to about 600%, and more narrowly from about 300% to about 500%, and more narrowly at or about 400%. In combination with other blow molding process parameters, Applicant found that when RE % exceeded 600%, there was no significant increase in radial strength while more cracks were found along the axial direction of the stent as a result of use, especially in stents that have aged prior to use. In combination with other blow molding process parameters, Applicant found that when RE % is about 100% or less, the radial strength was too low for a stent having a strut thickness of 0.006 inches, making the stent highly susceptible to fracture during crimping, delivery, and deployment. 
     TABLE I shows the effect of radial expansion on stent functional performance as measured by the number of cracks or broken struts. The stents that were tested had the strut pattern of  FIG. 5 . There were four groups of stents tested. Each group of stents were made from a precursor construct made of PLLA (“PLLA tube”) that had been deformed radially and axially by blow molding. For each group, stents cut from the deformed PLLA tubes were expanded from a crimped state to a deployed (expanded) diameter to simulate what occurs during implantation in a patient. The number of stents with at least one broken struts and the number of strut cracks per stent were noted for deployed diameters of 3.0 mm, 3.5 mm and 4.0 mm. A strut was counted as broken when a crack propagated all the way through the strut. A size criteria was used when counting cracks that did not go all the way through the strut: only cracks that propagated at least 50% of the strut width were counted. Thus, TABLE I shows that for stents made from a 300% radially expanded PLLA tube then deployed to 3.0 mm, the number of cracks satisfying the size criteria ranged from 2 cracks per stent to 39 cracks per stent. For stents made from a 500% radially expanded PLLA tube then deployed to 4.0 mm, three stents exhibited broken struts and the number of cracks satisfying the size criteria ranged from 9 per stent to 30 per stent. 
     
       
         
           
               
               
               
               
             
               
                   
                 TABLE I 
               
             
            
               
                   
                   
               
               
                   
                   
                   
                 Stent deployed to 
               
               
                   
                 Stent deployed to 
                 Stent deployed to 
                 4.0 mm diameter 
               
            
           
           
               
               
               
               
               
            
               
                 Radial 
                 3.0 mm diameter 
                 3.5 mm diameter 
                   
                 # of 
               
            
           
           
               
               
               
               
               
               
               
            
               
                 Expansion 
                 # stents 
                 # of 
                 # stents 
                 # of 
                 # stents 
                 cracks 
               
               
                 of 
                 with 
                 cracks 
                 with 
                 cracks 
                 with 
                 per 
               
               
                 Precursor 
                 broken 
                 per stent 
                 broken 
                 per stent 
                 broken 
                 stent 
               
               
                 Construct 
                 struts 
                 (note 1) 
                 struts 
                 (note 1) 
                 struts 
                 (note 1) 
               
               
                   
               
               
                 300% 
                 0 
                 2 to 39 
                 0 
                 2 to 22 
                 6 
                 1 to 17 
               
               
                 400% 
                 0 
                 0 
                 0 
                 0 
                 0 
                 0 
               
               
                 450% 
                 0 
                 0 to 8  
                 0 
                 0 to 13 
                 0 
                 1 to 6  
               
               
                 500% 
                 1 
                 1 to 23 
                 1 
                 18 to 37  
                 3 
                 9 to 30 
               
               
                   
               
               
                 (note 1) Number of cracks having a size that is at least 50% of the strut width, per stent. 
               
            
           
         
       
