Patent Publication Number: US-2022226133-A1

Title: Balloon expanded polymer stent

Description:
PRIORITY CLAIM 
     This application is a continuation of Ser. No. 16/172,599, filed Oct. 26, 2018, which is a continuation of Ser. No. 15/699,938, filed Sep. 8, 2017, which is a continuation of application Ser. No. 14/085,716, filed on Nov. 20, 2013, which is a continuation of application Ser. No. 14/042,512, filed on Sep. 30, 2013, which is a continuation of application Ser. No. 13/015,488, filed on Jan. 27, 2011, which claimed the benefit of U.S. provisional application no. 61/385,891 filed on Sep. 23, 2010, U.S. provisional application no. 61/385,902 filed Sep. 23, 2010 and U.S. provisional application no. 61/299,968 filed on Jan. 30, 2010. application Ser. Nos. 14/085,716, 14/042,512, and 13/015,488 and provisional applications 61/385,891, 61/385,902, and 61/299,968 are incorporated herein by reference. 
    
    
     FIELD OF THE INVENTION 
     The present invention relates to drug-eluting medical devices; more particularly, this invention relates to polymeric scaffolds that are expanded by a delivery balloon. 
     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 (one example of a stent is found in U.S. Pat. No. 6,066,167 to Lau et al). 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 basic, functional requirements. The stent must be capable of withstanding the structural loads, for example, 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. After deployment, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it. In particular, the stent must adequately maintain a vessel at a prescribed diameter for a desired treatment time despite these forces. The treatment time may correspond to the time required for the vessel walls to remodel, after which the stent is no longer necessary for the vessel to maintain a desired diameter. 
     Radial strength, which is the ability of a stent to resist radial compressive forces, relates to a stent&#39;s radial yield strength and radial stiffness around a circumferential direction of the stent. A stent&#39;s “radial yield strength” or “radial strength” (for purposes of this application) may be understood as the compressive loading, which if exceeded, creates a yield stress condition resulting in the stent diameter not returning to its unloaded diameter, i.e., there is irrecoverable deformation of the stent. When the radial yield strength is exceeded the stent is expected to yield more severely and only a minimal force is required to cause major deformation. 
     Even before the radial yield strength is exceeded there may be permanent deformation in the stent a following radial compressive load, but this degree of permanent deformation somewhere in the stent is not severe enough to have a significant effect on the stent&#39;s overall ability to radially support a vessel. Therefore, in some cases the art may view “radial yield strength” as the maximum radial loading, beyond which the scaffold stiffness changes dramatically. “Radial yield strength” units are sometimes force-divided-by-length, which is an expression of radial yield strength on a per-unit-length basis. Thus, for a radial yield strength per unit length, e.g., F N/mm, the radial load which, if it exceeds this value, would result in significant change in stiffness for a stent having two different lengths, L 1  and L 2 , would therefore be the product F*L 1  and F*L 2 , respectively. The value F, however, is the same in both cases, so that a convenient expression can be used to appreciate the radial yield strength independent of the length of the stent. Typically, the radial force that identifies the point where stiffness is lost does not change much on a per-unit-length basis when the stent length changes. 
     Stents implanted in coronary arteries are primarily subjected to radial loads, typically cyclic in nature, which are due to the periodic contraction and expansion of vessels as blood is pumped to and from a beating heart. Stents implanted in peripheral blood vessels, or blood vessels outside the coronary arteries, e.g., iliac, femoral, popliteal, renal and subclavian arteries, however, must be capable of sustaining both radial forces and crushing or pinching loads. These stent types are implanted in vessels that are closer to the surface of the body. Because these stents are close to the surface of the body, they are particularly vulnerable to crushing or pinching loads, which can partially or completely collapse the stent and thereby block fluid flow in the vessel. 
     As compared to a coronary stent, which is limited to radial loads, a peripheral stent must take into account the significant differences between pinching or crushing loads and radial loads, as documented in Duerig, Tolomeo, Wholey,  Overview of superelastic stent Design , Min Invas Ther &amp; Allied Technol 9(3/4), pp. 235-246 (2000) and Stoeckel, Pelton, Duerig,  Self - Expanding Nitinol Stents—Material and Design Considerations , European Radiology (2003). The corresponding crushing and radial stiffness properties of the stent also can vary dramatically. As such, a stent that possesses a certain degree of radial stiffness does not, generally speaking, also indicate the degree of pinching stiffness possessed by the stent. The two stiffness properties are not the same, or even similar. 
     The amount of cross-sectional crush expected for a peripheral stent implanted within the femoral artery has been estimated to be about 5.8+/−7%, 6.5+/−4.9% and 5.1+/−6.4% at the top, middle and bottom portions of the femoral artery in older patients and 2.5+/−7.7%, −0.8+/−9.4% and −1.5+/−10.5% for younger patients. Other considerations for peripheral stents are the degree of bending and axial compression the stent can withstand without mechanical loss of strength/stiffness. As compared to coronary stents, a peripheral stent usually has lengths of between about 36 and 40 mm when implanted in the superficial femoral artery, as an example. As such, the stent must be flexible enough to withstand axial compression and bending loading without failure. The amount of bending and axial compression expected has been studied and reported in Nikanorov, Alexander, M. D. et al.,  Assessment of self - expanding Nitinol stent deformation after chronic implantation into the superficial femoral artery.    
     To date the most commonly used type of peripheral stent are self-expanding stents made from super-elastic material, such as Nitinol. This type of material is known for its ability to return to its original configuration after severe deformation, such as a crushing load or longitudinal bending. However, this variety of self-expanding stents have undesired qualities; most notably, the high resiliency of super-elastic material produces what is commonly referred to as a “chronic outward force” (COF) on the blood vessel supported by the stent. Complications resulting from COF are discussed in Schwartz, Lewis B. et al.  Does Stent Placement have a learning curve: what mistakes do we as operators have to make and how can they be avoided ?, Abbott Laboratories; Abbott Park, Ill., USA. It is believed that a COF exerted on a blood vessel by a self-expending stent is a main contributor to high degrees of restenosis of lesions treated by the self-expanding stent. It has been shown that not even an anti-proliferative drug delivered from drug eluting self-expandable stents can mitigate the restenosis caused by the stent&#39;s COF. 
     Stents that are plastically deformed by a balloon to support a vessel do not suffer from this drawback. Indeed, balloon expanded stents, in contrast to self-expanding stents made from a super-elastic material, have the desirable quality of being deployable to the desired diameter for supporting the vessel without exerting residual outward forces on the vessel. However, the prior art has concluded that plastically deformed stents, once collapsed, pinched or crushed in a peripheral artery will remain so, permanently blocking the vessel. The prior art has concluded, therefore, that plastically deformed stents pose an undesirable condition to the patient and should not be used to treat peripheral blood vessels. 
     A polymer scaffold, such as that described in US 2010/0004735 is made from a biodegradable, bioabsorbable, bioresorbable, or bioerodable polymer. The terms biodegradable, bioabsorbable, bioresorbable, biosoluble or bioerodable refer to the property of a material or stent to degrade, absorb, resorb, or erode away from an implant site. The polymer scaffold described in US 2010/0004735, as opposed to a metal stent, is intended to remain in the body for only a limited period of time. The scaffold is made from 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. Moreover, it is believed that biodegradable scaffolds allow for improved healing of the anatomical lumen as compared to metal stents, which may lead to a reduced incidence of late stage thrombosis. In these cases, there is a desire to treat a vessel using a polymer scaffold, in particular a bioerodible polymer scaffold, as opposed to a metal stent, so that the prosthesis&#39;s presence in the vessel is for a limited duration. However, there are numerous challenges to overcome when developing a polymer scaffold. 
     The art recognizes a variety of factors that affect a polymeric scaffold&#39;s ability to retain its structural integrity and/or shape when subjected to external loadings, such as crimping and balloon expansion forces. These interactions are complex and the mechanisms of action not fully understood. According to the art, characteristics differentiating a polymeric, bio-absorbable scaffold of the type expanded to a deployed state by plastic deformation from a similarly functioning metal scaffold are many and significant. Indeed, several of the accepted analytic or empirical methods/models used to predict the behavior of metallic scaffolds tend to be unreliable, if not inappropriate, as methods/models for reliably and consistently predicting the highly non-linear, time dependent behavior of a polymeric load-bearing structure of a balloon-expandable scaffold. The models are not generally capable of providing an acceptable degree of certainty required for purposes of implanting the scaffold within a body, or predicting/anticipating the empirical data. 
     Moreover, it is recognized that the state of the art in medical device-related balloon fabrication, e.g., non-compliant balloons for scaffold deployment and/or angioplasty, provide only limited information about how a polymeric material might behave when used to support a lumen within a living being via plastic deformation of a network of rings interconnected by struts. In short, methods devised to improve mechanical features of an inflated, thin-walled balloon structure, most analogous to mechanical properties of a pre-loaded membrane when the balloon is inflated and supporting a lumen, simply provides little, if any insight into the behavior of a deployed polymeric scaffold. One difference, for example, is the propensity for fracture or cracks to develop in a polymer scaffold. The art recognizes the mechanical problem as too different to provide helpful insights, therefore, despite a shared similarity in class of material. At best, the balloon fabrication art provides only general guidance for one seeking to improve characteristics of a balloon-expanded, bio-absorbable polymeric scaffold. 
     Polymer material considered for use as a polymeric scaffold, e.g. poly(L-lactide) (“PLLA”), poly(L-lactide-co-glycolide) (“PLGA”), poly(D-lactide-co-glycolide) or poly(L-lactide-co-D-lactide) (“PLLA-co-PDLA”) with less than 10% D-lactide, and PLLD/PDLA stereo complex, may be described, through comparison with a metallic material used to form a stent, in some of the following ways. A suitable polymer has a low strength to weight ratio, which means more material is needed to provide an equivalent mechanical property to that of a metal. Therefore, struts must be made thicker and wider to have the required strength for a stent to support lumen walls at a desired radius. The scaffold made from such polymers also tends to be brittle or have limited fracture toughness. The anisotropic and rate-dependant inelastic properties (i.e., strength/stiffness of the material varies depending upon the rate at which the material is deformed) inherent in the material, only compound this complexity in working with a polymer, particularly, bio-absorbable polymer such as PLLA or PLGA. 
     Processing steps performed on, and design changes made to a metal stent that have not typically raised concerns for, or required careful attention to unanticipated changes in the average mechanical properties of the material, therefore, may not also apply to a polymer scaffold due to the non-linear and sometimes unpredictable nature of the mechanical properties of the polymer under a similar loading condition. It is sometimes the case that one needs to undertake extensive validation before it even becomes possible to predict more generally whether a particular condition is due to one factor or another—e.g., was a defect the result of one or more steps of a fabrication process, or one or more steps in a process that takes place after scaffold fabrication, e.g., crimping? As a consequence, a change to a fabrication process, post-fabrication process or even relatively minor changes to a scaffold pattern design must, generally speaking, be investigated more thoroughly than if a metallic material were used instead of the polymer. It follows, therefore, that when choosing among different polymeric scaffold designs for improvement thereof, there are far less inferences, theories, or systematic methods of discovery available, as a tool for steering one clear of unproductive paths, and towards more productive paths for improvement, than when making changes in a metal stent. 
     The present inventors recognize, therefore, that, whereas inferences previously accepted in the art for stent validation or feasibility when an isotropic and ductile metallic material was used, those inferences would be inappropriate for a polymeric scaffold. A change in a polymeric scaffold pattern may affect not only the stiffness or lumen coverage of the scaffold in its deployed state supporting a lumen, but also the propensity for fractures to develop when the scaffold is crimped or being deployed. This means that, in comparison to a metallic stent, there is generally no assumption that can be made as to whether a changed scaffold pattern may not produce an adverse outcome, or require a significant change in a processing step (e.g., tube forming, laser cutting, crimping, etc.). Simply put, the highly favorable, inherent properties of a metal (generally invariant stress/strain properties with respect to the rate of deformation or the direction of loading, and the material&#39;s ductile nature), which simplify the stent fabrication process, allow for inferences to be more easily drawn between a changed stent pattern and/or a processing step and the ability for the stent to be reliably manufactured with the new pattern and without defects when implanted within a living being. 
     A change in the pattern of the struts and rings of a polymeric scaffold that is plastically deformed, both when crimped to, and when later deployed by a balloon, unfortunately, is not predictable to the same or similar degree as for a metal stent. Indeed, it is recognized that unexpected problems may arise in polymer scaffold fabrication steps as a result of a changed pattern that would not have necessitated any changes if the pattern was instead formed from a metal tube. In contrast to changes in a metallic stent pattern, a change in polymer scaffold pattern may necessitate other modifications in fabrication steps or post-fabrication processing, such as crimping and sterilization. 
     In addition to meeting the requirements described above, it is desirable for a scaffold to be radiopaque, or fluoroscopically visible under x-rays. Accurate placement is facilitated by real time visualization of the delivery of a scaffold. A cardiologist or interventional radiologist can track the delivery catheter through the patient&#39;s vasculature and precisely place the scaffold at the site of a lesion. This is typically accomplished by fluoroscopy or similar x-ray visualization procedures. For a scaffold to be fluoroscopically visible it must be more absorptive of x-rays than the surrounding tissue. Radiopaque materials in a scaffold may allow for its direct visualization. However, a significant shortcoming of a biodegradable polymer scaffold (and polymers generally composed of carbon, hydrogen, oxygen, and nitrogen) is that they are radiolucent with no radiopacity. Biodegradable polymers tend to have x-ray absorption similar to body tissue. One way of addressing this problem is to attach radiopaque markers to structural elements of the stent. A radiopaque marker can be disposed within a structural element in such a way that the marker is secured to the structural element. However, the use of stent markers on polymeric stents entails a number of challenges. One challenge relates to the difficulty of insertion of markers. These and related difficulties are discussed in US 2007/0156230. 
     There is a need to develop a prosthesis for treating peripheral blood vessels that possesses the desirable qualities of a balloon expanded stent, which does not exert residual outward forces on the vessel (as in the case of a self-expanding stent) while, at the same time, being sufficiently resilient to recover from a pinching or crushing load in a peripheral blood vessel, in addition to the other loading events expected within a peripheral blood vessel that are not typically experienced by a coronary scaffold. There is also a need to fabricate such a polymer scaffold so that the prosthesis also is capable of possessing at least a minimum radial strength and stiffness required to support a peripheral blood vessel; a low crossing profile; and a limited presence in the blood vessel. There is also a need for a scaffold that is easily monitored during its pendency using standard imaging techniques, and is capable of high yield production. 
     SUMMARY OF THE INVENTION 
     The invention provides a polymer scaffold suited to address the foregoing needs including high crush recoverability, e.g., at least about 90-95% after a 50% crushing load. The scaffold is cut from a polymer tube and crimped to a balloon. Accordingly, the invention provides a balloon expandable, plastically deformed scaffold cut from a tube and being suitable for use as a peripheral scaffold. As such, the drawbacks of self-expanding stents can be obviated by practicing the invention. 
     To date the art has relied on metals or alloys for support and treatment of peripheral blood vessels. As mentioned earlier, once a metallic stent is implanted it remains in the body permanently, which is not desired. A scaffold made from a material that dissolves after it treats an occluded vessel, therefore, would be preferred over a metal stent. A polymer, however, is much softer than a metal. If it will serve as a replacement to metal, a new design approach is needed. 
     High radial force, small crimped profile and crush recovery is needed in the polymer scaffold. If the material cannot be modified enough to meet these needs, then a modification to the design of the scaffold network of struts is required. There are a few known approaches to increase the radial yield strength. One is to increase the wall thickness and another is to increase the strut width. Both of these modifications, however, will result in greater profile of the device at the crimped state. A small crimped profile of the device and increased stiffness and strength is therefore necessary and heretofore not addressed in the art. 
     As will be appreciated, aspects of a polymer scaffold disclosed herein contradict conclusions that have been previously made in the art concerning the suitability of a balloon-expandable stent, or scaffold for use in peripheral blood vessels. The problems concerning self-expanding stents are known. Therefore a replacement is sought. However, the conventional wisdom is that a balloon expanded stent having sufficient radial strength and stiffness, as opposed to a self-expanding stent, is not a suitable replacement, especially in vessels that will impose high bending and/or crushing forces on the implanted prosthesis. 
     According to the invention, crush-recoverable polymer scaffolds possessing a desired radial stiffness and strength, fracture toughness and capability of being crimped down to a target delivery diameter will properly balance three competing design attributes: radial strength/stiffness verses toughness, in-vivo performance verses compactness for delivery to a vessel site, and crush recovery verses radial strength/stiffness. 
     Disclosed herein are embodiments of a scaffold that can effectively balance these competing needs, thereby providing an alternative to prostheses that suffer from chronic outward force. As will be appreciated from the disclosure, various polymer scaffold combinations were fabricated and tested in order to better understand the characteristics of a scaffold that might address at least the following needs: 
     Crush recoverability of the scaffold without sacrificing a desired minimal radial stiffness and strength, recoil, deploy-ability and crimping profile; 
     Acute recoil at deployment—the amount of diameter reduction within ½ hour of deployment by the balloon; 
     Delivery/deployed profile—i.e., the amount the scaffold could be reduced in size during crimping while maintaining structural integrity; 
     In vitro radial yield strength and radial stiffness; 
     Crack formation/propagation/fracture when crimped and expanded by the balloon, or when implanted within a vessel and subjected to a combination of bending, axial crush and radial compressive loads; 
     Uniformity of deployment of scaffold rings when expanded by the balloon; and 
     Pinching/crushing stiffness. 
     Based on these studies, which have included in-vivo animal testing of a peripherally implanted scaffold, the invention provides the following relationships characterizing a polymer scaffold that exhibits the desired characteristics including crush recoverability: 
     a ratio of outer diameter to wall thickness; 
     a ratio of outer diameter to strut width; 
     a ratio of radial stiffness to pinching stiffness; 
     a ratio of pinching stiffness to scaffold diameter; 
     a ratio of radial stiffness to scaffold diameter; 
     a ratio of strut or link thickness to its width; and 
     a ratio of pre-crimp scaffold diameter to strut moment of inertia. 
     Additional relationships characterizing mechanical properties of a scaffold meeting the above needs may be inferred from the disclosure. 
     According to one aspect of the invention, polymer scaffolds having crush recovery and good radial strength and stiffness possess one or more of the following relations between material properties and/or scaffold dimensions. It will be understood that these relationships, as disclosed herein and throughout the disclosure, include previously unknown relationships among scaffold structural properties, material and dimensions that reveal key characteristics of a scaffold needed for a crush-recoverable scaffold uniquely suited to achieve the clinical objective. As such, the invention includes the identification of a particular relationship, e.g., a dimensionless number used in combination with one or more additional scaffold dimensions, e.g., inflated diameter, aspect ratio, crown angle, wall thickness, to produce a crush-recoverable scaffold having the desired stiffness and strength property needed to support the vessel. 
     According to one aspect of the invention a strut forming a ring of the crush recoverable scaffold has an aspect ratio (AR) of between about 0.8 and 1.4. Aspect ratio (AR) is defined as the ratio of cross-sectional width to thickness. Thus for a strut having a width of 0.0116 and a wall thickness of 0.011 the AR is 1.05. 
     According to another aspect of the invention, the links connect rings of the scaffold. The AR of a link may be between about 0.4 and 0.9. 
     According to another aspect of the invention, the AR of both the link and the strut may between about 0.9 and 1.1, or about 1. 
     According to another aspect of the invention, a crush recoverable scaffold is crimped to a delivery balloon of a balloon catheter. The balloon has a maximum expanded diameter less than the diameter of the scaffold before crimping. The scaffold has a pre-crimping diameter of between 7-10 mm, or more narrowly 7-8 mm, and possesses a desired pinching stiffness while retaining at least a 80% recoverability from a 50% crush. 
     According to another aspect of the invention a crush-recoverable scaffold has a desirable pinching stiffness of at least 0.5 N/mm, radial strength of at least 0.3 N/mm and a wall thickness of at least 0.008″, or between about 0.008″ and 0.012″. The scaffold is capable of recovering at least 80% of its diameter after at least an about 30% crush. 
     According to another aspect of invention a 9 mm scaffold (pre-crimp diameter) with wall thickness of between 0.008″ and 0.014″, or more narrowly 0.008″ and 0.011″ providing the desired pinching stiffness while retaining 50% crush recoverability. More generally, it was found that a ratio of pre-crimp or tube diameter to wall thickness of between about 30 and 60, or between about 20 and 45 provided 50% crush recoverability while exhibiting a satisfactory pinching stiffness and radial stiffness. And in some embodiments it was found that a ratio of inflated diameter to wall thickness of between about 25 and 50, or between about 20 and 35. 
     According to another aspect of the disclosure a crush-recoverable scaffold has a desirable pinching stiffness to wall thickness ratio of 0.6-1.8 N/mm 2 . 
     According to another aspect of the disclosure a crush-recoverable scaffold has a desirable pinching stiffness to wall thickness*tube diameter ratio of 0.08-0.18 N/mm 3 . 
     According to another aspect of invention a crush-recoverable scaffold has a ratios of pinching stiffness to radial stiffness of between about 4 to 1, 3 to 1, or more narrowly about 2 to 1; ratios of pinching stiffness to wall thickness of between about 10 to 70, or more narrowly 20 to 50, or still more narrowly between about 25 and 50; and ratios of scaffold inflated diameter to pinching stiffness of between about 15 and 60 or more narrowly between about 20 to 40. 
     According to another aspect of the invention a crush recoverable polymer scaffold has rings comprising 9 or 8 crowns. For a 9 crown pattern and 7-9 mm outer diameter a crown angle is less than 115 degrees and more preferably crown angles between 105 and 95 degrees. For a 8 crown pattern and 7-9 mm outer diameter the angle is about less than 110 degrees. 
     According to another aspect of invention, a crush-recoverable scaffold has a radial strength of greater than about 0.3 N/mm, or between about 0.32 and 0.68 N/mm, and a radial stiffness of greater than about 0.5 N/mm or between about 0.54 N/mm and 1.2 N/mm. The scaffold may have a wall thickness of about 0.008″ to 0.014″ and configured for being deployed by a 6.5 mm non-compliant balloon from about a 2 mm crimped profile, or deployed to a diameter of between about 6.5 mm and 7 mm from about a 2 mm crossing profile on a balloon catheter. The scaffold strut and/or link elements may have an AR of equal to or greater than 1.0. 
     According to another aspect of the invention, a crush-recoverable polymer scaffold recovers greater than 80% of its diameter after being pinched by an amount equal to 50% of its diameter (50% crush) and the pinched state is maintained for 1-5 minutes. 
     According to another aspect of the invention, a crush-recoverable polymer scaffold recovers greater than 90% of its diameter after being pinched to 25% of its diameter (75% crush) and the pinched state is maintained for 1-5 minutes. 
     According to another aspect of the invention, a crush-recoverable polymer scaffold includes a marker structure including a pair of markers arranged circumferentially on a connecting link and spaced from adjacent rings of the scaffold so that the crimped profile is the same with or without the markers. Alternatively, according to another aspect of the invention, a radiopaque foil is wrapped around a link of the scaffold and held in place 
     In another aspect of invention, a polymer scaffold having a wall thickness of between about 0.008″ and 0.014″ and outer diameter of between about 7 mm and 10 mm was capable of meeting the foregoing needs. 
     In another aspect of the invention, a crush recoverable scaffold was crimped from a 7 mm, 8 mm and 9 mm outer diameter to a 2 mm outer diameter and deployed without fracture and/or excessive cracking of struts that are a typical concern when a polymer, especially a brittle polymer like PLLA, is used to form the scaffold structure. 
     A scaffold has a pre-crimp diameter (SD PC ) meaning the diameter of the scaffold before it is crimped to its delivery balloon, and an inflated diameter (SD I ). The scaffold is crimped to the balloon-catheter and intended for delivery to a vessel within the body. The average vessel diameter where the scaffold is to be implanted is VD. SD I  is about 1.2 times greater than VD. For purposes of the disclosure, VD can range from about 5 mm to 10 mm and SD PC  can range between about 6 to 12 mm. According to another aspect of invention: 
       1.1×(VD)≤SD PC ≤1.7×(VD)  (EQ. 1)
 
