Implantable medical devices

TECHNICAL FIELD

This invention relates to implantable medical devices and methods of delivering the same.

BACKGROUND

The body includes various passageways such as arteries, other blood vessels, and other body lumens. These passageways sometimes become occluded or weakened. For example, the passageways can be occluded by a tumor, restricted by plaque, or weakened by an aneurysm. When this occurs, the passageway can be reopened or reinforced, or even replaced, with a medical endoprosthesis. An endoprosthesis is typically a tubular member that is placed in a lumen in the body. Examples of endoprosthesis include stents and covered stents, sometimes called "stent-grafts".

An endoprosthesis can be delivered inside the body by a catheter that supports the endoprosthesis in a compacted or reduced-size form as the endoprosthesis is transported to a desired site. Upon reaching the site, the endoprosthesis is expanded, for example, so that it can contact the walls of the lumen.

Prostate enlargement, also known as benign prostate hyperplasia or benign prostate hypertrophy, is a common affliction among older men. The condition involves swelling of the prostate. The prostate surrounds the urethra, or urinary tract, and enlargement of the prostate may restrict passage of urine from the bladder towards the urethra. Benign prostate hyperplasia is uncomfortable because it makes urination difficult or impossible. The condition is also dangerous because it can lead to infection of the bladder and kidneys.

Prostate enlargement can be treated with surgery known as resection. Resection can be accomplished by cutting away a large portion of the prostate gland. Prostate enlargement can also be treated with heat treatment, cold treatment, or ablation.

Sometimes a restricted urethra can be treated with a prostatic stent to support the urethra and keep it open despite pressure from the enlarged prostate. A prostatic stent may be implanted permanently or as an interim Multi-stage expandable stent-grafts are disclosed in <patcit id="pcit0001" dnum="WO0033770A"><text>WO-A-00/33770</text></patcit>.

SUMMARY

The invention relates to an implantable medical device as described in claim 1.

An example which is not part of the invention features a medical device. The medical device includes a balloon catheter having an expandable member, e.g., an inflatable balloon, at its distal end and a stent or other endoprosthesis. The stent is an apertured tubular member formed of a polymer and is assembled about the balloon. The stent has an initial diameter for delivery into the body and can be expanded to a larger diameter by inflating the balloon. The polymer does not flow substantially during expansion and substantial stress relaxation or creep does not occur so that the geometry of the stent is maintained.

In another example, a tubular endoprosthesis including a polymer body is provided and delivered into a body lumen. The endoprosthesis is expanded in the body lumen under conditions of expanding pressure and temperature so that the wall thickness of the polymer body is substantially maintained.

In another example, a polymer tube is formed to a first, large diameter. An aperture pattern is cut into the tube wall. The polymer is crosslinked or crystallized. The polymer tube is deformed to a second, small diameter. The polymer tube is expanded in a body lumen to a diameter larger than the second diameter by application of pressure and heat.

In another example, a polymer tube is formed to a first, small diameter. An aperture pattern is provided in the tube wall. The polymer is crystallized or crosslinked. The tube is expanded in a body lumen by application of pressure and heat.

In another aspect, an implantable medical apparatus includes an element operable for movement within the body by mechanical force applied to the element. The element includes a polymer having a melt or glass transition temperature in the range above body temperature to about 50°C or 60 °C and exhibiting a plateau in a plot of storage modulus as a function of temperature at melt or glass transition. In embodiments, the element is a stent. The stent may be generally a tubular body that includes an apertured wall. The stent may be operable for expansion from a first, smaller diameter to a second larger diameter for implantation in a lumen. The thickness of the stent wall varies by about 1% or less between the first and second diameter.

In another aspect, the invention features a medical device including a polymer having a melt or glass transition temperature above body temperature and exhibiting an approximate plateau in a plot of storage modulus as a function of temperature at melt or glass transition. The melt or glass transition temperature may be, for example, above about 37 °C. The medical device may undergo a triggerable event at about the plateau. The triggerable event may be, for example, a change in the flexibility, a change in the porosity, a change in the coefficient of friction or a change in the surface roughness. The medical device may be, for example, a stent that has a portion that has a collapsed position that can be reverted to an expanded position by a trigger subsequent to insertion into the body.

Aspects may include one or more of the following features. The polymer body includes apertures. The polymer body has a ratio of aperture open area to wall area of about 0.5 or more or 0.7 or more. The endoprosthesis is expanded by simultaneously applying an expanding pressure and heat to the endoprosthesis. The polymer body is heated above the melt or glass transition temperature of polymer in the polymer body. The polymer body is elastomeric at the melt or glass transition temperature. The polymer is elastomeric at body temperature. The polymer is crystalline. The polymer is crosslinked. The polymer is radiation crosslinked. The melt or glass transition temperature is about 40 to 50° C. The melt or glass transition temperature has a transition range of about 5° C or less. The polymer exhibits a plateau in the melt or glass transition range in a plot of storage modulus as a function of temperature. The polymer body includes a drug, radiopaque agent or magnetic heating agent. The polymer is a shape memory polymer, e.g. capable of remembering a smaller diameter configuration after expansion. The polymer is, for example, polynorbornene, polycaprolactone, polyenes, nylons, polycyclooctene (PCO), blends of PCO and styrene-butadiene rubber, polyvinyl acetate/polyvinylidinefluoride (PVAc/PVDF), blends of PVAc/PVDF/polymethylmethacrylate (PMMA), polyurethanes, styrene-butadiene copolymers, polyethylene, trans-isoprene, blends of polycaprolactone and n-butylacrylate and blends thereof. An expansion pressure of about 1 atm or more is applied. The endoprosthesis is delivered on a catheter. The endoprosthesis is delivered to a site of occlusion and the site is simultaneously dilated while expanding the endoprosthesis. The endoprosthesis is delivered to a site of lumen curvature and the endoprosthesis is expanded at the site. The endoprosthesis is delivered to a vascular lumen. The endoprosthesis is delivered adjacent (into) the prostate.

Aspects may include one or more of the following. A heat applicator applies heat to the stent during inflation of the balloon to expand the balloon to the expanded diameter. The polymer has a melt or glass transition temperature in the range of about 40 to 50°C and a modulus at the melt or glass transition temperature sufficient to maintain the stent geometry or under application of pressure and/or heat. The polymer exhibits a plateau in the storage modulus in the range of melt or glass transition temperatures. The stent has a wall thickness of about 0.005 to 5 mm. The stent has an initial unexpanded inner diameter in the range of about 1 mm to 5 mm. The stent has an expanded inner diameter of about 1 mm to 20 mm. The stent may be expandable to about 100% or 400% or more of the initial inner diameter. An example of a coronary stent has an initial inner diameter of about 2 mm, and expanded inner diameter of about 4 mm and the wall thickness is about 0.005 mm to 0.1 mm. The stent can be in the form of a tube including aperture areas provided in the tube. The aperture are in the shape of elongate slots, e.g., when the stent is in the small diameter condition. The apertures have a dimension of about 1 mm or less in the small diameter condition. The apertures are in the shape of diamond-like openings, e.g. when the stent is in an expanded condition. The stent can be a wire-form formed of one or more filaments configured to generally define a tube.