     
     TABLE I shows that stents cut from PLLA tubes that were radially expanded to 400% performed best, as this group exhibited no broken struts and no cracks after being deployed, whether deployed to a diameter of 3.0 mm, 3.5 mm, or 4.0 mm. “No cracks” means that there were no cracks of a size that was at least 50% of the strut width. By contrast, radial expansion below 400% (to 300%) and above 400% (to 450% and 500%) resulted in cracks greater than 50% of strut width. Broken struts occurred with radial expansion of 300% and 500%. 
     When the number of broken struts is weighted more than the number cracks, the column with the worst performance corresponds to stents deployed to 4.0 mm diameter. Notably within in this column, stents formed from PLLA tubes radially expanded to 400% exhibited no broken struts and no cracks of a size greater than 50% of strut width. 
     We turn now to the axial extension processing parameter. In combination with other blow molding process parameters, improved performance in PLLA stents was seen with percent axial extension (AE %) from about 10% to about 400%, and more narrowly from about 20% to about 200%, and more narrowly from about 20% to about 70%, and more narrowly at about 20%. In combination with other blow molding process parameters, Applicant found that when AE % is about 100% or more, the stent exhibited more cracks and broken struts along the circumferential direction during stent deployment. 
     The selected level for AE % may depend on the degree of axial orientation that is already present in an extruded tube that is used as the polymer precursor construct. As previously indicated, a significant amount of axial orientation may already be induced in the precursor construct as a result of extrusion and draw down. In combination with other blow molding process parameters, Applicant has unexpectedly found improved stent functionality when the stent is formed from an extruded tube subjected to AE % of about 20% to 70% during blow molding, wherein prior to blow molding the tube extrusion process used a draw down ratio in the range of about 8:1 to about 2:1, more narrowly from about 7:1 to about 3:1, and more narrowly about 7:1. 
     As previously indicated, the stent is subject to deformation during stent deployment. Some portions of the stent are stretched while other portions of the stent are compressed. Deformation during stent deployment is believed to occur mostly in the circumferential direction, though some deformation also occurs in the axial direction and in directions other than axial and circumferential. Therefore, Applicant believes that at least some axial orientation of polymer molecule chains is desirable. In one study, axial extension of the precursor construct was varied from 0% to 300%. Many cracks and broken struts were observed after deployment of stents made from a precursor construct that was axially expanded above 100%. Above 100%, the incidence of cracks and broken struts generally increased proportionally with greater axial extension. A lower incidence of cracks and broken struts was observed with axial extension in the range of about 20% to about 70%. 
     We turn next to the tensile force processing parameter. In combination with other blow molding process parameters, improved performance in PLLA stents was seen with a tensile force corresponding to about 84 grams applied to one end of the tube during blow molding. 
     We turn now to the propagation rate processing parameter, which corresponds to the rate at which a deforming section of the polymer tube travels along the length of the polymer tube, and may also correspond to the rate at which heating nozzles are linearly translated across the mold. In combination with other blow molding process parameters, improved performance in PLLA stents was seen with an axial propagation rate no greater than about 0.3 mm/minute compared to rates from about 0.6 mm/minute to about 2 mm/minute. 
     In combination with other blow molding process parameters, improved performance in PLLA stents was seen with an expansion pressure in the tubular construct in the blow mold at a gauge pressure of about 130 pounds per square inch (psi) or less, and more narrowly in the range of about 110 psi to about 130 psi. In combination with the other blow molding process parameters, an expansion pressure below 70 psi is often insufficient to expand the polymer tube, while an expansion pressure above 180 psi may produce air bubbles in the polymer. Air bubbles are believed to increase the incidence of broken struts and cracks. 
     Next we turn to the expansion temperature processing parameter. In combination with other blow molding process parameters, improved performance in PLLA stents was seen with an expansion temperature between about 160 deg F. to about 220 deg F., and more narrowly between about 160 deg F. and 190 deg F., and more narrowly between about 170 deg F. and about 180 deg F., and more narrowly at about 170 deg F. 
     In some embodiments, the expansion temperature is at a selected level above Tg of the polymer of the tubular construct in the blow mold. As with other polymers, Tg for PLLA may vary depending on the processing history of the polymer. For PLLA, Tg may range from 122 deg F. to 176 deg F. (50 deg. C. to 80 deg. C.) and, more narrowly, between about 136 deg F. to about 140 deg F. (58 deg. C. to about 60 deg. C.). In combination with other blow molding process parameters, improved performance in PLLA stents was seen with an expansion temperature that is between 20 to 50 deg. C. above Tg, and more narrowly at or about 20 deg. C. above Tg. 
     A precursor construct may also be made from other polymers, such as poly(lactic-co-glycolic acid) (“PLGA”). PLGA is a copolymer of LLA and GA. When the proportion of GA is increased, the maximum crystallinity of PLGA decreases and the degradation rate increases. Different forms of PLGA may be used in a precursor construct for a stent. The different forms may be identified with regard to the selected monomer ratio. The precursor construct can be made from PLGA including any molar ratio of L-lactide (LLA) to glycolide (GA). For example, without limitation, the precursor construct can be made from PLGA with a molar ratio of (LA:GA) including 85:15 (or a range of 82:18 to 88:12), 95:5 (or a range of 93:7 to 97:3), or commercially available PLGA products identified as having these molar ratios. Tg for various forms of PLGA ranges from about 104 deg F. to 140 deg F. (40 deg C. to 60 deg. C.). 