       1.1×(SD I )×(1.2) −1 ≤SD PC ≤1.7×(SD I )×(1.2) −1   (EQ. 2)
 
     Scaffold satisfying EQS. 1 and 2 can yield a crush-recoverable scaffold having at least 90% recovery after at least a 25% crush, while also having favorable radial stiffness, pinching stiffness, acceptable recoil, radial strength and/or crossing profile. In a preferred embodiment the scaffold is made from PLLA. The partial inequalities in EQS. 1 and 2 are intended to refer to approximate ranges. 
     It is contemplated that a polymer scaffold according to the invention may be used to treat conditions in the Femoral artery, Popliteal artery, Tibial artery, Pudendal artery, Brachial artery, Caroitid artery, Jugular vein, Abdominal arteries and veins. 
     In another aspect of the invention a symmetric, closed cell for a scaffold improves deployment uniformity and reduces fracture problems for a scaffold having crush recoverability. 
     In another aspect of invention a balloon-expandable medical device for being implanted in a peripheral vessel of the body includes a scaffold formed from a polymer tube,—configured for being crimped to a balloon,—the scaffold having a pattern of interconnected elements and—the scaffold having an expanded diameter when expanded from a crimped state by the balloon, wherein the scaffold attains greater than about 90% of its diameter after being crushed by an amount equal to at least 33% of its expanded diameter (33% crush); and wherein the scaffold has a radial stiffness greater than 0.3 N/mm. 
     In another aspect of invention a balloon-expandable medical device for being implanted in a peripheral vessel of the body includes a crimped scaffold that when deployed by a balloon forms a scaffold having an expanded diameter; wherein the scaffold is capable of regaining more than 90% of its diameter after being crushed to at least 75% of its expanded diameter or crushed by an amount equal to at least 25% of its expanded diameter; and wherein the scaffold comprises—a radial stiffness greater than about 0.3 N/mm, and—a radial strength, pinching strength, pinching stiffness and fracture toughness of a pre-crimp scaffold having a pre-crimp diameter between 300-400% greater than a diameter of the crimped scaffold. 
     In another aspect of invention a radially expandable stent includes a balloon expandable scaffold formed from a PLLA tube,—the scaffold including a plurality of radially expandable undulating cylindrical rings of struts, wherein the undulating rings of struts comprise crowns, wherein adjacent rings of struts are connected by longitudinal links, wherein a ring has no more than 9 crowns and 3 links around its circumference, and the angle at any crown is less than 115 degrees;—the scaffold has an outer diameter of 8 to 10 mm; and—the scaffold has a wall thickness at least about 0.008″. 
     In another aspect of invention a peripherally implantable medical device includes a crimped scaffold that when expanded by a balloon forms a scaffold having a diameter;—the scaffold regains more than 90% of the diameter after being crushed to at least 67% of the diameter (or crushed by an amount equal to at least 33% of its expanded diameter),—the scaffold is formed from PLLA,—the scaffold has a diameter to wall thickness ratio of between about 30 and 60;—the scaffold has struts and links, wherein a strut and/or link has a width to thickness ratio of between about 0.8 and 1.4, and—the scaffold has a radial stiffness greater than or equal to about 0.3 N/mm. 
     INCORPORATION BY REFERENCE 
     All publications and patent applications mentioned in this specification are herein incorporated by reference to the same extent as if each individual publication or patent application was specifically and individually indicated to be incorporated by reference, and as if each said individual publication or patent application was fully set forth, including any figures, herein. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a perspective view of a deformed polymer tube. The tube is formed into a scaffold. 
         FIG. 2  is a partial planar view of a scaffold pattern according to a first embodiment of a scaffold. 
         FIG. 3  is a partial perspective view of a scaffold structure. 
         FIG. 4  is a partial planar view of a scaffold pattern according to a second embodiment of a scaffold. 
         FIG. 5A  is a planar view of a portion of the scaffold pattern of  FIG. 4  taken at section VA-VA. 
         FIG. 5B  is a planar view of a portion of the scaffold pattern of  FIG. 2  taken at section VB-VB. 
         FIGS. 6A and 6B  are tables showing examples of scaffold features in accordance with aspects of the disclosure. 
         FIG. 7A-7B  shows a scaffold crown formation in its expanded and crimped states. 
         FIG. 7C-7D  shows a scaffold crown formation in its expanded and crimped states for a scaffold according to the first embodiment. 
         FIG. 7E-7F  shows a scaffold crown formation in its expanded and crimped states for a scaffold according to an alternative embodiment. 
         FIGS. 8B, 8C and 8D  are scanning electron microscope (SEM) photographs of scaffold crowns. The crowns have an inner radius of about 0.00025 inches. The photographs are taken after the scaffold was expanded by a balloon. 
         FIGS. 8A, 8F and 8G  are scanning electron microscope (SEM) photographs of scaffold crowns having an inner radius substantially higher than the inner radius of the scaffold crowns in  FIGS. 8B, 8C and 8D . The photographs are taken after the scaffold was expanded by a balloon.  FIG. 8E  is another scanning electron microscope (SEM) photograph of a scaffold crown. 
         FIGS. 9A-9B  show the first embodiment of a scaffold including a radiopaque marker structure formed on a link connecting rings.  FIG. 9A  shows the expanded configuration and  FIG. 9B  shows the location of the radiopaque markers relative to folded struts of the scaffold rings in the crimped configuration. 
         FIGS. 10A-10B  show an alternative embodiment of a scaffold including a radiopaque marker disposed on a link connecting rings.  FIG. 10A  shows the expanded configuration and  FIG. 10B  shows the location of the radiopaque marker relative to folded struts of the scaffold rings in the crimped configuration. 
         FIGS. 11A-11E  are several alternative embodiments of a scaffold including a radiopaque marker. For these embodiments the radiopaque marker(s) are located on or near the crown of a crown, as opposed to on a link connecting rings.  FIGS. 11A, 11B and 11E  depict examples of locations for a cylindrical marker while  FIGS. 11C and 11D  depict locations for a strip of marker material. 
         FIG. 11F  depicts an alternative embodiment of a scaffold having a radiopaque marker. In this example the radiopacity is provided through material used to strengthen the crown at an end ring. As such, the embodiments provide more visibility at the end ring while also strengthening the end ring. 
         FIGS. 12A, 12B and 12C  are diagrams describing a relationship between crush recoverability and wall thickness for a scaffold.  FIG. 12A  shows a cross-section of a scaffold in its un-deformed (unloaded) state and deformed state when subjected to a pinching load (drawn in phantom).  FIGS. 12B-12C  are models of equivalent half-cylinder shells of different thickness to show the effects of wall thickness on crush-recoverability when a scaffold is subject a pinching load. 
         FIG. 13  is a plot showing the crush-recovery for a scaffold after a 50% crush. The plot shows the percentage recovered over a 24 hour period following a brief, 1 minute and 5 minute crush at 50% crush. 
         FIGS. 14A and 14B  are partial planar views of a scaffold pattern according to an alternate embodiment of a scaffold including a first embodiment of a weakened or flexible link element connecting rings. 
         FIG. 14C  is a second embodiment of a weakened or flexible link element connecting rings of the scaffold. 
         FIGS. 14D and 14F  shows an alternate embodiment of a weakened portion of a link connecting rings.  FIG. 14D  shows an asymmetric weakened link portion and  FIG. 14F  shows a symmetric weakened link portion. 
         FIG. 14E  shows an example of a link structure where voids are formed in the link to create a point of fracture for the link at the voids. 
         FIG. 15  is a partial planar view of a scaffold pattern according to an alternate ring structure for a scaffold where the ring structure has curved struts extending between crowns. 
         FIGS. 16-23  are plots showing results from a first animal study for an implanted scaffold at 30, 90 and 180 days following implantation. The scaffold performance is compared to a self-expanding metal stent implanted within the same animal. 
         FIGS. 24-26  are plots showing results from a second animal study comparing the performance of scaffold having different wall thickness. 
     