Embodiments may include one or more of the following advantages. A balloon expandable stent made of a polymer can be provided that maintains the integrity of the stent geometry on expansion and heating. Maintenance of stent geometry is desirable since geometry affects, for example, the resistance to compression in the body and a predictable geometry is important to avoid irregular surfaces, kinking, or extensions of material into the body lumen which can interfere with the flow of body fluid. The polymers can be elastomers that have melting or glass transitions at temperatures safe for use in the body and exhibit elastomeric properties at both the melted or glass transition stage and the solid or crystalline phase. The stent body exhibits high resistance to inward compressive forces when the polymer is in the solid or crystalline phase. The elastomeric nature of the polymer in the melted or glass state enhances the ability to maintain geometry as the stent is expanded. For example, the polymer exhibits minimal flow during expansion and the thickness of the stent remains substantially constant. Elastomeric properties in the crystalline or solid state enhance the ability to conform to torturous curvature in narrow body lumens. High compression resistance allows the stent to maintain the body lumen open and resist occluding forces such as elastic recoil or the growth of thrombus from the vessel wall.

In another aspect, the invention features a polymeric stent having a portion that has a collapsed position that can be reverted to an expanded position by heating above a first temperature subsequent to insertion of the stent into a cavity or lumen. The stent may be in the form, for example, of a coiled elongated element (for example, a strand, a tape or a flattened tube). The stent may be further heated to a second temperature that is higher than the first temperature and removed as a substantially uncoiled element. When the stent is in the form, for example, of a coiled elongated flattened tube, the flattened tube may include a central opening that includes a medicament that can be released by the inserted stent. In some implementations, the medicament is compounded into the plastic or is a coating on the plastic. In some implementations, the portion is at an end of the stent and the portion is flared or stepped. In other implementations, the portion includes less than 50 % of the length of the stent.

In another aspect, the invention features a polymeric stent in the form of a coiled elongated element, and having a portion that has a collapsed position that can be reverted to an expanded position by heating above a first temperature subsequent to insertion of the stent into a cavity or lumen. When the stent is heated to a second temperature higher than the first temperature, the modulus of the element lowers sufficiently that the stent can be removed from the cavity or lumen as a substantially uncoiled element.

Another example which is not part of the invention features a polymeric stent including metal particles. A portion of the stent has a collapsed position that can be reverted to an expanded position by heating. The heating may be performed using inductive heating to revert the portion in the collapsed position to the expanded position.

Another example which is not part of the invention features a stent having an exterior surface that includes a plurality of protruding elements that extend outwardly from the surface. The protruding elements may be useful in helping the stent retain its position, for example, after insertion into the prostatic urethra.

In some embodiments, the protruding elements are formed of monofilament. The monofilament may include a plurality of constrictions along its length.

In some implementations, the stent is a polymeric stent and the stent has a portion that has a collapsed position that can be reverted to an expanded position by heating above a first temperature subsequent to insertion of the stent into a cavity or lumen.

The details of one or more embodiments of the invention are set forth in the accompanying drawings and the description below. Other features, and advantages of the invention will be apparent from the description and drawings and from the claims.

DESCRIPTION OF DRAWINGS

<ul id="ul0001" list-style="none" compact="compact"><li><figref idref="f0001">Figs. 1A and 1B</figref> are side views of a portion of a stent in a small diameter and expanded condition, respectively.</li><li><figref idref="f0001">Figs. 1C and 1D</figref> are cross-sectional views of a portion of a stent in a small diameter and expanded condition, respectively.</li><li><figref idref="f0002">Figs. 2A-2C</figref> illustrate delivery of a stent into a body lumen.</li><li><figref idref="f0003">Fig. 3</figref> is a plot of storage modulus as a function of temperature.</li><li><figref idref="f0004">Fig. 3A</figref> is a plot of storage modulus as a function of temperature for samples of PCO with varying degrees of crosslinking.</li><li><figref idref="f0005">Fig. 3B</figref> is a WAXS 2θ plot for samples of PCO with varying degrees of crosslinking.</li><li><figref idref="f0006">Figs. 4</figref> illustrates manufacture and use of a stent.</li><li><figref idref="f0007">Fig. 5</figref> illustrates manufacture and use of a stent.</li><li><figref idref="f0008">Fig. 6</figref> is a perspective view of a stent with an end in an expanded position.</li><li><figref idref="f0008">Fig. 7</figref> is a side view of the stent shown in <figref idref="f0008">Fig. 6</figref>.</li><li><figref idref="f0008">Fig. 8</figref> is a side view of the stent shown in <figref idref="f0008">Fig. 6</figref> with the end in a collapsed position.</li><li><figref idref="f0009">Fig. 8A</figref> is a graph of heat flow as a function of temperature for several POSS polyurethanes.</li><li><figref idref="f0010">Fig. 8B</figref> is a graph of storage modulus as a function of temperature for several POSS polyurethanes.</li><li><figref idref="f0011">Fig. 9</figref> is a perspective view of a wrapping fixture.</li><li><figref idref="f0012">Figs. 10</figref> is a cross-sectional view of a restricted prostatic urethra.</li><li><figref idref="f0013">Fig. 11</figref> is a cross-sectional view illustrating delivery of a stent to the prostatic urethra.</li><li><figref idref="f0013">Fig. 12</figref> is a cross-sectional view of a prostatic stent deployed in the urethra.</li><li><figref idref="f0014">Fig. 13</figref> is a side view of an alternative delivery system.</li><li><figref idref="f0015">Fig. 13A-13B</figref> are side views of an alternative delivery system.</li><li><figref idref="f0016">Fig. 14</figref> is a cross-sectional view illustrating removal of a prostatic stent.</li><li><figref idref="f0017">Fig. 15</figref> is a graph of storage modulus (E') VS temperature for PCO.</li><li><figref idref="f0018">Fig. 16</figref> is a graph of storage modulus (E') VS temperature for</li><li>PVAc/PVDF/PMMA blends.</li><li><figref idref="f0019">Fig. 17</figref> is a side view of an alternative stent with two ends portions in expanded positions.</li><li><figref idref="f0019">Fig. 18</figref> is the stent shown in <figref idref="f0019">Fig. 17</figref> with ends in collapsed positions.</li><li><figref idref="f0019">Fig. 19</figref> is a side view of an alternative stent with three portions in expanded positions.</li><li><figref idref="f0020">Fig. 20</figref> is a side view of an alternative stent made with a flattened tube.</li><li><figref idref="f0020">Fig. 21</figref> is a cross-sectional view of the stent shown in <figref idref="f0020">Fig. 20</figref>, taken along 21-21.</li><li><figref idref="f0020">Fig. 22</figref> is a side view of an alternative stent made with a tape.</li><li><figref idref="f0021">Fig. 23</figref> is a perspective view of a stent with a plurality of protruding elements, the end of the stent is in an expanded position.</li><li><figref idref="f0021">Fig. 24</figref> is a side view of the stent shown in <figref idref="f0021">Fig. 23</figref>.</li><li><figref idref="f0021">Fig. 25</figref> is a side view of the stent shown in <figref idref="f0021">Fig. 23</figref> with the end in a collapsed position.</li><li><figref idref="f0022">Fig. 26</figref> is a perspective view of an alternative stent with a plurality of protruding elements, the end of the stent in an expanded position.</li><li><figref idref="f0022">Fig. 27</figref> is a side view of the stent shown in <figref idref="f0022">Fig. 26</figref>.</li><li><figref idref="f0022">Fig. 28</figref> is a side view of the stent shown in <figref idref="f0022">Fig. 26</figref> with an end in a collapsed position.</li><li><figref idref="f0023">Fig. 29</figref> DSC traces of PLA (top) quenched from T = 180 °C or (bottom) annealed at T = 110 °C for 1 hr.</li><li><figref idref="f0024">Fig. 30</figref> DSC traces for PLA/PVAc blends following annealing for 1 hour at T = 110 °C. A heating rate of 10 °C/min was employed. PLA wt-% is indicated with each trace.</li><li><figref idref="f0025">Fig. 31</figref> Glass transition temperatures measured following quenching of the PLA/PVAc blends (solid points). Solid line is best fit to the Fox equation, 1/T_g = w_a/T_g^a + w_b/T_g^b.</li><li><figref idref="f0026">Fig. 32</figref> Tensile storage modulus versus temperature for a range of PLA/PVAc blends whose composition is indicated in the plot.</li></ul>