     For PLGA with a molar ratio (LA:GA) of 85:15, Tm is about 40 deg. C. lower than that of PLLA, so PLGA 85:15 can be extruded to form a precursor tube at about 20 deg. C. to about 40 deg. C. lower than the extrusion temperature for PLLA. Also, Tg for PLGA 85:15 is about 10 deg. C. lower than that of PLLA, so a precursor tube made of PLGA 85:15 can normally be expanded at a relatively low lower expansion pressure (i.e., process pressure) of about 110 psi. For PLGA 85:15, an axial propagation rate no greater than about 0.3 mm/minute is preferred. The axial propagation rate corresponds to the speed at which heat sources or nozzles are linearly translated over a blow mold containing the precursor tube. 
     It is contemplated that alternative polymers formulations, such as PLLA-based bioabsorbable copolymers or blends containing rigid and soft segments, might have less stiffness and better toughness. Examples for the rigid segment include without limitation PLA and PLGA. Examples for the soft segment include without limitation polycaprolactone (“PCL”) and polytrimethylcarbonate (“PTMC”). An example of a PLLA-based bioabsorbable copolymer containing rigid and soft segments is, without limitation, poly(L-lactide-co-caprolactone) copolymer. An examples of a PLLA-based bioabsorbable blend containing rigid and soft segments is poly poly(L-lactide)/poly(L-lactide)-block-polycaprolactone. A precursor tube made from any one or a combination of these alternative polymer formulations may be processed in the same manner as described above for a PLLA precursor tube. For example, and not limitation, expansion temperature during blow molding can be between 20 to 50 deg. C. above Tg, and more narrowly at or about 20 deg. C. above Tg of the polymer formulation. Deformation of a precursor tube made from any one or a combination of these alternative polymer formulations can involve any one or any combination of the following process steps: 
     (a) maintaining fluid pressure in the precursor tube at a process pressure from about 50 psi to about 200 psi, or more narrowly in the range of about 75 psi to about 175 psi, or more narrowly in the range of about 100 psi to about 150 psi, or in the range of about 110 psi to about 130 psi, or in the range of about 50 psi to about 75 psi, or in the range of about 75 psi to about 100 psi, or in the range of about 100 psi to about 125 psi, or in the range of about 125 psi to about 150 psi, or in the range of about 150 psi to about 175 psi, or in the range of about 175 psi to about 200 psi; 
     (b) heating the precursor tube to a process temperature from about 100 deg F. to about 300 deg F., more narrowly in the range of about 125 deg F. to about 275 deg F., or in the range of about 150 deg F. to about 250 deg F., or in the range of about 160 deg F. to about 220 deg F., or in the range of about 100 deg F. to about 150 deg F., or in the range of about 150 deg F. to about 200 deg F., or in the range of about 200 deg F. to about 250 deg. F, or in the range of about 250 deg F. to about 300 deg F.; 
     (c) radially expanding the precursor tube during the maintaining of fluid pressure and the heating, the radial expansion being according to a radial expansion ratio between about 100% and about 600%, or in the range of about 150% to about 550%, or in the range of about 200% to about 500%, or in the range of about 250% to about 500%, or in the range of about 300% to about 450%, or in the range of about 100% to about 200%, or in the range of about 200% to about 300%, or in the range of about 300% to about 400%, or in the range of about 400% to about 500%, or in the range of about 500% to about 600%; 
     (d) axially extending the precursor tube during the maintaining of fluid pressure and the heating, the axial extension being according to an axial extension ratio from about 10% to about 200%, or from about 15% to about 150%, or from about 18% to about 120%, or from about 20% to about 100%, or in the range of about 10% to about 50%, or in the range of about 50% to 100%, or in the range of about 100% to about 150%, or in the range of about 150% to about 200%; 
     (e) heating the precursor tube may include heating a tubular mold containing the precursor tube, the heating including moving a heat source disposed outside the precursor tube at a linear rate of movement parallel to the central axis of the mold, the linear rate of movement being from about 0.05 mm per minute to about 1.5 mm per minute, or from about 0.07 mm per minute to about 1.0 mm per minute, or from about 0. 1 mm per minute to about 0.7 mm per minute, or in the range of about 0.1 mm per minute to about 0.3 mm per minute, or in the range of about 0.3 mm per minute to about 0.6 mm per minute; and 
     (f) heating the precursor tube may further include applying a load to an end of the precursor tube during the maintaining of fluid pressure and the heating, the load being from about 20 grams to 200 grams, or from about 40 grams to about 175 grams, or from about 50 grams to about 150 grams, or in the range of about 20 grams to about 50 grams, or in the range of about 50 grams to about 100 grams, or in the range of about 100 grams to about 150 grams, or in the range of about 150 grams to about 200 grams. 
     While several particular forms of the invention have been illustrated and described, it will also be apparent that various modifications can be made without departing from the scope of the invention. It is also contemplated that various combinations or subcombinations of the specific features and aspects of the disclosed embodiments can be combined with or substituted for one another in order to form varying modes of the invention. Accordingly, it is not intended that the invention be limited, except as by the appended claims.