    
    
     DETAILED DESCRIPTION OF EMBODIMENTS 
     The disclosure proceeds as follows. First, definitions of terms that may be used during the course of the subsequent disclosure are explained. Embodiments of processes for forming a deformed polymer tube from a precursor are provided. According to the disclosure, the crush recoverable and balloon expandable scaffold is cut from a tube ( FIG. 1 ) formed through a process intended to enhance mechanical properties of the scaffold including fracture toughness. Discussion of the scaffold patterns according to several embodiments are discussed next. Examples of the scaffold patterns are provided. During this discussion, reference is made to aspects of a scaffold found to play an important role in the stiffness, strength, crimping and deployment of a polymer scaffold, as well as other properties as they relate to crush recoverability of a load-bearing polymer structure. Included herein are aspects of the scaffold that are contrary and, in some cases, surprising and unexpected, particularly when compared to aspects of a comparable, peripheral metal stent having a similar pattern of struts. Finally, bench and in-vivo test results are discussed, including exemplary examples of embodiments of invention and explanation of the results observed and problems overcome. In these examples there may be gained a further appreciation of aspects of invention—a crush recoverable and balloon-expandable polymer scaffold having desirable radial strength and stiffness properties and capable of being crimped to a diameter suitable for delivery through a blood vessel via a balloon catheter. 
     For purposes of this disclosure, the following terms and definitions apply: 
     “Inflated diameter” or “expanded diameter” refers to the maximum diameter the scaffold attains when its supporting balloon is inflated to expand the scaffold from its crimped configuration to implant the scaffold within a vessel. The inflated diameter may refer to a post-dilation diameter which is beyond the nominal balloon diameter, e.g., a 6.5 mm semi-compliant PEBAX balloon has about a 7.4 mm post-dilation diameter. The scaffold diameter, after attaining its inflated diameter by balloon pressure, will to some degree decrease in diameter due to recoil effects and/or compressive forces imposed by the wall of the vessel after the balloon is removed. For instance, referring to an expansion of the V59 scaffold having the properties in Table 6B, when placed on a 6.5 mm PEBAX balloon and the balloon is expanded to a post-dilation condition outside a vessel, the scaffold inner diameter will be about 7.4 mm and about (0.955)×(7.4 mm) before and after, respectively, acute-recoil has occurred. The inflated diameter may be about 1.2 times the average vessel diameter and peripheral vessel sizes typically range from about 4 to 10 mm for purposes of this disclosure. 
     “Theoretical minimum diameter” means the smallest diameter for a scaffold based on its geometry of strut lengths, thickness and widths. A “theoretical minimum diameter” is not defined in terms of a minimum crimped profile for a scaffold or stent that can be later deployed and work properly as a balloon-expanded prosthesis. Rather, it is only a definition defined by the geometry, or minimum volume of space that a device can occupy following a uniform reduction in diameter. As a formula, the “theoretical minimum diameter” (Dmin) may be expressed as follows: 
       Dmin=(ΣSwi+ΣCrj+ΣLwk)*(1/π)+2*WT  (EQ. 3)
 
     Where the quantities above are taken from a cross-sectional slice of the scaffold, 
     ΣSwi (i=1 . . . n) is the sum of n ring struts having width Swi; 
     ΣCrj (j=1 . . . m) is the sum of m crown inner radii having radii Crj (times 2); 
     ΣLwk (k=1 . . . p) is the sum of p links having width Lwk; and 
     WT is the scaffold wall thickness. 
     EQ. 3 assumes the width for a folded pair of struts, e.g., struts  420 ,  422  in  FIG. 7A , is the same whether measured near the crown  410  or the strut mid width. When the crown is built up more, so that the width is wider there than ring strut mid-width, Swi would be measured by the width at the crown. Also, the minimum space between struts is defined by twice the inner radius of the adjacent crown (or valley), i.e., Crj. 
     For the scaffold dimensions of  FIG. 6B  the crown width is wider than the strut mid-width. Therefore, using EQ. 3 Dmin is [16*(0.013)+12*(0.0005)+4*(0.0115)] *(1/π)+2*(0.011)=0.1048″ or 2.662 mm (minimum diameter computed at cross-section passing through crowns). If, instead the cross-section were taken at the strut mid width (0.0116 instead of 0.013) EQ. 3 gives 0.0976″ or 2.479 mm. 
     It should be noted that EQ. 3 assumes the struts have essentially a square cross-section. This is the case for the scaffold of  FIG. 6B  (strut cross-sectional dimension at the crown is 0.011×0.013). For a scaffold having struts with a trapezoidal cross section, e.g., a scaffold cut from a smaller diameter so that the ratio of wall thickness to outer diameter is much higher than in the case of  FIG. 1 , a more accurate approximation for Dmin would be (ΣSwi+ΣCrj+ΣLwk)*(1/π) since the edges of the struts at the outer surface would abut at Dmin before the surfaces extending over the thickness of a strut abut each other. 
     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 result from the applied force. For example, a material has both a tensile and a compressive modulus. 
     “Toughness”, or “fracture 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 and is sometimes used to describe a circumferential property, i.e radial strength. 
     The term “crush recovery” is used to describe how the scaffold recovers from a pinch or crush load, while the term “crush resistance” is used to describe the force required to cause a permanent deformation of a scaffold. A scaffold or stent that does not possess good crush recovery does not substantially return to its original diameter following removal of a crushing force. As noted earlier, a scaffold or stent having a desired radial force can have an unacceptable crush recovery. And a scaffold or stent having a desired crush recovery can have an unacceptable radial force. 
     The polymer scaffold illustrated in  FIG. 2  is formed from a poly(L-lactide) (“PLLA”) tube. The process for forming this PLLA tube may be the process described in U.S. patent application Ser. No. 12/558,105. Reference is made to a precursor that is “deformed” in order to produce the tube of  FIG. 1  having the desired scaffold diameter, thickness and material properties as set forth below. Before the tube is deformed or, in some embodiments, expanded to produce the desired properties in the starting tube for the scaffold, the precursor is formed. The precursor may be formed by an extrusion process which starts with raw PLLA resin material heated above the melt temperature of the polymer which is then extruded through a die. Then, in one example, an expansion process for forming an expanded PLLA tube includes heating a PLLA precursor above the PLLA glass transition temperature (i.e., 60-70 degrees C.) but below the melt temperature (165-175 degrees C.), e.g., around 110-120 degrees C. 
     A precursor tube is deformed in radial and axial directions by a blow molding process wherein deformation occurs progressively at a predetermined longitudinal speed along the longitudinal axis of the tube. As explained below, the deformation improves the mechanical properties of the tube before it is formed into the scaffold of  FIG. 2 . The tube deformation process is intended to orient polymer chains in radial and/or biaxial directions. The orientation or deformation causing re-alignment is performed according to a precise selection of processing parameters, e.g. pressure, heat (i.e., temperature), deformation rate, to affect material crystallinity and type of crystalline formation during the deformation process. 
     In an alternative embodiment the tube may be made of poly(L-lactide-co-glycolide), poly(D-lactide-co-glycolide) (“PLGA”), polycaprolactone (“PCL”), any semi-crystalline copolymers combining any of these monomers, or any blends of these polymers. Material choices for the scaffold should take into consideration the complex loading environment associated with many peripheral vessel locations, particularly those located close to limbs. 
     The femoral artery provides a dynamic environment for vascular implants as various forces may crush, twist, extend, or shorten the device simultaneously. The force application may vary between point load to distributed load or a combination thereof and also as a function of time. Recent results have shown that bioresorbable scaffolds made from highly crystalline PLLA can provide crush recovery without causing a permanent and constant outward radial force on the vessel. The permanent and constant outward radial force may be the cause of late clinical issues with nitinol self-expandable stents. However, a remaining challenge with bioresorbable scaffolds is to make them optimally fracture resistant as a function of time; that is, to improve their fatigue life or survivability under a variety of dynamic loading environments. There is a continuing need to improve fracture toughness for a scaffold; and in particular a peripherally implanted scaffold. 
     The fracture resistance of a vascular scaffold depends not only on the design and the material, but is also the manufacturing process and deployment parameters. Therefore it is in particular necessary to have a process, design, and a delivery system that allows the scaffold to be uniformly expanded and deployed. As a consequence of non-uniform deployment the various struts and crowns of a scaffold will potentially be exposed to very different forces and motions, which has a deleterious effect on the fatigue life. 
     An useful dimensionless number useful for characterizing a material&#39;s fracture toughness is called a Deborah number (Ratio of intrinsic material damping time constant and time constant of external applied force). The higher the Deborah number, the greater is the expected potential of an implant to fracture under a transient load or fatigue load of a given amplitude. 
     Toughening domains can be introduced into an implant design in several ways: a) backbone alteration to include low Tg blocks, e.g. block copolymers, b) polymer blends and c) introducing light crosslinks into the backbone. 
     Fracture toughness of a homopolymer such as PLLA can also be improved by controlling the microstructure of the final implant. Variables such as % crystallinity, size and/or distribution of crystallites, spatial distribution, and gradient and shape of the crystalline domains. A combination of these micro-structural controls in combination with a macroscopic design, e.g., scaffold pattern, crimping process, etc. may improve fracture toughness without significant adverse affects on other scaffold material properties, e.g., radial and/or pinching stiffness. 
     An alternative to providing elastomeric properties is the use of a multilayered structure having “soft” and “hard” layers, where the soft layer/layers would be made from a low Tg material and the hard layers would have a high Tg material. In a similar way high and low Tg domains can generate typical rubber-toughened morphologies through the use of block copolymers or polymer blends. The Tg of a given domain/block could be generated from a given monomer or by the use of several monomers in a random co-polymer. Typical low Tg materials can be made from caprolactone, lactone derivatives, carbonate, butylsuccinate, trimethylene carbonate, dioxanone or other known monomers in accordance with the disclosure. Other low Tg materials that could be used, would be a material that clears the kidneys through dissolution rather than degradation. Such material may include polyethylene glycol (PEG), polyvinylpyrrolidone (PVP), or polyvinylalochol (PVA), or other known polymers in accordance with the disclosure. 
     Alternative ways to improve the fatigue properties are through introduction of axial flexibility and the use of pre-designed fracture points, in particular in the connector links. The fracture points could function as precursors of actual fractures, e.g., crazes and cracks or small dimension of fracture distributed in the implant. A distribution or pattern of cracks or crazes may dictate or inform one of an expected toughness of the scaffold when subjected to a particular loading, e.g., torsion, radial force, tensile etc. Although it is understood that, due to the generally highly non-linear relationship between crack formation and a coupled loading environment, that is, simultaneously applied and time varying bending, torsion and axial loading, such predictive methods may not be applicable to all situations. 
     Alternative ways to improve the fatigue properties are through introduction of axial flexibility and the use of pre-designed fracture points, in particular, fracture points in or near connector links as discussed in greater detail below. 
     For a tube of  FIG. 1  having a diameter about 7 mm and a wall thickness above 200 micro-meters and more specifically a diameter of 8 mm and a wall thickness of 280 micro-meters, the temperature at expansion is 235+/−5 degrees Fahrenheit, the expansion pressure is 110+/−10 psi and the expansion speed is 0.68+/−0.20 mm/sec. 
     The degree of radial expansion that the polymer tube undergoes can partially characterize the degree of induced circumferential molecular and crystal orientation as well as strength 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).times.100%. The degree of axial extension that the polymer tube undergoes can partially characterize induced axial molecular or crystal orientation as well as strength 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).times.100%. In a preferred embodiment the RE is about 400% and the AE is 40-50%. 
     The strengthened and toughened cylindrical, polymer tube of  FIG. 1  is formed into a scaffold structure, in one embodiment a structure having a plurality of struts  230  and links  234  forming a pattern  200  as shown in  FIG. 2  (pattern  200  is illustrated in a planar or flattened view), which is about the pattern for the scaffold before crimping and after the scaffold is plastically, or irreversibly deformed from its crimped state to its deployed state within a vessel by balloon expansion. The pattern  200  of  FIG. 2 , therefore, represents a tubular scaffold structure (as partially shown in three dimensional space in  FIG. 3 ), so that an axis A-A is parallel to the central or longitudinal axis of the scaffold.  FIG. 3  shows the scaffold in a state prior to crimping or after deployment. As can be seen from  FIG. 3 , the scaffold comprises an open framework of struts and links that define a generally tubular body. The cylindrical, deformed tube of  FIG. 1  may be formed into this open framework of struts and links described in  FIGS. 2-3  by a laser cutting device, preferably, a pico-second green light laser that uses Helium gas as a coolant during cutting. 
     Details of a suitable laser process can be found in U.S. application Ser. No. 12/797,950. The Helium gas is necessary to avoid melting or altering properties of the scaffold structure adjacent the laser&#39;s cutting path. Exemplary laser machining parameters are provided in Table 1. 
     