DETAILED DESCRIPTION

Referring to <figref idref="f0001">Figs. 1A and 1B</figref>, a stent 10 includes a polymer body 12 generally defining a tube. The stent includes open areas 14. Referring particularly to <figref idref="f0001">Fig. 1A</figref>, in a small diameter condition, such as for delivery to a treatment site in the body, the open areas are relatively small and defined as slots cut through the wall of the stent. Referring particularly to <figref idref="f0001">Fig. 1B</figref>, in the expanded condition, the slots are widened to diamond-like shapes. The expansion mechanism of the stent utilizes a deformation (arrows) of the wall material about the open areas. As illustrated, the expansion results in a generally regular, symmetric, geometric pattern that resists and distributes inward compression of the stent by forces imposed by the lumen wall. Referring as well to <figref idref="f0001">Figs. 1B and 1D</figref>, the wall thickness T of the stent does not substantially change upon expansion from the small diameter collapsed (T_C) condition to the expanded condition (T_E)(T_C=T_E). The polymer does not substantially flow or thin out on expansion, so that a reliable expansion geometry and wall thickness can be achieved. Other stent constructions are suitable. For example, filament-form stents in which filaments of polymer material are arranged to define a generally tubular structure can be used. Open areas are defined between the filaments. An example of a stent design including helical filaments is provided in <patcit id="pcit0002" dnum="US4655771A"><text>Wallsten, U.S. 4,655,771</text></patcit>. Suitable aperture wall designs are also described in <patcit id="pcit0003" dnum="US4733665A"><text>Palmaz U.S. 4,733,665</text></patcit>. Another suitable arrangement is exemplified by the Express stent, commercially available from Boston Scientific, Natick, MA.

Referring to <figref idref="f0002">Figs. 2A-2C</figref>, the delivery of a stent into the body is illustrated. The stent is delivered utilizing a catheter 24, which includes a catheter body 25 that carries a balloon 26 at its distal end. At the proximal end, the catheter includes an inflation apparatus 28 such as a syringe or pump apparatus which can be used to inject and circulate inflation fluid into the catheter, where it is directed by a lumen to the interior of the balloon so that the balloon can be inflated. In addition, the inflation apparatus can include a heating apparatus 30, to heat the inflation fluid directed to the balloon. The catheter is delivered into a vessel 20 to the site of an obstruction 31 typically utilizing a guidewire 32. The guidewire 32 extends through a lumen within the body 25 of the catheter.

Referring particularly to <figref idref="f0002">Figs. 2A and 2B</figref>, the stent 10 is positioned over the inflatable balloon 26. Referring particularly to <figref idref="f0002">Figs. 2A</figref> for delivery into the body, the balloon is initially in a small diameter deflated condition. The stent is in a small diameter condition over the balloon. Referring particularly to <figref idref="f0002">Fig. 2B</figref>, when the treatment site is reached, the balloon is inflated by actuating the inflation apparatus 28. The inflation fluid is heated to heat the polymer body of the stent 10. By providing outward radial force while heating the stent, the stent is expanded into contact with the body lumen. The stent can be expanded simultaneously with the widening of the obstructed region. After expansion to the desired diameter, the temperature of the inflation fluid is typically decreased to reverse the softening of the stent body 10. Referring particularly to <figref idref="f0002">Fig. 2C</figref>, after the temperature of the stent has been reduced in this manner, it remains implanted in the vessel to resist vessel recoil and reduce restenosis after the balloon is deflated and the catheter is removed from the body.

Suitable polymers include those that maintain stent geometry under expansion conditions, allowing for intricate stent geometries such as apertured tubes having high open area to wall ratios. At temperatures above body temperature and under conditions of radial expanding pressure, the stent can be expanded without fracture or substantial irreversible stress relaxation or creep. Typically, the stent is heated to or above the melt or glass transition temperature during expansion. In this condition, the polymer is in a softened state. In this state, the polymer can be predictably deformed, typically about aperture regions during expansion. In addition, the soft condition permits proper apposition of the stent to the lumen wall without kinking and without damage due to excessive stiffness, which could straighten the lumen from its native curvature and lead to dissections or other trauma. After the stent is fully expanded and cooled, the polymer substantially sets in the proper apposition, e.g. about a native curvature. Excessive recoil of the stent to a linear configuration is avoided, reducing trauma about the vessel. At the same time, the polymer can have some elastomeric properties in the cooled, hardened state so that the stent can flex with natural vessel motion. After cooling, the stent exhibits sufficient resistance to inward radial force to reduce restenosis due to, e.g., lumen wall recoil. The polymer has sufficient strength so that the stent wall can be kept relatively thin while resisting restenosis from lumen wall forces.

Suitable polymers include elastomers that are crosslinked and/or crystalline. The crosslink and/or crystalline nature is sufficient to resist excessive creep or stress relaxation when the polymer is heated and expanded. The polymer can be crosslinked so that it exhibits the desired elastomeric properties but not crosslinked to the degree that it becomes excessively brittle. Too little crosslinking does not establish sufficient resistance to flow during heating and expansion to maintain stent geometry. In addition, crosslinking can be adjusted to adjust the melt or glass transition temperature and transition temperature range. A narrow melt transition range is desirable, e.g. 5°C or 10°C or less. Crosslinking can be achieved by application of radiation such as e-beam, UV, gamma, x-ray radiation or by heat-activated chemical crosslinking techniques. Chemical crosslinking agents include peroxides, such as benzoyl peroxide or dicumyl peroxide (DCP), and azo compounds, such as 2,2'-azobis(2,4-dimethyl valeronitrile) or 2,2'-azobis[N-(2-propenyl)-2-methylpropionamide]. Radiation techniques provide the advantage that the polymer typically does not have to be substantially heated to achieve crosslinking. An intricate aperture pattern provided in a stent precursor tube can be maintained and heat-induced flow of pre-crosslinked polymer can be avoided. For gamma radiation, an exposure of about 50-300, e.g. 250 kilograys typically provides sufficient crosslinking. Melting and crystallization temperatures are measured using a differential scanning calorimetry.

The polymer can have elastomeric properties in the melted or softened state. Elastomeric properties at melt or glass transition can be investigated by measuring the modulus of elasticity or storage modulus as a function of temperature and determining the elastomeric nature of the material in the desired expansion temperature range. Referring to <figref idref="f0003">Fig. 3</figref>, a plot of storage modulus as a function of temperature is provided. Storage modulus decreases as the material is heated. At the melt or glass transition, a plateau "P" is typically consistent with an elastomeric nature. At much higher temperatures, the modulus drops off more quickly, indicating a material which could flow under pressure. To determine storage modulus, a dynamic mechanical analyzer (Perkin Elmer) can be used. Dynamic mechanical analysis was carried out in tensile mode at an operating frequency of 1 Hz, a static force of 10mN, and oscillation amplitude of 5 µm (approximately 0.1 % strain) and an automatic tension setting of 125 %. Temperature ramps were conducted at 4 °C/minute over the range -100 °C to 100 °C.