       
         
           
               
             
               
                 TABLE 1 
               
             
            
               
                   
               
               
                 Laser Machining Parameters for a crush recoverable polymer scaffold 
               
               
                 having a wall thickness of between about .008″ and .014″ 
               
            
           
           
               
               
               
            
               
                   
                 Parameter 
                 Range 
               
               
                   
                   
               
               
                   
                 Scaffold length (mm) 
                  8-200 
               
               
                   
                 No. of passes to cut 
                 2-4 
               
               
                   
                 Cutting speed (in/min) 
                  4-10 
               
               
                   
                 Fast jog speed (in/min) 
                 10-14 
               
               
                   
                 Max accel/decal (in/min 2 ) 
                 0-6 
               
               
                   
                 Tube outer diameter 
                  6-12 
               
               
                   
                 Laser spot size 
                 14-20 
               
               
                   
                 Laser rep rate (kHz) 
                 25-50 
               
               
                   
                 Laser power setting (W) 
                  .8-1.22 
               
               
                   
                 Helium gas flow (scfh) 
                 11-17 
               
               
                   
                   
               
            
           
         
       
     
     Referring to  FIG. 2 , the pattern  200  includes longitudinally-spaced rings  212  formed by struts  230 . A ring  212  is connected to an adjacent ring by several links  234 , each of which extends parallel to axis A-A. In this first embodiment of a scaffold pattern (pattern  200 ) four links  234  connect the interior ring  212 , which refers to a ring having a ring to its left and right in  FIG. 2 , to each of the two adjacent rings. Thus, ring  212   b  is connected by four links  234  to ring  212   c  and four links  234  to ring  212   a.  Ring  212   d  is an end ring connected to only the ring to its left in  FIG. 2 . 
     A ring  212  is formed by struts  230  connected at crowns  207 ,  209  and  210 . A link  234  is joined with struts  230  at a crown  209  (W-crown) and at a crown  210  (Y-crown). A crown  207  (free-crown) does not have a link  234  connected to it. Preferably the struts  230  that extend from a crown  207 ,  209  and  230  at a constant angle from the crown center, i.e., the rings  212  are approximately zig-zag in shape, as opposed to sinusoidal for pattern  200 , although in other embodiments a ring having curved struts is contemplated. As such, in this embodiment a ring  212  height, which is the longitudinal distance between adjacent crowns  207  and  209 / 210  may be derived from the lengths of the two struts  230  connecting at the crown and a crown angle θ. In some embodiments the angle θ at different crowns will vary, depending on whether a link  234  is connected to a free or unconnected crown, W-crown or Y-crown. 
     The zig-zag variation of the rings  212  occurs primarily about the circumference of the scaffold (i.e., along direction B-B in  FIG. 2 ). The struts  212  centroidal axes lie primarily at about the same radial distance from the scaffold&#39;s longitudinal axis. Ideally, substantially all relative movement among struts forming rings also occurs axially, but not radially, during crimping and deployment. Although, as explained in greater detail, below, polymer scaffolds often times do not deform in this manner due to misalignments and/or uneven radial loads being applied. 
     The rings  212  are capable of being collapsed to a smaller diameter during crimping and expanded to a larger diameter during deployment in a vessel. According to one aspect of the disclosure, the pre-crimp diameter (e.g., the diameter of the axially and radially expanded tube from which the scaffold is cut) is always greater than a maximum expanded scaffold diameter that the delivery balloon can, or is capable of producing when inflated. According to one embodiment, a pre-crimp diameter is greater than the scaffold expanded diameter, even when the delivery balloon is hyper-inflated, or inflated beyond its maximum use diameter for the balloon-catheter. 
     Pattern  200  includes four links  237  (two at each end, only one end shown in  FIG. 2 ) having structure formed to receive a radiopaque material in each of a pair of transversely-spaced holes formed by the link  237 . These links are constructed in such a manner as to avoid interfering with the folding of struts over the link during crimping, which, as explained in greater detail below, is necessary for a scaffold capable of being crimped to a diameter of about at most Dmin or for a scaffold that when crimped has virtually no space available for a radiopaque marker-holding structure. 
     A second embodiment of a scaffold structure has the pattern  300  illustrated in  FIG. 4 . Like the pattern  200 , the pattern  300  includes longitudinally-spaced rings  312  formed by struts  330 . A ring  312  is connected to an adjacent ring by several links  334 , each of which extends parallel to axis A-A. The description of the structure associated with rings  212 , struts  230 , links  234 , and crowns  207 ,  209 ,  210  in connection with  FIG. 2 , above, also applies to the respective rings  312 , struts  330 , links  334  and crowns  307 ,  309  and  310  of the second embodiment, except that in the second embodiment there are only three struts  334  connecting each adjacent pair of rings, rather than four. Thus, in the second embodiment the ring  312   b  is connected to the ring  312   c  by only three links  234  and to the ring  312   a  by only three links  334 . A link formed to receive a radiopaque marker, similar to link  237 , may be included between  312   c  and ring  312   d.    
       FIGS. 5A and 5B  depict aspects of the repeating pattern of closed cell elements associated with each of the patterns  300  and  200 , respectively.  FIG. 5A  shows the portion of pattern  300  bounded by the phantom box VA and  FIG. 5B  shows the portion of pattern  200  bounded by the phantom box VB. Therein are shown cell  304  and cell  204 , respectively. In  FIGS. 5A, 5B  the vertical axis reference is indicated by the axis B-B and the longitudinal axis A-A. There are four cells  204  formed by each pair of rings  212  in pattern  200 , e.g., four cells  204  are formed by rings  212   b  and  212   c  and the links  234  connecting this ring pair, another four cells  204  are formed by rings  212   a  and  212   b  and the links connecting this ring pair, etc. In contrast, there are three cells  304  formed by a ring pair and their connecting links in pattern  300 . 
     Referring to  FIG. 5A , the space  336  and  336   a  of cell  304  is bounded by the longitudinally spaced rings  312   b  and  312   c  portions shown, and the circumferentially spaced and parallel links  334   a  and  334   c  connecting rings  312   b  and  312   c.  Links  334   b  and  334   d  connect the cell  304  to the right and left adjacent ring in  FIG. 3 , respectively. Link  334   b  connects to cell  304  at a W-crown  309 . Link  334   d  connects to cell  304  at a Y-crown  310 . A “Y-crown” refers to a crown where the angle extending between a strut  330  and the link  334  at the crown  310  is an obtuse angle (greater than 90 degrees). A “W-crown” refers to a crown where the angle extending between a strut  330  and the link  334  at the crown  309  is an acute angle (less than 90 degrees). The same definitions for Y-crown and W-crown also apply to the cell  204 . There are eight connected or free crowns  307  for cell  304 , which may be understood as eight crowns devoid of a link  334  connected at the crown. There are one or three free crowns between a Y-crown and W-crown for the cell  304 . 
     Additional aspects of the cell  304  of  FIG. 5A  include angles for the respective crowns  307 ,  309  and  310 . Those angles, which are in general not equal to each other (see e.g.,  FIG. 6A  for the “V2” and “V23” embodiments of scaffold having the pattern  300 ), are identified in  FIG. 5A  as angles  366 ,  367  and  368 , respectively associated with crowns  307 ,  309  and  310 . For the scaffold having the pattern  300  the struts  330  have strut widths  361  and strut lengths  364 , the crowns  307 ,  309 ,  310  have crown widths  362 , and the links  334  have link widths  363 . Each of the rings  312  has a ring height  365 . The radii at the crowns are, in general, not equal to each other. The radii of the crowns are identified in  FIG. 5A  as radii  369 ,  370 ,  371 ,  372 ,  373  and  374 . 
     Cell  304  may be thought of as a W-V closed cell element. The “V” portion refers to the shaded area  336   a  that resembles the letter “V” in  FIG. 6A . The remaining un-shaded portion  336 , i.e., the “W” portion, resembles the letter “W”. 
     Referring to  FIG. 5B , the space  236  of cell  204  is bounded by the portions of longitudinally spaced rings  212   b  and  212   c  as shown, and the circumferentially spaced and parallel links  234   a  and  234   c  connecting these rings. Links  234   b  and  234   d  connect the cell  204  to the right and left adjacent rings in  FIG. 2 , respectively. Link  234   b  connects to cell  236  at a W-crown  209 . Link  234   d  connects to cell  236  at a Y-crown  210 . There are four crowns  207  for cell  204 , which may be understood as four crowns devoid of a link  234  connected at the crown. There is only one free crown between each Y-crown and W-crown for the cell  204 . 
     Additional aspects of the cell  204  of  FIG. 5B  include angles for the respective crowns  207 ,  209  and  210 . Those angles, which are in general not equal to each other (see e.g.,  FIG. 6B  for the “V59” embodiment of a scaffold having the pattern  200 ), are identified in  FIG. 5B  as angles  267 ,  269  and  268 , respectively associated with crowns  207 ,  209  and  210 . For the scaffold having the pattern  200  the struts  230  have strut widths  261  and strut lengths  266 , the crowns  207 ,  209 ,  210  have crown widths  270 , and the links  234  have link widths  264 . Each of the rings  212  has a ring height  265 . The radii at the crowns are, in general, not equal to each other. The radii of the crowns are identified in  FIG. 5B  as inner radii  262  and outer radii  263 . 
     Cell  204  may be thought of as a W closed cell element. The space  236  bounded by the cell  204  resembles the letter “W”. 
     Comparing  FIG. 5A  to  FIG. 5B  one can appreciate that the W cell  204  is symmetric about the axes B-B and A-A whereas the W-V cell  304  is asymmetric about both of these axes. The W cell  204  is characterized as having no more than one crown  207  between links  234 . Thus, a Y-crown crown or W-crown is always between each crown  207  for each closed cell of pattern  200 . In this sense, pattern  200  may be understood as having repeating closed cell patterns, each having no more than one crown that is not supported by a link  234 . In contrast, the W-V cell  304  has three unsupported crowns  307  between a W-crown and a Y-crown. As can be appreciated from  FIG. 5A , there are three unsupported crowns  307  to the left of link  334   d  and three unsupported crowns  307  to the right of link  334   b.    
     The mechanical behavior of a scaffold having a pattern  200  verses  300  differs in the following ways. These differences, along with others to be discussed later, have been observed in comparisons between the scaffold of  FIGS. 6A-6B , which include in-vivo testing. In certain regards, these tests demonstrated mechanical aspects of scaffold according to the invention that were both unexpected and contrary to conventional wisdom, such as when the conventional wisdom originated from state of the art metallic stents, or coronary scaffold. For a particular design choice, whether driven by a clinical, production yield, and/or delivery profile requirement, therefore, the following characteristics should be kept in mind. 
     In general, a polymer scaffold that is crush-recoverable, possesses a desired radial stiffness and strength, fracture toughness and is capable of being crimped down to a target delivery diameter, e.g., at least about Dmin, balances the three competing design attributes of radial strength/stiffness verses toughness, in-vivo performance verses compactness for delivery to a vessel site, and crush recovery verses radial strength/stiffness. 
     In-vivo performance verses compactness for delivery to the vessel site refers to the ability to crimp the scaffold down to the delivery diameter. The ring struts  230  connecting crowns to form the W-cell  204  are more restrained from rotating about an axis tangent to the abluminal surface (axis A-A). In the case of the W-V cell the V portion, the crown may tend to twist about the axis A-A under particular configurations due to the reduced number of connecting links  336 . The ring portions can in effect “flip”, which means rotate or deflects out-of-plane as a result of buckling (please note: “out-of-plane” refers to deflections outside of the arcuate, cylindrical-like surface of the scaffold; referring to  FIG. 5A  “out-of-plane” means a strut that deflects normal to the surface of this figure). When there is a link  234  at each of a crown or valley as in  FIG. 5B , any tendency for the crown to buckle or flip is reduced because the ring struts are more restrained by the link  236 . Essentially, the link serves to balance the load across a ring more evenly. 
     The “flipping” phenomenon for a scaffold constructed according to pattern  300  has been observed during crimping, as explained and illustrated in greater detail in U.S. application Ser. No. 12/861,719. The W-V cell  304  is devoid of a nearby link  334  at a crown  307  to restrain excessive twisting of the adjacent crown or valley. In essence, when there are two crowns  307  between a link  334  the restraint preventing flipping or buckling of the V portion of the ring depends on the buckling strength of the individual ring strut  330 , i.e., the strength and stiffness of the polymer strut in torsion. When there is a link  234  connected to each adjacent crown/valley ( FIG. 5B ), however, out of plane deflections at the crown  207  is restrained more, due to the bending stiffness added by the connected link  234 , which restrains twisting at the adjacent crown  207 . 
     A scaffold according to pattern  200  is correspondingly stiffer than a similarly constructed scaffold according to pattern  300 . The scaffold according to pattern  200  will be stiffer both axially and in longitudinal bending, since there are more links  236  used. Increased stiffness may not, however, be desirable. Greater stiffness can produce greater crack formation over a less stiff scaffold. For example, the stiffness added by the additional links can induce more stress on rings interconnected by the additional links  234 , especially when the scaffold is subjected to a combined bending (rings moving relative to each other) and radial compression and/or pinching (crushing). The presence of the link  234  introduces an additional load path into a ring, in addition to making the ring more stiff. 
     In-vivo requirements can favor a scaffold according to pattern  200 , but a scaffold according to pattern  300  may be more easily crimped down to the delivery diameter. Other factors also affect the ability to crimp a scaffold. According to the disclosure, it was found that crown angles less than about 115 degrees for the pre-crimp scaffold can produce less fracture and related deployment problems (e.g., uneven folding/unfolding of ring struts) than scaffold with higher crown angles (relative to the inflated diameter, in one case 6.5 mm). The scaffold is crimped to a balloon that can be inflated up to about 7.4 mm. Thus, when the balloon is hyper-inflated the scaffold attains about up to about a 7 mm inflated diameter. For a balloon catheter-scaffold assembly according to the disclosure the largest inflated diameter for the balloon is less than or equal to the scaffold diameter before crimping. As mentioned above, it is preferred that the maximum inflated diameter for the scaffold is less than the scaffold diameter before crimping. 
     During the course of designing a crush recoverable polymer scaffold having a desired crimped profile, it was found that when forming the scaffold at the 8 mm diameter it was difficult to crimp the scaffold to a desired crimped profile, e.g., to crimp the scaffold from the 8 mm diameter to about 2 mm profile, for two reasons. First, by imposing the 350-400% diameter reduction requirement, the polymer material was more susceptible to crack formation and propagation, simply due to strain levels experienced by the scaffold when subjected to this extensive diameter reduction. This concern was addressed by adjusting stiffness, e.g., reducing the strut angle, wall thickness and/or number of crowns. Additionally, the process steps used to form the tube ( FIG. 1 ) was found to help improve the scaffold&#39;s resistance to crack formation and propagation, as explained earlier. 
     Second, even when the scaffold dimensions were adjusted to limit crack formation, there was the problem of limited space for scaffold within the crimped profile. Due to the mass of material associated with the crimped scaffold, the available space for compression of the rings to the desired crimped profile was not achievable without creating unacceptable yield stresses or fracture. Thus, even when a 350-400% diameter reduction was achievable without crack or deployment problems, the scaffold pattern would not allow further reduction without exceeding the range of articulation that the scaffold design would allow. 
     According to another aspect of the disclosure, there are modified crown designs for a scaffold intended to improve the fracture toughness and/or reduce the delivery diameter of the scaffold. It was discovered that a design change to an existing scaffold pattern that would overcome a limitation on reduced profile, and which could be implemented using a brittle polymer like PLLA of PLGA, was a significant reduction in the size of the inner radius of the crown or valley bridging the struts that form the crown/valley. 
       FIGS. 7A and 7B  illustrate a pair of struts  420 ,  422  near a crown  410 . In the pre-crimp state, the struts  420 ,  422  are separated by the crown angle ϕ and the crown is formed with an inner radius r a . This is a typical design for a crown. The inner radius is selected to avoid stress concentrations at the crown. As the art has taught when there is a dramatic change in geometry at a hinge point, such as a crown, there is a greater likelihood cracks or yielding will form at the hinge point (thereby affecting radial strength) since the moment of inertia in bending across the crown is discontinuous. 
     In the case of a metal stent, the angle ϕ before crimping is less than the angle when the stent is deployed. By forming the stent with the reduced diameter, the stent may be more easily crimped to a small profile. Due to the presence of the inner radius, the angle ϕ is capable of being exceeded at deployment without loss of radial stiffness. If this radius is too small, however, and the strut angle at deployment exceeds ϕ, there is a greater chance of yielding or other problems to develop due to stress concentrations at the inner radius. Due to the ductility and resiliency of metal, stents made from metal may also be crimped down further than shown in  FIG. 7B . The struts  420 ,  422  may touch each other, i.e., S is less than 2×r a , and yet the stent can still recover and maintain its radial stiffness despite the over crimped condition. 
     For polymer scaffold, however, it has been found that the distance S ( FIG. 7B ) should not generally be smaller than allowed for the radius r a , i.e., S greater than or equal to 2 r a . For a polymer scaffold, if the struts  420 ,  422  are brought closer to each other, i.e., S becomes less than 2×r a , the brittleness of the material can likely result in fracture problems when the scaffold is deployed. The scaffold may not therefore be able to maintain its radial stiffness if crimped beyond the allowable distance for the radius. The scanning electron microscope (SEM) photographs included as  FIGS. 8A, 8F and 8G  show fractures at crowns when the distance S in  FIG. 7B  is less than 2×r a . As can be seen in these photographs, there is significant material failure in a W crown, free crown and Y crown. 
     With the objective of decreasing the distance S between struts  420 ,  422  ( FIG. 7B ) the inventors decided to reduce down the radius r a  as small as possible, despite the advice offered by the art. It was discovered, to their surprise, that the scaffold was able to recover from the crimped condition to the expanded condition without significant, noticeable, reoccurring or prohibitive loss in radial strength. The SEMs provided as  FIGS. 8B, 8C and 8D  show crowns/valleys having reduced radii after being crimped, then expanded by the balloon. In these examples the crown inner radii were made as small as the cutting tool (a green light pico-second laser, described above) was able to produce. As can be seen by comparing  FIGS. 8A, 8F and 8G  with  FIGS. 8B, 8C and 8D  the scaffold having reduced radii produced some voids but there is no crack propagation. Structural integrity was maintained. The deployed scaffold in these photos maintained good radial stiffness. 
       FIGS. 7C and 7D  illustrate embodiments of a crown formation that produced these unexpected results. An example of a W cell having a reduced radii type of crown formation just described is illustrated in  FIG. 5B and 6B . The radius r b  is about 0.00025 inches, which corresponds to the smallest radius that could be formed by the laser. The 0.00025 inch radius is not contemplated as a target radius or limit on the radius size, although it has produced the desired result for this embodiment. Rather, it is contemplated that the radius may be as close to zero as possible to achieve a reduced profile size. The radius, therefore, in the embodiments can be about 0.00025 (depending on the cutting tool), greater than this radius, or less than this radius to practice the invention in accordance with the disclosure, as will be appreciated by one of ordinary skill in the art. For instance, it is contemplated that the radii may be selected to reduce down the crimped size as desired. 
     An inner radius at about zero, for purposes of the disclosure, means the minimum radius possible for the tool that forms the crown structure. An inner radius in accordance with some embodiments means the radius that allows the distance S to reduce to about zero, i.e., struts are adjacent and/or touch each other as shown in  FIG. 7D  (S′ is about, or zero). 
     Without wishing to be tied to a particular theory for how the scaffold according to the invention is capable of being reduced down to the theoretical minimum diameter and then expanded without loss of strength, it is believed that the selection of starting diameter being greater than the inflated diameter played a role in the favorable outcome. In contrast to the previous example where a metal stent is formed from a diameter less than its inflated diameter, which smaller diameter may be selected to facilitate a smaller crimped profile, a polymer scaffold according to preferred embodiments is formed from a starting diameter greater than the maximum inflated diameter for the balloon catheter-scaffold assembly (a larger starting diameter may be preferred to reduce acute recoil, as explained below, and/or to enhance radial strength characteristics in the deployed state as explained earlier in the tube processing steps for the tube of  FIG. 1 ). As such, the strut angle pre-crimp is preferably greater than the maximum crown/strut angle when the scaffold is deployed. Stated differently, the crown angle in  FIG. 7C  (pre-crimp angle) is never exceeded when the balloon expands the scaffold from the crimped to deployed state. This characteristic of the crush recoverable polymer scaffold, i.e., pre-crimp crown angle greater than the deployed crown angle, is believed to provide clues as to how the polymer scaffold in the SEM photographs was able to retain radial strength when a minimum inner radius was used for the crown formation, contrary to the prior art. Compression, but not expansion of the scaffold when loaded by the vessel, it is believed, will not induce further weakening, despite the presence of voids. When the crown experiences only a compressive deformation relative to its pre-crimp shape ( FIG. 7C ), the potentially weakened area near the inner radius is subjected to only compressive stresses, which do not tend to tear the crown apart, i.e., induce crack propagation. 
     Crimping of the scaffold, as detailed in U.S. application Ser. No. 12/861,719, includes heating the polymer material to a temperature less than, but near to the glass transition temperature of the polymer. In one embodiment the temperature of the scaffold during crimping is raised to about 5 to 10 degrees below the glass transition temperature for PLLA. When crimped to the final, crimped diameter, the crimping jaws are held at the final crimp diameter for final dwell period. This method for crimping a polymer scaffold having crush recovery is advantageous to reduce recoil when the crimp jaws are released. Another, unexpected outcome, however, was found relating to the reduced inner radius aspect of the disclosure. It was found that during the dwell period the polymer scaffold crimped profile could be reduced to a profile less than the theoretical minimum profile. 
     From the example given earlier for the scaffold of  FIG. 6B , the value for Dmin is 0.1048″ or 2.662 mm. When crimping this scaffold according to the crimping procedure summarized above and described in U.S. application Ser. No. 12/861,719 (docket no. 62571.448), it was found that the scaffold could be reduced down to a crimped profile of 0.079″ or 2.0066 mm. Hence, the crimped profile was less than Dmin for this scaffold. With this profile a protective sheath of 0.085″ OD could be placed over the scaffold. When a drug coating was disposed over the scaffold, the profile of the scaffold with sheath was 0.092″. For this scaffold the range of radial strength was 0.45-0.65 N/mm, range of radial stiffness was 1.00-1.20 N/mm and the crush recoverability was about 90% (50% crush). 
     It is believed that a reduced profile less than Dmin was achieved due to a compression of the material during the dwell period. Essentially, the pressure imposed by the crimping jaws during the dwell period at the raised temperature caused the struts forming the ring to be squeezed together to further reduced the crimped scaffold profile. According to these embodiments, the crimped scaffold having a profile less than its theoretical minimum profile was successfully deployed and tested in vivo. This scaffold possessed the desired radial stiffness properties, in addition to the desired crush recovery of above about 90% following a 50% reduction in diameter. 
     In another aspect of this disclosure, the strut and crown formation for a crush recoverable polymer scaffold is formed to take the shape depicted in  FIG. 7E , for purposes of achieving a crimped profile less than the crimped profile for the scaffold having the crown formation shown in  FIG. 7A . According to these embodiments, the crown is formed with a radius r c  as shown. When this scaffold is crimped, the struts may be brought close together so that the distance separating them is near zero (S″ is about, or zero). In contrast to the embodiments of  FIG. 7C , the radius r c  is made some finite or larger radii than by forming a hole or enlarged area between the ends of the struts and crown. The thickness at the crown, t c ′ forming the inner radius along its inner surface may be less than the strut width (in the example of  FIG. 7C  and  FIG. 16  the crown thickness may be larger than the strut width). This can allow a larger inner radius to be used at the crown without increasing the crimped profile. 
     In these embodiments, a scaffold having the crown formation depicted in  FIGS. 7E-7F  is referred to as a “key-hole” crown formation. The name will be understood without further clarification by reference to  FIG. 7F , which shows a key-hole slot or opening formed by the inner wall surfaces. In the crimped profile, the struts near the crown may be brought closer together while a hole or opening having radius r c  is more or less maintained at the crown. The distance S″ is less than twice the radius r c  for the “key-hole” crown formation. 
     Examples of scaffold embodying patterns  300  and  200  are provided in  FIGS. 6A-6B  (referred to as the V2 embodiment, which has a 0.008 inch wall thickness, V23 embodiments having 0.008 and 0.014 inch wall thickness and the V59 embodiment, which has a 0.011 inch wall thickness). Specific values for the various cell attributes of  FIGS. 5A-5B  are provided. 
     The scaffold V59 (pattern  200 ) having a pre-crimp diameter of 8 mm is capable of being crimped to a non-compliant balloon wherein the crimped profile is about 2 mm. The inflated diameter is about 6.5 mm in this example. The scaffold V2, V23 having pre-crimp diameters 7 and 9, respectively, are expanded to about 6.5 mm by a non-compliant balloon. The V2 and V23 scaffold are capable of being crimped to diameters of about 0.092 inches (2.3 mm). 
     According to the disclosure, it was found that the aspect ratio (AR) of a strut of a scaffold may be between about 0.8 and 1.4, the AR of a link may be between about 0.4 and 0.9, or the AR of both a link and a strut may between about 0.9 and 1.1, or about 1. Aspect ratio (AR) is defined as the ratio of width to thickness. Thus for a strut having a width of 0.0116 and a wall thickness of 0.011 the AR is 1.05. 
     According to the disclosure, the radial strength of a balloon expanded polymer scaffold having crush recoverability has a radial strength of greater than about 0.3 N/mm, or between about 0.32 and 0.68 N/mm, and a radial stiffness of greater than about 0.5 N/mm or between about 0.54 N/mm and 1.2 N/mm. According to the disclosure, a crush-recoverable scaffold has these ranges of stiffness and strength for a scaffold having a wall thickness of about 0.008″ to 0.014″ and configured for being deployed by a 6.5 mm non-compliant balloon from about a 2 mm crimped profile, or deployed to a diameter of between about 6.5 mm and 7 mm from about a 2 mm crossing profile on a balloon catheter. 
     A biodegradable polymer, such as PLLA (and polymers generally composed of carbon, hydrogen, oxygen, and nitrogen) is radiolucent with no radiopacity. It is desirable for a scaffold to be radiopaque, or fluoroscopically visible under x-rays, so that accurate placement within the vessel may be facilitated by real time visualization of the scaffold body, preferably the end rings. A cardiologist or interventional radiologist typically will track a delivery catheter through the patient&#39;s vasculature and precisely place the scaffold at the site of a lesion using fluoroscopy or similar x-ray visualization procedures. For a scaffold to be fluoroscopically visible it must be more absorptive of x-rays than the surrounding tissue. Radiopaque materials in a scaffold may allow for its direct visualization. One way of including these materials with a biodegradable polymer scaffold is by attaching radiopaque markers to structural elements of the scaffold, such as by using techniques discussed in U.S. application Ser. No. 11/325,973. However, in contrast to other stent or scaffold, a biodegradable, bioabsorbable, bioresorbable, or bioerodable, and peripherally implanted scaffold having crush recoverability according to the disclosure has special requirements not adequately addressed in the known art. 
     There is the unmet need of maintaining a desired stiffness property in the vicinity of the marker-holding material (marker structure) without increasing the minimum crimped diameter, e.g., Dmin. The marker-holding material must not interfere with the extremely-limited space available for achieving the required crossing profile or delivery diameter for the crimped scaffold on the delivery catheter, particularly in the case of a scaffold that has a diameter reduction of 300-400% or more when crimped from the starting, pre-crimp diameter to the delivery diameter, and/or where the target delivery diameter is about at most a theoretical minimum diameter (Dmin) for the scaffold. It has been found that in order to be capable of achieving a desired delivery diameter, e.g., 300-400% or more diameter reduction during crimping, the marker material (when located on a link) should not interfere with the folding of the struts forming rings of the scaffold. However, when addressing this need without consideration for the effect on radial stiffness, it was found that there was an unacceptable loss in stiffness in the vicinity of the marker structure. 
     Referring to  FIGS. 9A and 9B  there are shown portions of the scaffold according to pattern  200 .  FIG. 9A  shows the portion of the scaffold where the link  237  holding a radiopaque material  500  (marker  500 ) is located.  FIG. 9B  shows this same portion of the scaffold when configured in a crimped configuration. The rings  212   b ,  212   c,    212   d  and  212   f  are shown in their compressed, folded or compact configuration as crimped rings  212   b ′,  212   c ′,  212   d ′ and  212   f ′, respectively. So that each of the rings  212  may have the same radial stiffness properties (ignoring link connections), the pair of markers  500  is preferably located on the link  237 , as opposed to on a ring strut  230 . In other embodiments the marker  500  may be located on the ring  212  by making suitable accommodation in the ring structure. 
     As can be appreciated from  FIG. 9B , in order to maintain the minimum diameter, e.g., about at least the theoretical minimum crimped diameter (Dmin) for the crimped scaffold, the presence of marker structure preferably has no effect on the distance between folded struts  230 . To achieve this result, the length of the link  237  may be increased, (L 237 =L 1 +L 2 ,) over the length L 1  of the other links  234  that do not have the markers to carry (the length L 2  being about the length needed to accommodate marker structure (depots  502  and the pair of markers  500 ), without interfering or limiting the folding of struts  230  as necessary to achieve a 300-400% or more diameter reduction. Stents or scaffold that do not have a tight crimped diameter requirement or minimum space between structural elements of a scaffold, by contrast, may have the link connecting rings increased in size beneath the fold struts to hold a marker  500 , since there remains available space for marker structure in the crimped configuration. 
     The depots  502  may be formed when the scaffold is cut from the tube. The depots  502  provide a hole sized slightly smaller than a diameter of a marker  500  sphere, e.g., a platinum sphere, so that the sphere may be placed in the hole and secured therein as a drug-polymer coating is applied over the scaffold. The drug-polymer coating can serve as an adhesive or barrier retaining the marker  500  within the hole of a depot  502 . 
     In one aspect of the disclosure the diameter of a sphere forming the marker  500  necessary to achieve adequate illumination is less than the wall thickness ( 235 ,  FIG. 3 ) of the polymer scaffold. As such, the sphere may be placed within the hole and then a coating applied over it. Since the sphere diameter is about equal to or less than the wall thickness  235  no reforming, or shaping of the sphere is necessary to achieve a flat profile. A process of applying the marker, therefore, is simplified. 
     When the length of a link having marker structure is increased to maintain the minimum crimped diameter according to the embodiments of  FIG. 9 , however, the combined radial stiffness properties of the nearby rings is reduced since they are spaced further apart. To minimize this loss in stiffness, particularly with respect to the end ring (which is inherently less stiff since it is connected to only one neighboring ring), the marker structure is located between links  212   c  and  212   f,  as opposed to rings  212   d  and  212   f.  Additionally, the marker structure is arranged so that the marker pair  500  is placed in depots  502   a,    502   b  orientated along the vertical axis B-B as opposed to longitudinally (axis A-A). By placing the depots  502   a  and  502   b  along axis B-B the length L 2  is preferably less than if the markers  500  were disposed longitudinally, so that the undesirable loss in the combined radial stiffness of the adjacent rings  212   c,    212   f  (resulting from the increased length of link  237 ) and the end ring  212   d  is minimal. 
     According to another embodiment of a marker for the polymer scaffold, a scaffold according to the pattern  200  may be devoid of link  237  having the marker structure and increased length needed to accommodate crimping requirements. Referring to  FIGS. 10A and 10B , instead, a radiopaque sheet of material  504 , e.g., a 0.025″ length and 0.004″ thick gold, platinum or Iridium foil, is wrapped around a link  234  and held in place by, e.g., a drug-polymer coating deposited over the scaffold. Since the thickness of the foil may be negligible, or the material compressible during crimping, the scaffold may be capable of maintaining at least about a Dmin crimped diameter despite the presence of the marker  504  between folded struts  230 . According to these embodiments, since the foil does not affect scaffold function—the link length may be about the same as other links  234 —the foil may be preferably placed nearer to the end of the scaffold to facilitate easier identification of the scaffold end within the vessel. For example, the marker  504  may be located on the link connecting ring  212   d  to ring  212   f  since stiffness properties are not affected by the presence of the marker  504 . 
     According to other embodiments of a marker for the polymer scaffold, as depicted in  FIGS. 11A and 11E , a scaffold according to the pattern  200  is modified in the ring structure to hold a radiopaque marker. By placing marker(s) material on or near a crown  207  of the end ring  212   d,  as shown in  FIGS. 11A-11E , the location of the end ring  212   d  in the vessel can be more easily located (since the marker is located on the end ring). According to the embodiments depicted in  FIGS. 11A, 11B and 11E  one or more cylindrically shaped markers  511 ,  516 ,  531 , respectively, may be located at the crown  207  ( FIG. 11A and 11B ) or near the crown  207  as in  FIG. 11E . According to the embodiments depicted in  FIGS. 11C and 11D  one or more strips of radiopaque material  521 ,  526  are placed near the crown ( FIG. 11C ) or around the crown ( FIG. 11D ). 
     A single marker  511  may be received in a hole  512  formed at the crown  207 , in the case of  FIG. 11A , or received in a hole provided by an eyelet  519  that extends from the crown  207 , as shown in  FIG. 11B . In the later case, it may be necessary to increase the ductility, or fracture toughness of the material forming the extension  519  to avoid the eyelet breaking off from the crown  207 . Since there is no strength/stiffness requirements for this eyelet, it may be practical to alter the material locally so that it is more fracture resistant without affecting the stiffness of the crown. For example, Toughness could be achieved by local heat treatment, local plasticization, or a local coating application. Local heat treatment could be particularly useful if a polymer blend or a block copolymer is used in the backbone of the scaffold.  FIG. 11E  shows three radiopaque pieces  531  received in three holes formed in the strut  532 , which has been made thicker to accommodate for the loss in strength of the strut  230  due to the presence of the holes  534 . 
       FIGS. 11C and 11D  show examples of a strip of radiopaque material  521 ,  526  received in slots  522 ,  528 , respectively, formed in the ring  212   b.  The strip  521  may be located in the strut  524 , or the strip  526  may be located about the crown  207  to increase visibility of the crown. These design choices should also take into account the affect on the bending stiffness of this crown, which is also true of the embodiment of  FIG. 11A . Preferably the slot  522 ,  528  coincides with the neutral axis of the strut and/crown to minimize the effect on bending stiffness at the crown. 
     In other embodiments the strips  521 ,  526  may be made from a material consisting of radiopaque particles dispersed in a bioresorbable material, e.g., 60% Tungsten particles. This embodiment has the advantage of dispersing the radiopaque material within the vessel after the scaffold has biodegraded. 
     In another embodiment a scaffold may have links connecting the end rings lengthened to accommodate a marker, e.g., as shown in  FIG. 9A-9B , without losing substantial radial strength or stiffness at the end ring (due to the increased length of the link) by having metallic spring elements inserted into the crowns. Thus, according to this embodiment there is a marker element. 
     In another embodiment a metallic or composite metal-polymer spring may serve a dual role of providing greater visibility and strengthening the end ring. Referring to  FIG. 11F  there is shown a modified form of the non-symmetric cell  304 ′ when the end ring  312   b ′ forms one of its sides. At the free crown  307 , Y-crown  309  and W crown  310  there is an arched strengthening element, or spring  460  embedded in the crown. The material for member  460  may be, or may include, e.g., Iron, Magnesium, Tungsten to provide, in addition to added strength/stiffness at the crown, greater visibility of the end of the scaffold when implanted within the body as these materials are radiopaque. The positioning of the member  460  relative to a strut&#39;s neutral axis may be closest to its edge such as nearest the outer end of the crown, e.g., furthest from the inner radius of the crown so that the tensile ultimate stress across the strut when the ring is under compression is increased mostly due to the presence of the member  460 . The member  460  is preferably located at each of the crowns at the end ring to serve a dual role of providing greater visibility and adding additional radial strength and stiffness to the end ring (which would otherwise have less radial stiff than interior ring structure since the end ring is connected to only one neighboring ring). 
     Design Process 
     As mentioned earlier, the problem may be stated in general terms as achieving the right balance among three competing design drivers: radial strength/stiffness verses toughness, in-vivo performance verses compactness for delivery to a vessel site, and crush recovery verses radial strength/stiffness. 
     Embodiments having patterns  200  or  300  were found to produce desired results with particular combinations of parameters disclosed herein, or readily reproducible in light of the disclosure. It will be recognized there were no known predecessor balloon-expandable stents having adequate crush recovery to use as a guide (indeed, the art had discouraged such a path of development for a peripheral stent). As such, various polymer scaffold combinations were fabricated based and the following properties evaluated to understand the relationships best suited to achieve the following objectives: 
     Crush recoverability of the scaffold without sacrificing a desired minimal radial stiffness and strength, recoil, deploy-ability and crimping profile; 
     Acute recoil at deployment—the amount of diameter reduction within ½ hour of deployment by the balloon; 
     Delivery/deployed profile—i.e., the amount the scaffold could be reduced in size during crimping while maintaining structural integrity; 
     In vitro radial yield strength and radial stiffness; 
     Crack formation/propagation/fracture when crimped and expanded by the balloon, or when implanted within a vessel and subjected to a combination of bending, axial crush and radial compressive loads; 
     Uniformity of deployment of scaffold rings when expanded by the balloon; and 
     Pinching/crushing stiffness. 
     These topics have been discussed earlier. The following provides additional examples and conclusions on the behavior of a scaffold according to the disclosure, so as to gain additional insight into aspects of the disclosed embodiments. 
     A scaffold fabricated with a pattern similar to pattern  300  ( FIG. 4 ) possessed a good amount of crush recoverability, however, this scaffold&#39;s other properties were not ideal due to memory in the material following balloon expansion. The scaffold, which was initially formed from a 6.5 mm tube and deployed to about the same diameter, had acute recoil problems—after deployment to 6.5 mm it recoiled to about a 5.8 mm diameter. The scaffold also exhibited problems during deployment, such as irregular expansion of scaffold rings. 
     One attempt at solving the design problem proceeded in the following manner. The scaffold&#39;s properties were altered to address stiffness, strength, structural integrity, deployment and recoil problems while maintaining the desired crush recoverability. Ultimately, a scaffold was designed (in accordance with the disclosure) having the desired set of scaffold properties while maintaining good crush recovery properties after a 50% pinch deformation, which refers to the scaffold&#39;s ability to recover its outer diameter sufficiently, e.g., to about 90-95%, following a crushing load that depresses the scaffold to a height about equal to 50% of its un-deformed height. 
     The pinching stiffness (as opposed to the radial stiffness) is most influenced or most sensitive to changes in the wall thickness of the scaffold. As the wall thickness increases, the pinching stiffness increases. Moreover, the crush recoverability of a scaffold is most affected by the stresses created at the regions that deflect most outward in response to the applied load. As explained below, as the wall thickness is increased, the crush recoverability decreases due to an increased concentration of strain energy at the outwardly deflected ends of the scaffold. A design for a crush recoverable scaffold, therefore, must balance the wall thickness for increased pinching stiffness against the reduction in crush recoverability resulting from an increased pinching stiffness. Similarly, although radial stiffness is less affected by changes in wall thickness (since loads are more predominantly in-plane loading as opposed to out of plane during pinching) when wall thickness is altered to affect crush recoverability the radial stiffness must be taken into consideration. Radial stiffness changes when the wall thickness changes. 
     The diagrams drawn in  FIGS. 12A, 12B and 12C  are offered to assist with explaining a relationship between wall thicknesses and crush recoverability.  FIG. 12A  shows a cross-section of a scaffold in its un-deformed (unloaded) state and deformed state when subjected to a pinching load (drawn in phantom). The ends of the scaffold designated by “S” and “S” refer to regions with the highest strain energy, as one can appreciate by the high degree of curvature in these areas when the scaffold is under the pinching load. If the scaffold will not recover or have reduction in recovery from the pinching load (F), it will be because in these regions the material has yielded, which precludes or reduces recovery back to the pre-crush diameter. The equal and opposite crushing forces in force F in  FIG. 12A  deflect the scaffold height from its un-deformed height , i.e., the scaffold diameter, to a deformed height as indicated by δ. The region of the scaffold that will contain the highest degree of strain energy when the crushing force F is being applied is near the axis of symmetry for the deformed shape, which is shown in phantom. In the following discussion, the load reaction or material stress/strain state at the scaffold regions S and S′ will be expressed in terms of the strain energy. 
       FIGS. 12B and 12C  are simplified models of the loaded structure intended to illustrate the effects on the strain energy in region S when the scaffold has different wall thickness. Essentially, the model attempts to exploit the symmetry of the deformed shape in  FIG. 12A  to construct a linear stress-strain representation at region S in terms of a spring have a spring constant K. Accordingly, the scaffold properties are modeled as arcs  10 / 20  (½ of a hoop or ring) or half-cylinder shells supported at the ends. The arc cannot displace downward (Y-direction) when the enforced displacement δ is applied, which is believed acceptable as a boundary condition due to the symmetry in  FIG. 12A . Movement in the x-direction is restrained by the spring having spring constant K. The hemispherical arc  10  in  FIG. 12C  has a thickness ti and the hemispherical arc  20  in  FIG. 12B  has a thickness of t 2 &gt;&gt;t 1 . 
     As the pinching load is applied in  FIGS. 12B and 12C , the arcs  10  and  20  are deformed (as shown in phantom). This is modeled by an enforced displacement of the arcs  10 / 20  at their center by about the amount delta (δ) as in  FIG. 12A . The arc  10  deforms less than arc  20 , however, in terms of its curvature when the enforced displacement is applied, because its flexural rigidity is higher than arc  20 . Since the curvature is less changed in arc  10 , more of the % strain energy resulting from the enforced displacement will be carried by the spring at the ends, where the spring force is restraining outward movement at S. For arc  20  more % strain energy is carried in the arc, as the greater changes of curvature are intended to show, as opposed to the spring restraining movement at the ends. 
     Consequently, for a given applied force the % strain energy at the ends will be greater for arc  10 , since the flexural rigidity of the arc  10  is greater than the arc  20 . This is depicted by the displacement of the spring (x 2 &gt;x 1 ). The % strain energy in the spring restraining arc  20  (i.e., ½ K(x 2 ) 2 /(total strain energy in arc  20 )×100) is greater than the % strain energy in the arc  10  restraining spring (i.e., ½ K(x 1 ) 2 /(total strain energy in arc  10 )×100). From this example, therefore, one can gain a basic appreciation for the relationship between wall thicknesses and crush recoverability. 
     In a preferred embodiment it was found that for a 9 mm scaffold pre-crimp diameter a wall thickness of between 0.008″ and 0.014″, or more narrowly 0.008″ and 0.011″ provided the desired pinching stiffness while retaining 50% crush recoverability. More generally, it was found that a ratio of pre-crimp (or tube) diameter to wall thickness of between about 30 and 60, or between about 20 and 45 provided 50% crush recoverability while exhibiting a satisfactory pinching stiffness and radial stiffness. And in some embodiments it was found that a ratio of inflated diameter to wall thickness of between about 25 and 50, or between about 20 and 35 provided 50% crush recoverability while exhibiting a satisfactory pinching stiffness and radial stiffness. 
     Wall thickness increases for increasing pinching stiffness may also be limited to maintain the desired crimped profile. As the wall thickness is increased, the minimum profile of the crimped scaffold can increase. It was found, therefore, that a wall thickness may be limited both by the adverse effects it can have on crush recoverability, as just explained, as well as an undesired increase in crimped profile. 
     Testing 
     Provided below are results from various tests conducted on scaffolds and stents for purposes of measuring different mechanical properties and making comparisons between the properties of the stents and scaffolds. The stents used in the tests were the Cordis® S.M.A.R.T.® CONTROL® Iliac self-expanding stent (8×40 mm) (“Control stent”), the REMEDY Stent (6×40 mm) by Igaki-Tamai (“Igaki-Tamai stent”), and the Omnilink Elite® stent (6×40 mm). 
     The data presented in Tables 2-6 for the scaffolds V2, V23 and V59 are for scaffolds having the properties listed in Tables  6 A and  6 B, respectively. The scaffolds were crimped to a delivery balloon, then expanded to their inflated diameter using a process similar to the process described at paragraphs [0071]-[0091] of U.S. application Ser. No. 12/861,719. 
     The data presented in Tables 2-6 refer to scaffolds and stent properties after they were expanded by their delivery balloons. For each of the tests reported in Tables 2-6, infra, unless stated otherwise the statistic is a mean value. 
     Table 2 presents data showing the percentage of crush recovery for various scaffold compared with other types of stents. The scaffolds and stents were crushed using a pair of opposed flat metal plates moved together to crush or pinch the stents and scaffold by the respective amounts shown in the tables. The test was conducted at 20 degrees Celsius. 
     Table 2 compares the crush-recoverability of the V2, V23 and V59 scaffold to the Igaki-Tamai stent and Omnilink Elite® (6mm outer diameter and 40 mm length) balloon expandable stent. The crush period was brief (about 0 seconds). 
     