Chemically crosslinking PCO also has a direct impact on the thermomechanical properties, e.g. modulus versus temperature, through the establishment of a permanent network and indirectly through morphological transitions. Revealing such effects by the use of DMA, <figref idref="f0004">Fig. 3A</figref> shows plots of the tensile storage modulus (E') versus temperature for cured PCOs prepared with varying amounts of DCP. All of the PCO samples are characterized by a solid-like storage modulus (about 1.7 GPa) for temperatures below T = -70°C with this modulus value being invariant to the crosslinking density. For temperatures above T = -70°C, the apparent onset of T_g in the PCO samples, E' begins to decrease gradually to a level that is dependent on crosslink density, but spanning 0.05 to 0.5 GPa. The decrease in the modulus with crosslinking in this temperature region can be understood from the results of the DSC and wide angle x-ray scattering (WAXS), <figref idref="f0005">Fig. 3B</figref>, that showed crosslinking reduces the degree of crystallinity of PCO. It is to be expected that the crystalline phase will function as both the fixing mechanism for shape memory and a means of controlling room temperature modulus over a full order of magnitude. For temperatures nearing T = 62°C, close to the melting temperature measured by DSC, the storage modulus of neat PCO begins to decrease sharply to about 2 MPa at the completion of melting at 71°C. As found with DSC, this transition temperature is observed mechanically to decrease with increasing degree of crosslinking. For temperatures greater than T_m, the modulus of neat PCO, trace (i), continues to decrease to a point where the material flows like a viscous liquid, not showing a persistent rubbery plateau (<figref idref="f0003">Figure 3</figref>). This feature hampers the applicability of neat PCO for use as a shape memory polymer due to an inability to be deformed as a rubber above T_m without rapid stress relaxation. On the other hand, cured PCO, which contains just 1% peroxide, represented by trace (ii), will allow significant shape memory effects owing to its persistent rubbery plateau above 72°C. As the amount of peroxide increases, the rubbery plateau modulus increases, allowing for enhanced mechanical energy storage, but the transition temperature and the steepness of the transition decrease. In the case of PCO with 10% DCP, shown as trace (v) in <figref idref="f0004">Fig. 3A</figref>, the thermomechanical response that is observed is inconducive to shape memory effects as the fixing (crystallization) temperature is lower than room temperature so that shape fixing would require subambient cooling and the temporary shape would be expected to drift via partial melting. In addition, the melting transition is too broad for dramatic strain recovery to be expected.

Suitable polymers include elastomers that are typically crosslinked and/or crystalline and exhibit melt or glass transitions at temperatures that are above body temperature and safe for use in the body, e.g. at about 40 to 50°C. Suitable polymers can have an elastic modulus of about 60,000 or 70,000 psi or more at 25°C (ASTM D638M). Such polymers may have a variety of room temperature moduli, from rigid glassy materials having storage moduli of several GPa to compliant rubbers with moduli as low as tens of MPa. Moreover, the moduli may tuned over the range 0.5 &lt; E &lt; 10 MPa, as dictated by the end application. Suitable polymers include polynorbornene, polycaprolactone, polyenes, nylons, polycyclooctene (PCO), blends of PCO and styrene-butadiene rubber, polyvinyl acetate/polyvinylidinefluoride (PVAc/PVDF), blends of PVAc/PVDF/polymethylmethacrylate (PMMA), polyurethanes, styrene-butadiene copolymers, polyethylene (particularly, crosslinked polyethylene), trans-isoprene, block copolymers of polyethylene terephthalate (PET) and blends of polycaprolactone and n-butylacrylate. A suitable PVAc/PVDF tube is formed by compounding 60-80 parts (by weight) PVAc (B-100, mw 500,000, ChemPoint, Cleveland, OH) with 40 to 20 parts PVDF (Grade 1010, Solvay Fluoropolymers, Houston, TX). The PVAc/PVDF is a crystalline material that can be utilized with or without crosslinking.

The polymer body can be made of mixtures of polymers or multiple polymer layers. The polymer forming the stent body can be compounded to include a drug, which elutes from the polymer, or radiopaque material. The structural polymer body can be coated with other polymers to carry drug or control drug delivery from the structural polymer. The polymer body can also exhibit shape memory properties. This feature of the polymer can be used in combination with the expansion properties discussed above. For example, the polymer can be configured to remember an enlarged or reduced diameter configuration. For example, the stent can be delivered into the body, and expanded by a combination of heat and radial pressure as described above. After a time, the stent can be retrieved by reheating the stent. In this case, the heating causes the stent to revert its small diameter condition. The remembered stent diameter is less than the vessel diameter and the stent can be more easily removed from the vessel. Such an application might be useful, for example, for a stent delivered into the prostate where removal and replacement is desirable.

Refer to <figref idref="f0006">Figs. 4</figref> and <figref idref="f0007">5</figref>, manufacture and use of the stent is illustrated. Referring particularly to <figref idref="f0006">Fig. 4</figref>, in a first step, a polymer tube 40 is constructed by extrusion or molding a suitable polymer to an initial diameter di which in the same or greater than the target lumen diameter. (For stents made of polymer strands, the strands can be formed by extrusion, followed by arranging the strands into a tube, e.g. by weaving or knitting.) The tube wall is then cut to provide a pattern of open areas in a desirable geometric pattern, e.g. by laser cutting. The polymer can be recrystallized or crosslinked, if necessary. Next, the tube is heated typically near or above the melt or glass transition and mechanically deformed to a small diameter, suitable for delivery. The tube is cooled, e.g. to room temperature. The tube is assembled onto a catheter, delivered into the body, and expanded by application of heat, to the melt or glass transition, while inflating the balloon as discussed above. (If the polymer has shape memory properties, the polymer tends to expand upon heating to a larger, remembered diameter.

Referring particularly to <figref idref="f0007">Fig. 5</figref>, a polymer tube 50 is constructed to an initial diameter di, smaller than the vessel diameter. The tube wall is cut to provide an open area in a desirable pattern so that when the stent is expanded, a compression-resistant geometry will result. The polymer is recrystallized or crosslinked, if necessary. The tube is assembled in a catheter, delivered into the body, and expanded by heating to the melt or glass transition temperature while inflating the balloon to provide an outward radial expansion, as discussed above. If the polymer has shape memory properties, the stent can be subsequently re-heated so it reverts back to its remembered small diameter configuration and removed from the body.

In particular embodiments, the stent can have an expanded inner diameter of about 2 to about 20 mm. The initial inner diameter can be in the range of about 1 to about 3 mm. The wall thickness is in the range of about 0.005 mm to 20 mm. The wall thickness variation between delivery and expanded states is within about 10%, 5% or 1% or less. The ratio of the open area to the wall area in the expanded state is about 0.5 to 0.7 or more. (The tube wall area is the area of the tube defined by polymer. The open area is the open area defined by the apertures.) A particular stent for coronary use has an initial diameter of about 2 mm, an expanded diameter of about 4 mm, and a wall thickness of about 0.005 mm to 0.1 mm. The stent can be used for various applications including coronary, neuro, carotid, peripheral vasculature, ureteral, and prostate lumens. The stent is particularly useful where the lumen path is highly curved or irregular.

The catheter can be, e.g. an angioplasty balloon catheter with a relatively non-distendable inflating balloon suitable for expansion of occluded regions of a vascular lumen. The balloon may include a polymer such as PET, which has a burst pressure of about 1.5 to 5 atm or more. The stent can be heated by heating the balloon inflation fluid. The balloon inflation fluid can be heated by e.g., heating the fluid delivery device outside the body using e.g., resistive heating tape. Alternatively, the catheter can carry a heating device. For example, a resistive heater or RF heater can be provided in the interior of the balloon. A heated balloon catheter is described in <patcit id="pcit0004" dnum="US5496311A"><text>Abele et al. U.S. 5,496,311</text></patcit> and <patcit id="pcit0005" dnum="US4955377A"><text>U.S. 4,955,377</text></patcit>. Vessel heating as described in the '311 can be used in combination with stent delivery as discussed herein. Alternatively, the stent can be heated directly. For example, the polymer can be compounded to include a material, such as magnetic particles, which are susceptible to heating by magnetic effects, such as hysteresis effects. A magnetic field can be imposed on the stent body by a source on a catheter or outside the body. Particles are available as the Smartbond System from Triton Systems, Inc., Chelmsford, MA. Heating by magnetic effects is discussed in <patcit id="pcit0006" dnum="US6056844A"><text>U.S. 6,056,844</text></patcit>. The stent can also be heated during delivery without applying expansion force to soften the stent, improving its flexibility and thus improving delivery to a treatment site through a tortuous vessel path.

Example

A polycyclooctene polymer (Vistenemer 8012 pellets, mw 90,000, Degussa, NJ) is melt processed in an extruder to produce a tube having dimensions of about 0.118 inch O.D. and 0.070 inch I.D. (wall thickness about 0.024 inch). The tube is cut to a length of about 4cm. The tube is subject to UV excimer laser ablation cutting to provide an aperture pattern of rectangular slots having a width of about 0.2 mm and a length of about 8 mm. Beam energy and pulse rate are selected to avoid substantial heating or melting of the polymer. The polymer can be compounded with about 10% TiO₂ (T8141, Dupont) to enhance absorption of laser radiation. A suitable pattern is consistent with the Express stent (commercially available from Boston Scientific, Natick, MA). (Alternatively, a pattern as described in <patcit id="pcit0007" dnum="US4733665A"><text>Palmaz U.S. 4,733,665</text></patcit> can be used.) The tube is heated to a temperature below its melt point, e.g., to about 39 to 40° C in a water bath and expanded by balloon catheter to a diameter of about 5 mm and positioned on a mandrel (PTFE tube) to maintain the expanded shape and diameter. The tube is then cooled to room temperature. The polymer is then crosslinked by e-beam radiation at 250 K Grays (Steris Isomedics Services, Northborough, MA). Crosslinking fixes the stent in the condition. (The crosslinked PCO has an elastic (Youngs) modulus of about 74945 psi at about 25°C (ASTM D638M)). The stent is heated to the polymer melt temperature, about 45° C and collapsed over a deflated balloon (diameter of about 2 mm) with a 4 mm inflated maximum diameter and 2cm length. (A suitable balloon catheter is a 75 cm Meditech UltraThin Catheter, available from Boston Scientific, Natick, MA.). The balloon and stent are immersed in a water bath of about 42 to 45° C and water the same temperature is used to inflate the balloon. The stent is expanded to about 4 mm diameter (ID) at an inflation pressure of about 1 to 1.5 atm (measured at the delivery syringe). After expansion, the heating is discontinued and the balloon and inflation fluid allowed to cool to body temperature (while the balloon remains inflated). Alternatively, a cooled contrast fluid can be circulated to the balloon. The stent exhibits no visible reduction in wall thickness or irregular flow of polymer into the stent open areas. In addition, in the heated, expanded state, the stent can be bent around a mandrel of about 0.75 cm radius without kinking. After the stent is cooled, it maintains the curved form.

Referring to <figref idref="f0008 f0009 f0010">Figs. 6-8</figref>, stent 100 includes a coiled rod 120 composed of a polymer. Stent 100 includes an elongated portion 140 having a general diameter d₁ and an end portion 160 having a maximum diameter d₂. General diameter d₁, for example, may be between about 3 mm and about 25 mm, more preferably between about 6 mm and about 14 mm, and maximum diameter d₂, for example, may be between about 7 mm and about 30 mm, more preferably between about 10 mm and about 17 mm. When stent 100 is designed, for example, for insertion into a urethra, the stent may have an overall length, for example, of between about 3 mm and about 15 mm, preferably between about 6 mm and 10 mm, and end portion 160 may have a length, for example, of from about 3 mm to about 15 mm, preferably from about 5 mm to about 10 mm. End portion 160 is in a flared position. Referring to <figref idref="f0008">Fig. 8</figref>, stent 100 is shown with end portion 160 in a collapsed position that can be reverted with heating to the expanded position shown in <figref idref="f0008">Fig. 6 and 7</figref>.

Generally, the portion of the stent in the collapsed position that can be reverted to the expanded position is, for example, greater than 5%, 10%, or even 25% of the overall length L of the stent, and less than 80% or 65% of the overall length L of the stent. For example, the of the overall length L of the stent may be between 10% and 65% of the overall length L of the stent.

The polymers preferably are cross-linked and/or crystalline elastomers that have melt or glass transition temperatures that are above body temperature, for example, greater than 45 °C or 55 °C. The degree of cross-linking can be used to adjust, for example, the melt or glass transition temperature, and range, of the polymer. The polymer preferably has a relatively narrow, for example, less that 5 °C or 10 °C, melt or glass transition temperature range.

The polymer preferably has elastomer properties in its melted or softened state. Preferred polymers have an elastic modulus, for example, of about 60,000 psi or 70,000 psi or more at 25 °C (ASTM D638).

Examples of polymers include polynorbornene and copolymers of polynorbornene, blends of polynorbornene with KRATON® (thermoplastic elastomer) and polyethylene, styrenic block copolymer elastomers (e.g., styrene-butadiene), polymethylmethacrylate (PMMA), polyethylene, polyurethane, polyisoprene, polycaprolactone and copolymers of polycaprolactone, polylactic acid (PLA) and copolymers of polylactic acid, polyglycolic acid (PGA) and copolymers of polyglycolic acid, copolymers of PLA and PGA, polyenes, nylons, polycyclooctene (PCO), polyvinyl acetate (PVAc), polyvinylidene fluoride (PVDF), blends of polyvinyl acetate/polyvinylidine fluoride (PVAc/PVDF), blends of polymethylmethacrylate/ polyvinyl acetate/polyvinylidine fluoride (PVAc/PVDF/PMMA) and polyvinylchloride (PVC).

In some embodiments, the polymers above are also useful for the stents of <figref idref="f0001">Figs. 1</figref>, <figref idref="f0006">4</figref> and <figref idref="f0007">5</figref>.