       
         
           
               
             
               
                 TABLE 2 
               
             
            
               
                   
               
               
                 Approximate crush recovery using flat plate test 
               
               
                 at 20 Deg. Celsius (as percentage of starting 
               
               
                 diameter, measured 12 hours following crush) 
               
            
           
           
               
               
               
               
               
            
               
                   
                 when 
                 when 
                 when 
                 when 
               
               
                   
                 crushed by 
                 crushed by 
                 crushed by 
                 crushed by 
               
               
                   
                 an amount 
                 an amount 
                 an amount 
                 an amount 
               
               
                   
                 equal to 
                 equal to 
                 equal to 
                 equal to 
               
               
                   
                 18% of 
                 33% of 
                 50% of 
                 65% of 
               
               
                   
                 starting 
                 starting 
                 starting 
                 starting 
               
               
                 Stent/scaffold 
                 diameter 
                 diameter 
                 diameter 
                 diameter 
               
               
                 type 
                 (18% crush) 
                 (33% crush) 
                 (50% crush) 
                 (65% crush) 
               
               
                   
               
               
                 V23 (.008″ 
                 99% 
                 96% 
                 89% 
                 79% 
               
               
                 wall thickness) 
               
               
                 V23 (.014″ 
                 99% 
                 93% 
                 84% 
                 73% 
               
               
                 wall thickness) 
               
               
                 V59 (.011″ 
                 99% 
                 96% 
                 88% 
                 80% 
               
               
                 wall thickness) 
               
               
                 Igaki-Tamai 
                 99% 
                 94% 
                 88% 
                 79% 
               
               
                 Omnilink 
                 93% 
                 80% 
                 65% 
                 49% 
               
               
                 Elite  (R)   
               
               
                   
               
            
           
         
       
     
     As can be seen in the results there is a dramatic difference between the V2, V23 and V59 crush recovery compared with the Omnilink Elite® coronary stent. The best results are achieved by the V23 (0.008″ wall thickness) and V59 scaffold when taking into consideration the radial strength and stiffness properties of these scaffold compared with the Igaki-Tamai stent (see Table 5). 
     Table 3 compares the crush recovery behavior for a V23 scaffold with 0.008″ wall thickness (FIG. 6A) following a 50% crush. The data shows the percent crush recovery of the V23 scaffold following a brief (approximately 0 seconds), 1 minute and 5 minute crush by an amount equal to 50% of the starting diameter. 
     