Particular polyurethanes are made by reacting (A) a polyol, (B) a chain extender dihydroxyl-terminated POSS and (C) a diisocyanate, where POSS stands for a polyhedral oligomeric silsesquioxane diol. The polyol (A) can be polyethylene glycol (PEG), polycaprolactone (PCL), polycyclooctene (PCO), trans-1,4 butadiene, transisoprene, polynorbornene diol and polymethacrylate copolymer, the chain extender(B) can be TMP cyclopentyldiol-POSS, TMP cyclohexyldiol-POSS, TMP isobutyldiol-POSS, trans-cyclohexanediolcyclohexane-POSS, or transcyclohexanediolisobutyl-POSS and the diisocyanate (C) can be selected from a large number of diisocyanates and is preferably 4,4' diphenyl methylene diisocyanate (MDI). Other diisocyanates (C) that are suitable for use in the synthesis of hybrid polyurethane SMPs include: toluene-2,4-diisocyanate (TDI), toluene-2,6diisocyanate, hexamethylene-1,6-diisocyanate (HDI), 4,4'diphenylmethane diisocyanate (MDI), isophorone diisocyanate (IPDI), and hydrogenate 4,4'-diphenylmethane diisocyanate (H12MDI).

The particular polyurethanes described directly above may be prepared the nonlimiting schemes illustrated below. A graph of heat flow as a function of temperature for several POSS polyurethanes is shown in <figref idref="f0009">Fig. 8A</figref> and a graph of storage modulus as a function of temperature for several POSS polyurethanes is shown in <figref idref="f0010">Fig. 8B</figref>. <chemistry id="chem0001" num="0001"><img file="EP1560612B1_D0001.tif" /></chemistry>

This scheme shows an example of synthesis of TPU using polyethylene glycol as polyol, TMP Isobutyldiol-POSS as chain extender to react with 4,4' diphenyl methylene diisocyanate in toluene. <chemistry id="chem0002" num="0002"><img file="EP1560612B1_D0002.tif" /></chemistry>

This scheme shows an example of synthesis of TPU using polycaprolactone diol as polyol, TMP Isobutyldiol-POSS as chain extender to react with 4,4' diphenyl methylene diisocyanate. <chemistry id="chem0003" num="0003"><img file="EP1560612B1_D0003.tif" /></chemistry>

This scheme shows an example of synthesis of TPU using polyocyclooctene as polyol, TMP Isobutyldiol-POSS as chain extender to react with 4,4' diphenyl methylene diisocyanate.

Any of the polymers mentioned above may be filled with, for example, nanoparticles of clay and silica to, for example, increase the modulus of the plastic. Dispersing agents and/or compatibilizing agents maybe used, for example, to improve the blending of polymers and the blending of polymers with fillers. Dispersing agents and/or compatibilizing agents include, for example, ACRAWAX® (ethylene bis-stearamide), polyurethanes and ELVALOY® (acrylic-functionalized polyethylene). The polymers can be cross-linked by application of radiation such as e-beam, UV, gamma, x-ray radiation or by heat-activated chemical crosslinking techniques. Radiation techniques provide the advantage that the polymer typically does not have to be substantially heated to achieve crosslinking. For e-beam radiation, an exposure of about 200-300, e.g. 250 ldlograys, typically provides sufficient crosslinking.

Referring to <figref idref="f0011">Fig. 9</figref>, wrapping fixture 200 can be used for making coiled stent 100. Wrapping fixture 200 includes a base 220 for support, a mandrel 240 with a flared end 260, slits 280 for fixing the plastic rod 120, an aperture 290 for fixing the plastic rod 120 on the non-flared end and a fixing screw 299 for releasably fixing mandrel 240 to wrapping fixture 200.

Stent 100 is manufactured from plastic rod 120 made by a variety of methods known in the art (e.g., extrusion, coextrusion, injection molding, casting, compression molding). If casting is used, the polymers may have tunable critical temperatures and rubber modulus have been synthesized using a thermosetting random copolymer formed from two vinyl monomers that yield controlled T_g and casting-type processing. Such copolymers were crosslinked with a difunctional vinyl monomer (crosslinker), the concentration of crosslinker controlling the rubber modulus and thus the work potential during recovery. Rod 120 may have a diameter, for example, of about 0.25 mm to about 2.5 mm or more. Plastic rod 120 and wrapping fixture 200 are heated to the softening temperature of the polymer, making the plastic rod 120 malleable. The rod 120 is inserted into aperture 290 that is machined through mandrel 240 to fix the rod at the starting end. The rod 120 is tightly wrapped around mandrel 240, including flared end 260. To fix plastic rod 120 in place, plastic rod 120 is pushed into slits 280. The overall length of the stent 100 may be, for example, about 2 mm to about 150 mm or more. The overall length required depends upon the application. The plastic rod 120, now fixed in place on the mandrel 240, is heated to above the softening point of the material and maintained at that temperature long enough to anneal the rod and fix the shape. Typically, the time required to fix the shape is from about 0.25 hr to 10 hr or more. After cooling, stent 100 is removed from mandrel 240. Before packaging, the flared end of the coil is tapered down and collapsed so that the diameter along the entire length of the stent is approximately d₁. Collapsing the flared end 160 of stent 100 allows for ease of insertion, for example, into a restricted prostatic urethra.

EXAMPLE 1

A 56:24:20 mixture of PVAc/PVDF/PMMA is dry bended and loaded into the hopper of an extruder. The PVAc is grade B-100, the PVDF is Solvay SOLEF® 1010 and the PMMA is Atofina PLEXIGLAS® V045. The mixture is melt processed to produce 1.27 mm (0.05 inch) monofilament. The rod is made into a coil by winding it around wrapping fixture 200. The fixture and the rod are immersed into a 50 °C water bath. At this temperature, the rod becomes malleable enough to wind easily around the mandrel and secured in place to prohibit the uncoiling of the helical shape. The mandrel is removed from the fixture with the stent locked in place and placed into an oven at 110 °C for one hour to anneal the stent. This annealing process locks the permanent shape of the coil. The mandrel and coil are cooled to room temperature, and the stent is removed from the mandrel. The stent had on overall length of approximately 73 mm and a flared end portion length of approximately 7 mm. The diameter d₁ of the body is approximately 6 mm and the maximum diameter d₂ of the flare on the open end is approximately 11 mm. Before packaging, the flared end of the coil is tapered down with brief heating to 50 °C and manipulation, followed by cooling, so that the diameter is approximately 6 mm along the entire length of the stent.

EXAMPLE 2

A 70:30 mixture of PVAc/PVDF is dry bended and loaded into the hopper of an extruder. The mixture is melt processed to produce 1.27 mm (0.05 inch) monofilament. The rod is made into a coil by winding it around wrapping fixture 200. The fixture and the rod are immersed into a 50 °C water bath. At this temperature, the rod becomes malleable enough to wind easily around the mandrel and secured in place to prohibit the uncoiling of the helical shape. The mandrel is removed from the fixture with the stent locked in place and placed into an oven at 110 °C for one hour to anneal the stent. This annealing process locks the permanent shape of the coil. The mandrel and coil are cooled to room temperature, and the stent is removed from the mandrel. The stent had on overall length of approximately 73 mm and a flared end portion length of approximately 7 mm. The diameter d₁ of the body is approximately 6 mm and the maximum diameter d₂ of the flare on the open end is approximately 11 mm. Before packaging, the flared end of the coil is tapered down with brief heating to 50 °C and manipulation, followed by cooling, so that the diameter is approximately 6 mm along the entire length of the stent.