       
         
           
               
             
               
                 TABLE 3 
               
             
            
               
                   
               
               
                 Approximate crush recovery of V23 (.008″ wall thickness) 
               
               
                 using flat plate test at 20 Deg. Celsius (as percentage 
               
               
                 of starting diameter, measured 24 hours following crush) 
               
            
           
           
               
               
               
            
               
                   
                 when crushed by an 
                 when crushed by an 
               
               
                   
                 amount equal to 25% 
                 amount equal to 50% 
               
               
                 Crush duration 
                 of starting diameter 
                 of starting diameter 
               
               
                   
               
               
                 0 second crush 
                 100%  
                 99% 
               
               
                 1 minute crush 
                 99% 
                 86% 
               
               
                 5 minute crush 
                 92% 
                 83% 
               
               
                   
               
            
           
         
       
     
       FIG. 13  shows the crush recovery properties for the V59 scaffold when crushed by an amount equal to 50% of its starting diameter over a 24 hour period following removal of the flat plates. There are three plots shown corresponding to the recovery of the scaffold following a 0 second, 1 minute and 5 minute crush duration. The scaffold diameter was measured at different time points up to 24 hours after the flat plates were withdrawn. As can be seen in these plots, most of the recovery occurs within about 5 minutes after the flat plates are withdrawn. It is contemplated, therefore, that an about 90% crush recovery is possible for longer periods of crush, e.g., 10 minutes, ½ hour or one hour, for scaffold constructed according to the disclosure. 
     When the pinching or crushing force is applied for only a brief period (as indicated by “0 sec hold time (50%)” in  FIG. 13 ) tests indicate a recovery to about 95-99% of its initial diameter. When the force is held for 1 minute or 5 minute, tests indicate the recoverability is less. In the example of FIG. 13, it was found that the scaffold recovered to about 90% of its initial diameter. The 1 minute and 5 minute time periods being about the same suggests that any effects of the visco-elastic material succumbing to a plastic or irrecoverable strain when in a loaded state has mostly occurred. 
     In accordance with the disclosure, a crush-recoverable polymer scaffold (having adequate strength and stiffness properties, e.g., the stiffness and strength properties of the scaffold in Table 4, infra) has a greater than about 90% crush recoverability when crushed by an amount equal to about 33% of its starting diameter, and a greater than about 80% crush recoverability when crushed by an amount equal to about 50% of its starting diameter following an incidental crushing event (e.g., less than one minute); a crush-recoverable polymer scaffold has a greater than about 90% crush recoverability when crushed by an amount equal to about 25% of its starting diameter, and a greater than about 80% crush recoverability when crushed by an amount equal to about 50% of its starting diameter for longer duration crush periods (e.g., between about 1 minute and five minutes, or longer than about 5 minutes). 
     An acute recoil problem was observed. In one example, a scaffold was formed from a 7 mm deformed tube having a 0.008″ wall thickness. When the scaffold was balloon deployed to 6.5 mm, the scaffold recoiled to about 5.8 mm. To address this problem, the scaffold was formed from larger tubes of 8 mm, 9 mm and 10 mm. It was found that a larger pre-crimp diameter relative to the intended inflated diameter exhibited much less recoil when deployed to 6.5 mm. It is believed that the memory of the material, formed when the deformed tube was made, reduced the acute recoil. 
     A starting tube diameter of 10 mm, for example, for a scaffold having a 7.4 mm inflated diameter should exhibit less recoil than, say, a 8 mm tube, however, this larger diameter size introduced other problems which discouraged the use of a larger tube size. Due to the larger diameter it became difficult, if not infeasible to reduce the diameter during crimping to the desired crimped diameter of about 2 mm. Since there is more material and a greater diameter reduction, there is less space available to reduce the diameter. As such, when the starting diameter exceeds a threshold, it becomes infeasible to maintain the desired crimped profile. It was found that a 9 mm tube size produced acceptable results in that there was less recoil and a crimped profile of about 2 mm could still be obtained. 
     An excessive starting diameter can introduce other problems during deployment. First, when the diameter reduction from starting diameter to crimped diameter is too great, the local stresses in the scaffold hinge elements, crowns or troughs correspondingly increase. Since the polymer material tends to be brittle, the concern is with cracking or fracture of struts if stress levels are excessive. It was found that the diameter 9 mm starting diameter scaffold (in combination with other scaffold dimensions) could be reduced down to 2 mm then expanded to the 7.4 mm inflated diameter without excessive cracking or fracture. 
     Table 4 compares the acute recoil observed in the V2, V23 and V59 scaffold of  FIGS. 6A and 6B . 
     
       
         
           
               
             
               
                 TABLE 4 
               
             
            
               
                   
               
               
                 Acute recoil comparisons 
               
            
           
           
               
               
               
            
               
                   
                 Stent/scaffold type 
                 percent recoil 
               
               
                   
                   
               
            
           
           
               
               
               
            
               
                   
                 V2 (.008″ wall thickness) 
                 11.3% 
               
               
                   
                 V23 (.008″ wall thickness) 
                 3.9% 
               
               
                   
                 V23 (.014″ wall thickness) 
                 4.3% 
               
               
                   
                 V59 (.011″ wall thickness) 
                 4.5% 
               
               
                   
                   
               
            
           
         
       
     
     As discussed earlier, unlike a metal stent, a design for a polymer scaffold must take into consideration its fracture toughness both during crimping and when implanted within a vessel. For a scaffold located within a peripheral artery the types of loading encountered are in general more severe in terms of bending and axial loading than a coronary scaffold, in addition to the pinching or crush forces experienced by the scaffold, due to the scaffold&#39;s proximity to the surface of the skin, and/or its location within or near an appendage of the body. See e.g. Nikanorov, Alexander, M. D. et al.,  Assessment of self - expanding Nitinol stent deformation after chronic implantation into the superficial femoral artery.    
     As is known in the art, a scaffold designed to have increased radial stiffness and strength properties does not, generally speaking, also exhibit the fracture toughness needed for maintaining structural integrity. The need to have a peripherally implanted polymer scaffold with adequate fracture toughness refers both to the need to sustain relatively high degrees of strain in or between struts and links of the scaffold and to sustain repeated, cyclical loading events over a period of time , which refers to fatigue failure. 
     The methods of manufacture, discussed earlier, of the tube from which the scaffold is formed are intended to increase the inherent fracture toughness of the scaffold material. Additional measures may, however, be employed to reduce instances of fracture or crack propagation within the scaffold by reducing the stiffness of the scaffold in the links, or by adding additional hinge points or crown elements to the ring. Alternatively or in addition, pre-designated fracture points can be formed into the scaffold to prevent fracture or cracks from propagating in the more critical areas of the scaffold. Examples are provided. 
     As mentioned above, a peripherally implanted polymer scaffold is subjected, generally speaking, to a combination of radial compressive, pinching or crushing, bending and axial compression loads. Test results indicate that a majority of cracks can occur in the struts forming a ring, as opposed to the links connecting rings for a peripherally implanted polymer scaffold. Indeed, while bench data may suggest that a scaffold is quite capable of surviving cyclical radial, bending and axial loadings when implanted in a peripheral vessel, when the scaffold is in-vivo subjected to combined axial, flexural and radial loading in a peripheral vessel there is nonetheless unacceptable crack formation, fracture or significant weakening in radial strength. 
     With this in mind, alternative embodiments of a scaffold pattern seek to weaken, or make more flexible the scaffold in bending and axial compression without significantly affecting the radial strength or stiffness of the scaffold. By making links connecting rings more flexible, relative movement between a ring and its neighbor, which occurs when a scaffold is placed in bending or axial compression when rings are not axially aligned with each other, e.g., when the scaffold resides in a curved vessel, does not produce as a high a loading between the ring and its neighbor since the link tends to deflect more in response to the relative movement between the rings, rather than transfer the load directly from one ring to another. 
     Referring to an alternative embodiment of pattern  200 , a scaffold is constructed according to the pattern depicted in  FIGS. 14A and 14B . Pattern  400  is similar to pattern  200  except that a link  434 / 440  connecting the rings  212  is modified to create greater flexibility in the scaffold in bending and axial compression (or tension). Referring to  FIG. 14B , the link  434  includes a first portion  435  having a first moment of inertia in bending (MOI 1 ) nearest a Y-crown of a ring and a second portion  438  having a second moment of inertia (MOI 2 ) in bending nearest a W-crown of the neighboring ring, where MOI 1 &lt;MOI 2 . Additionally, a U-shaped portion  436  is formed in the portion  435  to create, in effect, a hinge or articulation point to reduce bending stiffness further. The U-shape portion  436  opens when the ring  212  rotates clockwise in  FIG. 14B . As such, the link is very flexible in clockwise bending since the bending stiffness about the hinge  436   a  is very low. For counterclockwise rotation, the ends of the U-shaped portion abut, which in effect negates the effect of the hinge  436   a.    
     To construct a scaffold that is equally flexible for both clockwise and counterclockwise bending of the scaffold, the U-shaped portions  434  may be removed so that the increased flexibility is provided solely by the reduced MOI portions of the links, such as by replacing the U-shaped portion  436  in  FIG. 14B  with a straight section having a reduced MOI. An alternative is depicted in  FIG. 14A , which shows alternating inverted U-Links  440   b  and U-links  440   a.  When the scaffold is subject to a clockwise bending moment (i.e., ring  212   d  is displaced downwardly in  FIG. 14A  relative to ring  212   e ) the U-shaped portions of the U-links  440   a  act as hinge points. The “U” opens in response to the relative movement between the adjacent rings  212 , whereas the inverted U-links  440   b  function, essentially, as straight sections since the ends of the inverted “U” will contact each other. Similarly, when the scaffold is subject to a counterclockwise bending moment (i.e., ring  212   d  is displaced upwardly in  FIG. 14A  relative to ring  212   e ) the inverted U-shaped portions of the inverted U-links  440   b  act as hinge points. The inverted “U” opens in response to the relative movement between the adjacent rings  212 , whereas the U-links  440   a  function, essentially, as straight sections since the ends of the “U” each other when the scaffold deflects. 
     In another embodiment a reduced MOI may be achieved by increasing the distance between each ring, or preferably every other ring. For example, the distance between ring  212   a  and  212   b  in  FIG. 2  may be increased (while the distance between rings  212   c,    212   b  remains the same). In this example, a link connecting rings  212   a  and  212   b  can have the same MOI as the link connecting rings  212   b  and  212   c  yet the former link will be less stiff in bending since its length is longer than the later link. 
     In another alternative, the pattern  400  includes a link  442  with opposing “U” shaped portions or an “S” portion, as depicted in  FIG. 14C . The S-link  442  has the MOI 1  and MOI 2  as before, except that the portion  435  of the link  442  has two hinge points,  444   a  and  444   b,  instead of the one in  FIG. 14B . With this arrangement, the link  442  provides a hinge point to increase bending flexibility for both clockwise and counterclockwise bending. As such, for a pattern  400  having links  442  the same link  442  may be used everywhere to achieve greater bending flexibility for both clockwise and counterclockwise bending. 
       FIGS. 14D through 14F  illustrate additional embodiments of a link  442  extending between and connecting the Y and W crowns. These examples show links having variable MOIs, either by shaping the link as the pattern is cut from a tube or by modifying the link after the scaffold has been cut from the tube. 
       FIGS. 14D and 14F  show links  450  and  454 , respectively, formed to have a section having a lower MOI than sections located adjacent the connecting crowns. In the case of link  450  the section  451  having the low MOI is offset from, or non-symmetric about the neutral axis “X” in bending for the sections adjacent crowns. In the case of link  454  the section  455  is symmetric about the neutral axis for the sections adjacent the crowns. This symmetry/non-symmetry contrast for sections  451  and  452  may also be described with respect to an axis of symmetry for the crowns. Thus, for an axis of symmetry “X” for a Y-crown ( 210 ) or W crown ( 209 ), which can be readily identified from the figures, section  451  is asymmetric about the X axis, whereas the necked section  455  of the link  454  is symmetric about this axis, which may be considered a crown axis. 
       FIG. 14E  illustrates an example where material between the ends of the links are removed to form two curved voids  433   a,    433   b  in the link  452 . These embodiments may function in a similar manner as the “S” link discussed earlier. According to this embodiment, a pre-designated fracture point (to fail before the rings fail) is between the voids  433   a,    433   b.  The material forming the voids  433   a,    433   b  may be about the same so as to retain symmetry about the axis X, or they may be a different size to cause the axis of symmetry to not be co-linear with this axis, as in the case of  FIG. 14D . The selection of the void size is based on the desired fracture characteristics relative to the ring and whether it is preferred to have a link less stiff for bending in the clockwise or counterclockwise direction, as explained earlier. 
     In another embodiment, greater fatigue and/or fracture toughness may be achieved by modifying the struts of the ring. Referring to  FIG. 15 , there is shown a pattern similar to pattern  300  except that the rings  450  are formed by curved struts  452  connected at crowns  451 . In this example the struts  452  have a shape approximating one sinusoidal period. By replacing the straight struts of  FIG. 4  with sinusoidal struts there is essentially additional hinge points created in the ring. 
     The number of modified link elements, as discussed in connection with  FIGS. 14A-14C  may be between 5-100% of the links used to connect rings of the scaffold. The U-links or S-links as described may be placed between each ring, or may be placed between every-other ring pair. Additionally, the links may be modified by having their MOI reduced without U or S links. Additionally, one or more connecting links can be removed. When there are less connecting links, e.g., 3 verses 4, the scaffold should generally have a reduced bending and axial stiffness (assuming everything else in the scaffold is unchanged). However, as mentioned earlier, the end-to-end or overall effects on performance, reproduceability, quality control and production capacity for such change in a scaffold is, unfortunately, not as easy to predict as in the case of a metal stent. 
     In another aspect of the disclosure, there is a scaffold pattern having rings formed by closed cells. Each of the closed cells of a ring share a link element that connects the longitudinally-spaced and circumferentially extending strut portions of the closed cell. Each of these closed cell rings are interconnected by a connecting link e.g., links,  434 ,  442 ,  450 ,  452  or  454 , having a reduced bending moment of inertia (MOI) to reduce the flexural rigidity of the structure connecting the closed cell rings. Alternatively, the connecting link can include a pre-designated fracture point, such as by forming an abrupt change in geometry near a high strain region. Returning again to  FIG. 14A , the scaffold pattern depicted has links  440   a  connected to each closed cell ring. For each closed cell  204  there is a first and second connecting link, which are co-linear with each other. The first link has a MOI 1  disposed adjacent the crown and the second link has a MOI 2  disposed distal the crown to produce the pattern shown in  FIG. 14A . Alternatively the links connecting the closed cell rings may have the MOI 1  disposed equidistant from the interconnected closed cell rings. 
     According to an additional aspect of the disclosure, there is a scaffold that includes pre-designated fracture points in the links connecting rings. The fracture points are intended to relive the inter-ring loading through crack formation in the links connecting rings. Since the loading on a crown is reduced or eliminated when there is sufficient crack propagation through the link (load cannot transfer across a crack), by including a pre-designated crack location, one may maintain the integrity of the ring structure at the expense of the links, e.g., links  450 ,  452 , in the event in-vivo loading exceeds the design, particularly with respect to fatigue loading. According to this aspect of the disclosure a link has a reduced MOI near a high strain region and includes an abrupt geometry change, e.g., about 90 degrees mid-span. These pre-designated fracture point in the scaffold may extend between closed cell rings, as described above, or between each ring strut. 
     Cracking/fracture problems are also observed as a consequence of irregular crimping and/or deployment of the scaffold. Irregular deployment is problematic, not only from the viewpoint of the scaffold not being able to provide a uniform radial support for a vessel, but also from the viewpoint of crack propagation, fracture and yielding of structure resulting in loss of strength and/or stiffness in vivo. Examples of irregular deployment include crowns being expanded beyond their design angles and in extreme cases, flipping or buckling of crowns during deployment or crimping. These problems were observed during crimping process and during deployment, examples of which are described in greater detail in U.S. application Ser. No. 12/861,719. 
     Pattern  300  may be susceptible to more of these types of problems than pattern  200 . The links of the pattern provide less support for the ring struts forming the V segment of the W-V closed cell  304 , as compared to pattern  200 . It is believed that the w-shaped closed cell  204  was more capable of deploying without irregularities, such as flipping, due to its symmetry. The asymmetric loading inherent in the W-V cell  304  was more susceptible to buckling problems during crimping or deployment. These potential problems, however, should they arise, may be addressed by adopting modifications to the crimping process. 
     For example, a scaffold having a diameter of 7 mm and asymmetric closed cells (pattern  300 ) was crimped then deployed without any flipping of struts observed. A second scaffold of 9 mm diameter was then crimped to a balloon and deployed. This scaffold had the same pattern  300  as the 7 mm scaffold. The strut or crown angle was increased by the ratio of the diameters, i.e., increased by a factor of 9/7, to compensate for the change in radial stiffness resulting from the increased diameter. When the 9 mm scaffold was crimped, however, flipping occurred in the scaffold struts (primarily in the V section of the W-V closed cell). To correct this problem the W closed cell (pattern  200 ) was tested. This modification helped to reduce instances of flipped struts. Surprisingly, the same irregular crimping/deployment problems have not been observed for the comparable metal stent having a W-V closed cell pattern. It was concluded, therefore, that the flipping problem (in particular) is a phenomenon unique to a polymer scaffold. 
     To avoid flipping phenomena, should it occur in a metal stent, one might consider simply adjusting the moment of inertia of a strut to prevent out of plane (outside of the arcuate, abluminal surface) deflection of a strut. However, as noted earlier, the polymer material introduces constraints or limitations that are not present with a metallic material. In the case of minimizing undesired motion of a strut by modifying bending inertia properties of the strut one needs to be mindful that polymer struts must, generally speaking, be thicker and/or wider than the equivalent metal strut. This means there is less space available between adjacent struts and already higher wall thicknesses than the metal counterpart. This problem of space is further compounded for embodiments that form a polymer scaffold from a tube that is the deployed, or larger than deployed size. It is desirable to have the scaffold reduced in diameter during crimping for passage to the same vessel sites as in the case of the metal stent. Thus, the delivery profile for the crimped scaffold should be about the same as the metal stent. 
     A metal stent may be cut from a tube that is between the deployed and crimped diameters. As such, the spacing between struts is greater and the stent is more easily compressed on the balloon because the stent pre-crimp has a diameter closer to the crimped diameter. A polymer scaffold, in contrast, may be cut from a diameter tube equal to or greater than the deployed state. This means there is more volume of material that must be packed into the delivery profile for a polymer scaffold. A polymer scaffold, therefore, has more restraints imposed on it, driven by the crimped profile and starting tube diameter, that limits design options on strut width or thickness. 
     A well known design requirement for a vessel supporting prosthesis, whether a stent or scaffold, is its ability to maintain a desired lumen diameter due to the inward radial forces of the lumen walls including the expected in vivo radial forces imparted by contractions of the blood vessel. Referring to the examples in  FIGS. 6A-6B , the radial stiffness and radial strength of the scaffold is influenced by the width of struts, crown radii and angles, length of ring struts extending between crowns and valleys, the number of crowns and the wall thickness (thickness  235 ,  FIG. 3 ) of the scaffold. The latter parameter (wall thickness) influences the pinching stiffness, as explained earlier. During the design process, therefore, this parameter was altered to affect pinching stiffness and crush recoverability, although it also has an effect on radial stiffness. In order to affect the radial stiffness, one or more of the foregoing parameters (crown angle, crown radius, ring strut length, crown number, and strut width) may be varied to increase or decrease the radial stiffness. 
     To take one example, when it was found that a 7 mm scaffold&#39;s recoil problem could be overcome by increasing the starting tube diameter to 8 mm, 9 mm or perhaps even 10 mm, an initial approximation to the corresponding changes to the scaffold pattern dimensions involved increasing characteristics such as ring strut length, crown angle and link by the ratio of the diameters, e.g., 8/7 when increasing OD from 7 mm to 8 mm. However, this rough approximation was found to be insufficient in retaining other desired properties, such as crush recoverability. Thus, further refinements were needed. 
     The relationships between radial stiffness and above mentioned parameters are well known. However, the relationship of these stiffness-altering parameters to crush recoverability of a balloon expandable stent, much less a balloon expandable scaffold is not well known, if known at all in the existing art. Accordingly, the design process required the constant comparison or evaluation among radial stiffness, pinching stiffness and crush recoverability (assuming the changes did not also introduce yield or fracture problems during crimping and deployment) when the stiffness parameters were altered to determine whether these and related scaffold properties could be improved upon without significant adverse effects to crush recoverability. 
     When varying these parameters to affect stiffness the following observations were made for a 9 crown and 8 crown scaffold. For a 9 crown pattern and 7-9 mm outer diameter an angle exceeding 115 degrees, while producing a high radial stiffness, also exhibited fracture problems when deployed and an unsatisfactory reduction in crush recoverability. Strut or crown angles found to produce acceptable results were between about 105 and 95 degrees. For a 8 crown scaffold a smaller angle than 115 degrees was preferred for the crown. For the 8 crown scaffold the angle is about less than 110 degrees. Generally speaking, the more crowns the more compliant becomes the scaffold radially and the higher the crown angle the less radially complaint becomes the scaffold. 
     Comparisons were made among mean radial strength (N/mm) and radial stiffness (N/mm) values after e-beam sterilization of a V2, V23 and V59 constructed scaffold (having the properties summarized in  FIGS. 6A-6B ) with the Control stent, Igaki-Tamai stent, and Absolute stent (8.5 mm outer diameter, 36 mm length). Table 5 summarizes the findings. 
     