Referring to <figref idref="f0012 f0013">Fig. 10-12</figref>, stent 100 may be, for example, inserted into restricted urethra 300 on delivery tube 320. During insertion, end portion 160 is in a collapsed position. After insertion, warm water (e.g., 45 °C-55 °C) is flushed through delivery tube 320 that is in thermal contact with stent 100. Heating reverts the collapsed end 160 to a flared, expanded position (<figref idref="f0013">Fig. 12</figref>). The flared, expanded position allows stent 100 to remain fixed in position, for example, in the prostatic urethra.

Referring to <figref idref="f0014">Fig. 13</figref> for a little more detail, delivery tube 320 is a long cylindrical tube into which a ureteral scope 380 is inserted. Delivery tube 320 has a distal end 330 over which stent 100 is placed. Delivery 320 is fitted with a side port 350 including a stopcock 370 through which saline can be flushed for irrigation. Delivery tube 320 with stent 100 in place is delivered into, for example, the prostatic urethra with the aid of a ureteral scope. Once stent 100 is in place, hot saline is flushed through port 350 to revert the collapsed end 160 to a flared, expanded position (<figref idref="f0013">Fig. 12</figref>). The flared, expanded end allows stent 100 to remain fixed in position, for example, in or adjacent the prostatic urethra or external sphincter between the prostate and the bladder to prevent migration. The direction of the flare can, of course, be oriented in other directions. The scope and delivery tube 320 are withdrawn, leaving stent 100 in place.

Referring to <figref idref="f0014 f0015">Figs. 13-13A</figref>, an alternative delivery system is illustrated that includes a tube 320 with a screw on tip 331 onto which a stent 100 is placed after collapsing the end portion 160. The assembly is inserted into, for example, the prostatic urethra.

Other delivery methods are within the claims. Stent 100 may be, for example, inserted into restricted urethra 300 on balloon catheter (not shown). During insertion, end portion 160 is in a collapsed position. After insertion, warm water is flushed through the guide wire lumen of the balloon catheter to flood the area and to heat the stent. Heating of the stent by the water reverts the collapsed end 160 to a flared, expanded position The flared, expanded position allows stent 100 to remain fixed in position, for example, in the prostatic urethra. If there is an obstruction in the lumen into which the stent is deployed the stricture can be dilated using the balloon to help the stent open fully and maintain a uniform diameter inside the vessel.

Referring to <figref idref="f0016">Fig. 14</figref>, coiled stent 100 with an end 160 in the expanded position can be removed with the aid of a catheter equipped with a grasping device 360 and a ureteral scope 380 for visualizing stent 100 in, for example, the prostatic urethra. Once end 400 of stent 100 has been visualized with ureteral scope 380, the stent is grasped with grasping device 360. Next, ureteral scope 380 is removed from the catheter and is replaced with a heating device (not shown), e.g. a catheter. Heating stent 100 above the softening point of the polymer, e.g., from about 45 °C to about 55 °C for polycyclooctene (PCO), and pulling the end 400 of stent 100 through the orifice 420 allows the stent to be removed in a substantially uncoiled state.

Although <figref idref="f0013">Fig. 11</figref> shows heating of stent 100 with a warm liquid on a delivery tube, other heating methods are possible. For example, heating may be accomplished with the use of IR, RF or inductive heating.

Although insertion into a prostatic urethra has been used as an example, insertion of stent 10 into other body lumens or cavities is possible. For example, other body lumens or cavities include the biliary duct, cystic duct, a ureter, a bulbar urethra or a hepatic duct.

Referring to <figref idref="f0017">Fig. 15</figref>, the modulus of polycyclooctene (PCO) that can, for example, be the polymer of stent 100 is shown as a function of temperature. Below approximately -65 °C (T_g, region A), PCO exists a rigid, glassy polymer. Above T_g, but below T_m, PCO exists as a flexible elastomer (region B). Above T_m, PCO exists as a relatively low modulus elastomer. Above T_m, for example, stent 100 composed of PCO can be removed from a lumen or cavity of the body, the prostatic urethra, for example, in a substantially uncoiled state.

Referring to <figref idref="f0018">Fig. 16</figref>, the modulus of a ternary blend of PVAc/PVDF/PMMA that can, for example, be the polymer of stent 100 is shown as a function of temperature. Adding PMMA offers the advantage, for example, of increasing the modulus of the blend.

Referring once again to <figref idref="f0017">Fig. 15</figref> and to <figref idref="f0008 f0009 f0010">Figs. 6-8</figref>, heating stent 100 in the expanded position briefly above T_m (region C) and then cooling rapidly below T_m (e.g., region B) "freezes" stress into stent 100. Stent 100 reverts from its collapsed position to its expanded position upon re-heating above T_m (region C) because the modulus of the material lowers sufficiently to release the residual stress that was "frozen" into stent 100 during the rapid cooling.

<figref idref="f0019 f0020 f0021 f0022">Figs. 17-28</figref> show other examples of stents.

Referring to <figref idref="f0019">Figs. 17 and 18</figref>, coiled stent 500 has two end portions 520 and 540 in expanded positions. End portions 520 and 540 can be collapsed (<figref idref="f0019">Fig. 18</figref>) for ease of insertion into a body cavity or lumen, and reverted with heat to expanded positions.

Referring to <figref idref="f0019">Fig. 19</figref>, coiled stent 600 has two end portions 620 and 660 and a central portion 640 in expanded positions. All three portions may be collapsed (not shown), and then reverted to expanded positions.

Referring to <figref idref="f0020">Figs. 20 and 21</figref>, coiled stent 700 made from a flattened tube 710 has an end portion 720 in an expanded position. The flattened tube can, for example, add strength to the stent. The tube can have a major diameter, for example, of between about 1.0 mm to about 3.0 mm, more preferably between about 1.5 mm to about 2.25 mm, and major inner diameter, for example, of between about 0.5 mm to about 2.5 mm, more preferably between about 1.25mm and about 1.75 mm. End portion 720 may be collapsed (not shown) and then reverted to an expanded position. Referring to <figref idref="f0020">Fig. 21</figref>, flattened tube 710 has an interior 740 that may be filled with, for example, a medicament. The medicament, for example, may be triclosan or salicylic acid. Release of medicament from flattened tube 710, for example, may reduce the risk of infection. Interior 740 may also be filled with, for example, paclitaxel or mitoxantrone. Release of the these medicaments from interior 740 may be, for example, useful for treating prostate cancer and reducing prostatic hyperplasia.

Referring to <figref idref="f0020">Fig. 22</figref>, coiled stent 800 made from tape 810 has an end portion 820 that may be collapsed and then reverted to an expanded position. Tape 810, for example, may have a thickness from about 0.5 mm to about 2.0 mm, more preferably from about 0.75 mm to about 1.25 mm, and a width, for example, of from bout 1.0 mm to about 3.0 mm, more preferably from about 1.75 mm to about 3.00 mm. In this particular embodiment, an aperture 840 is provided so that a trailing string (not shown) can be included for ease of removal with the grasping device shown in <figref idref="f0016">Fig. 14</figref>.

Referring to <figref idref="f0021">Figs. 23-25</figref>, coiled stent 900 is made from plastic rod and has an end portion 920 that may be collapsed and then reverted to an expanded position. Stent 900 includes a plurality of protruding elements 940 that are integral with and extend outwardly from the plastic rod from which the stent is made. Referring now to <figref idref="f0022">Figs. 26-28</figref>, coiled stent 1000 is made from plastic rod, for example, oriented monofilament, and has an end portion 1020 that may be collapsed and then reverted to an expanded position. Stent 1000 includes a plurality of protruding elements 1060 that extend outwardly from the plastic rod from which the stent is made. The friction provided by protruding elements 940,1060 can help to hold stent 900,1000 in place within, for example, the prostatic urethra.