       
         
           
               
             
               
                 TABLE 5 
               
             
            
               
                   
               
               
                 Radial strength and stiffness comparisons 
               
            
           
           
               
               
               
               
            
               
                   
                   
                 Radial strength 
                 Radial stiffness 
               
               
                   
                 Stent/scaffold type 
                 (sterilized) 
                 (sterilized) 
               
               
                   
                   
               
            
           
           
               
               
               
               
            
               
                   
                 Cordis ® 
                 0.82 
                 0.58 
               
               
                   
                 Igaki-Tamai 
                 0.04 
                 0.09 
               
               
                   
                 Absolute 8.5 ProLL 
                 0.51 
                 0.22 
               
               
                   
                 V2 (.008″ wall thickness) 
                 0.32 
                 0.54 
               
               
                   
                 V23 (.014″ wall thickness) 
                 0.49 
                 1.2 
               
               
                   
                 V23 (.008″ wall thickness) 
                 0.4 
                 0.59 
               
               
                   
                 V59 (.011″ wall thickness) 
                 0.6 
                 0.91 
               
               
                   
                   
               
            
           
         
       
     
     The V2, V23, and V59 had far superior strength and stiffness values over the Igaki-Tamai stent. The V23 with 0.014″ had the highest radial stiffness. The V2, V23 and V59 strength and stiffness values were comparable to the self-expanding stent. 
     Comparisons were also made between the pinching stiffness of scaffold according to the disclosure. The values represent average values in units of N/mm based on three samples. The stiffness values were computed from the measured force required to crush the scaffold to ½ or 50% of its starting diameter, e.g., expanded or inflated diameter, using a flat plate test at 20 Deg Celsius. 
     
       
         
           
               
             
               
                 TABLE 6 
               
             
            
               
                   
               
               
                 Pinching Stiffness 
               
            
           
           
               
               
               
            
               
                   
                 average 
                 standard 
               
               
                 Stent/scaffold type 
                 stiffness 
                 deviation 
               
               
                   
               
               
                 V2 (.008″ wall thickness; 36 mm length) 
                 0.151 
                 .005 
               
               
                 V23 (.008″ wall thickness; 38 mm length) 
                 0.202 
                 .004 
               
               
                 V23 (.014″ wall thickness; 38 mm length) 
                 0.394 
                 .052 
               
               
                 V59 (.011″ wall thickness; 36.5 mm length) 
                 0.537 
                 .037 
               
               
                   
               
            
           
         
       
     
     According to one aspect of the disclosure a crush-recoverable scaffold has a ratios of pinching stiffness to radial stiffness of between about 4 to 1, 3 to 1, or more narrowly about 2 to 1; ratios of pinching stiffness to wall thickness of between about 10 to 70, or more narrowly 20 to 50, or still more narrowly between about 25 and 50; and ratios of scaffold inflated diameter to pinching stiffness of between about 15 and 60 or more narrowly between about 20 to 40. 
     According to another aspect of the disclosure a crush-recoverable scaffold has a desirable pinching stiffness to wall thickness ratio of 0.6-1.8 N/mm 2 . 
     According to another aspect of the disclosure a crush-recoverable scaffold has a desirable pinching stiffness to wall thickness*tube diameter ratio of 0.08-0.18 N/mm 3 . 
     Animal Studies 
     Two animal studies (“Study 1” and “Study 2”) were conducted for the scaffolds described in  FIGS. 6A-6B . The scaffolds were implanted into the iliofemoral artery of a healthy porcine model at 28, 90 and 180 days to evaluate the effectiveness of the polymer scaffold. 
     Study 1: compares the V2 with a Cordis® S.M.A.R.T.® CONTROL® Iliac self-expanding stent having an 8 mm outer diameter and a 40 mm length (hereinafter the “control stent”). Among the features of the implanted V2 and control stent investigated in the study was the degree of, and related complications caused by a chronic outward force effect of the implanted prostheses on the healthy artery at 28, 90 and 180 days following implantation. 
     Study 2: compares the V23-008 and V23-014 to determine the effect wall thickness has on scaffold performance, principally loss in lumen area, scaffold area and growth in neointimal thickness. 
     During the course of the studies the implanted prostheses was subject to various degrees of hip extension and flexion by the swine, which is believed to impose about 10-12% bending, and about 13-18% axial compression of the implanted scaffold and control stent during a maximum hip and knee flexion. 
       FIG. 16  is a plot showing the mean minimal lumen diameter (MLD) of the artery for the control stent and scaffold as measured using optical coherence tomography (OCT). The measurements were taken after 28, 90 and 180 days. After 28 days, the scaffold average MLD was about 3.6 mm (15 samples) while the control stent mean MLD was about 4.7 mm (13 samples). After 90 days the scaffold mean MLD was about 4.4 mm (7 samples) and the control stent mean MLD was about 3.6 mm (5 samples). After 180 days the scaffold mean MLD was about 4.4 mm (9 samples) and the control stent mean MLD was about 4.0 mm (7 samples). The variance in mean MLD after 28, 90 and 180 days for the control stent was much larger than the variance in mean MLD for the scaffold. 
       FIG. 17  shows the mean neointimal thickness (as measured by OCT) after 28, 90 and 180 days. At 28 days the mean control stent neointimal thickness was about 0.4 mm (15 samples) while the mean scaffold neointimal thickness was less than 0.2 mm (13 samples). At 90 days the mean neointimal thickness for the control stent had increased to about 0.43 (7 samples) whereas the mean neointimal thickness for the scaffold had decreased to about 0.1 mm (5 samples). At 180 days the control stent&#39;s mean neointimal thickness had increased to 0.55 mm whereas the scaffold mean neointimal thickness had increased to about 0.19 mm. At the 28, 90 and 180 days the variance in neointimal thickness for the control stent was much higher than for the scaffold. It should be noted that the PLLA scaffold included a drug coating to reduce tissue growth, whereas the control stent did not have a similar drug coating on it. 
       FIG. 18  shows the amount of stenosis measured after 28, 90 and 180 days using OCT. The amount of stenosis was about 22% and 18% for the control stent and scaffold, respectively, after 28 days. After 90 days the amount of stenosis for the control stent had increased to about 25% whereas the scaffold stenosis had decreased to about 5%. After 180 days the amount of stenosis for the control stent remained at about 25% whereas the scaffold stenosis had decreased to about 4%. The variance in stenosis for the control stent was far greater than the scaffold after 28, 90 and 180 days. 
       FIG. 19  shows angiography images of the implanted scaffold and control stent, respectively, taken 180 days after implantation. The dark areas indicate the size of the lumen where the prostheses were implanted. As can be appreciated from these images, the lumen in the vicinity of the control stent has narrowed considerably. It is believed the increase in neointimal thickness, reduced MLD and increased stenosis measured in the vicinity of the control stent,  FIGS. 16, 17 and 18 , respectively, are symptoms of the chronic outward force imposed on the artery by the self-expanding control stent. 
       FIGS. 20 and 21  are plots at 28, 90 and 180 days using the QVA measurement technique (commonly used by physicians).  FIG. 20  depicts the mean and variance late loss (loss in lumen diameter after implantation) for the control stent and scaffold.  FIG. 21  shows the mean and variance for the MLD for the control stent and scaffold. 
       FIGS. 22 and 23  are plots at 28 and 90 days of the histomorphometric area stenosis and neointimal thickening by histomorphometry, respectively, for the control stent and V2 scaffold. Both these plots indicate a steady and less favorable increase in the area stenosis and neointimal growth whereas the area stenosis and neointimal growth for the scaffold was about constant and less than the control stent. 
       FIG. 24  compares the minimum lumen and minimum scaffold areas for the V23 having a 0.008″ wall thickness (“V23/008”) with the V23 having a 0.014″ wall thickness (“v23/014”) after 28 and 90 days. Both the minimum lumen area and minimum scaffold areas were higher for the V23 with a 0.014″ wall thickness.  FIG. 25  shows the lumen area loss after 28 and 90 days. The V23 with 0.014″ wall thickness had less lumen area loss than the V23 with 0.008″ wall thickness. Also, there was less variance among the samples for the V23 with 0.014″ wall thickness.  FIG. 26  shows the mean neointimal thickness between the V23 with 0.008″ and 0.014″ wall thickness. There was less tissue growth on the luminal surface of the scaffold when the 0.014″ wall thickness was implanted. 
     The 30, 90 and 180 day animal studies comparing the control stent to the V2 scaffold indicate that the scaffold exhibits noticeably less problems associated with a chronic outward force as compared to the control stent. The 30 and 90 day study comparing a 0.008″ and 0.014″ wall thickness scaffold indicate there is more likely a reduced loss in lumen diameter, scaffold diameter and less neointimal growth when a higher wall thickness is used for the scaffold. 
     While particular embodiments of the present invention have been shown and described, it will be obvious to those skilled in the art that changes and modifications can be made without departing from this invention in its broader aspects. Therefore, the appended claims are to encompass within their scope all such changes and modifications as fall within the true spirit and scope of this invention.