Protruding elements 940 are made, for example, by cutting into the plastic rod with, for example, a sharp edged instrument, for example, a knife, so as penetrate a depth into the plastic rod. The depth of penetration P is adjusted to provide acceptable frictional properties, while minimizing the impact on the mechanical properties of the plastic rod. In some implementations, the maximum depth of penetration into the plastic rod, measured inwardly from the outer surface of the plastic rod is, for example, from about 1 to about 50% of the average thickness of the plastic rod. If the depth penetration is too large, the mechanical properties of the plastic rod may be reduced and if the depth of penetration is too low, the resulting protruding elements may be too small to provide the appropriate frictional properties when expanded in a body cavity or lumen, for example, the prostatic urethra. Other cutting means are possible, for example, water knife and laser cutting means, to reduce the impact of the cutting on the mechanical properties of the plastic rod. The shape of the plastic rod from which the stent is made may be of other forms than that shown above. For example, it may be in the form of, for example, a coiled elongated flattened tube and the flattened tube may include a central opening that includes a medicament that can be released by the inserted stent.

In some implementations, stent 900 is manufactured from plastic rod made by a variety of methods known in the art (e.g., extrusion and coextrusion). The plastic rod may have a diameter, for example, of about 0.25 mm to about 2.5 mm or more. In a preferred method, the protruding elements are put onto the plastic rod before wrapping the mandrel shown and discussed above. After wrapping the mandrel, the plastic rod and wrapping fixture 200 are heated to the softening temperature of the polymer, making the plastic rod malleable. The protrusions are annealed in the "up" position, that is, with the protruding elements extending outwardly by "prying up" the protruding elements that results from cutting. Prying up the protruding elements may be achieved by, for example, running a surface across the protruding elements in a direction opposite the cut direction. Annealing is continued to fix the shape. After cooling, the stent is removed from mandrel. Before packaging, the flared end of the coil is tapered down along with protruding elements and collapsed so that the diameter along the entire length of the stent is approximately d₁. Collapsing the flared end and protruding elements allows for ease of insertion, for example, into a restricted prostatic urethra. In some implementations, the flared end and the main body are collapsed to have a diameter less than d₁. Upon heating, the end portion reverts to its expanded position and the protruding elements revert to their up positions. If a medicament is placed in the cavities 960 from which the protruding elements 940 are carved, it may be released upon expansion of the stent.

Stent 1000 with protruding elements 1060 is made by, for example, wrapping a thicker plastic rod with a thinner plastic rod, for example, a monofilament, that includes a plurality of constrictions, for example, knots along its length. The elevation E above an outer surface of the thicker plastic rod is adjusted to provide acceptable frictional properties. In some implementations, the maximum elevation above an outer surface of the thicker plastic rod is, for example, from about 1 to about 50% of the average thickness of thicker plastic rod. If the elevation E is too large, insertion of stent 1000 into, for example, a prostatic urethra may become difficult and if the maximum elevation above the protruding elements is too small, the protruding elements may not provide the appropriate frictional properties when expanded in a body cavity or lumen, for example, the prostatic urethra. The shape of the rods from which the stent is made may be of other forms than that shown above. For example, it may be in the form of, for example, a coiled elongated flattened tube and the flattened tube may include a central opening that includes a medicament that can be released by the inserted stent.

In some implementations, stent 1000 is manufactured from plastic rod made by a variety of methods known in the art (e.g., extrusion and coextrusion). The thicker plastic rod may have a diameter, for example, of about 0.25 mm to about 2.5 mm or more. The thinner plastic rod from which the protruding elements are fashioned may have a diameter of, for example, from about 0.2 mm to about 20 mm. In a preferred method, the constrictions, for example, knots, are placed on the thinner plastic rod and the thinner plastic rod is wrapped around the outer surface of the thicker plastic rod. The ends of the thinner plastic rod are heat staked to hold the thinner plastic rod onto the outer surface of the thicker plastic rod. Now, the assembly of the thinner and thicker plastic rod is wrapped around the mandrel shown and discussed above. After wrapping the mandrel, the plastic rods and wrapping fixture 200 are heated to the softening temperature of the polymer of the thicker plastic rod, making the plastic rod malleable. Annealing is continued to fix the shape. After cooling, the stent is removed from mandrel. Before packaging, the flared end of the coil is tapered down so that the diameter along the entire length of the stent is approximately d₁. Collapsing the flared end and protruding elements allows for ease of insertion, for example, into a restricted prostatic urethra. In some implementations, the thinner plastic rod may contain a medicament that is released upon expansion, for example, in the prostatic urethra. In other implementations, the thinner plastic rod is made of a degradable material and the degradable material is filled with a medicament.

In some embodiments, the entire stent, for example, the stents of <figref idref="f0006">Figs. 4</figref>, <figref idref="f0007">5</figref>, <figref idref="f0020">20</figref>, <figref idref="f0021">23</figref> and <figref idref="f0022">26</figref>, may have an expanded position.

Other Embodiments

In some of the embodiments of any of the above stents, only a portion or portions of the stent (e.g., the portion(s) having an expanded position) may be composed of the polymer. The remainder of the stent may be, for example, composed of a non-polymeric material (e.g., a metal or metal alloy, e.g., Ni/Ti alloy). Moreover, the stent may be composed of multiple layers of materials, for example, by coextruding the layers when making an elongated element. The stent may be a multiple segment stent.

The polymer in any of the above stents may be a blend of polymers, for example, miscible blends of a semicrystalline polymers with an amorphous polymer. For those blends that are miscible at the molecular level, a single glass transition results, without broadening. Additionally, in such miscible blends the equilibrium crystallinity (which controls the plateau modulus between T_g and T_m where shape fixing is performed) also changes dramatically and systematically with blend composition; i.e., relative levels of each component.

Polymers blends with a relatively high modulus in the fixed state at room temperature, having a tunable and sharp transition, the permanent shape of which can be remolded repeatedly above certain melting temperatures are prepared by the blending of crystalline polymers (C') with amorphous polymers (A'), such that they are a single miscible phase in the molten state (allowing processing to stress-free native states) but crystalline to a limited and tailored extent and which further vitrify on cooling to room temperature. The recovery of the polymer blend may be fast, for example, within seconds. Examples for (C') include poly(vinylidene fluoride) (PVDF) (T_g = -35 °C, T_m = 175 °C), polylactide (PLA) (T_g = 56 °C, T_m = 165 °C), poly(hydroxy butyrate), poly(ethylene glycol) (PEG), polyethylene, polyethylene-co-vinyl acetate, poly(vinyl chloride) (PVC), and poly(vinylidene chloride) (PVDC) and copolymers of poly vinylidene chloride (PVDC)/poly vinyl chloride (PVC). Examples for (A') include poly(vinyl acetate) (PVAc) (T_g = 35 °C), poly methyl acrylate (PMA), poly ethyl acrylate (PEA), atactic poly methyl methacrylate (aPMMA), isotactic poly methyl methacrylate (iPMMA), syndiotactic poly methyl methacrylate (sPMMA), and other poly alkyl methacrylates.

In some preferred embodiments formed from two miscible polymer blends, the blend is prepared by mixing amorphous poly(vinyl acetate) (PVAc) (T_g = 35 °C) with semicrystalline polylactide (PLA) (T_g = 56 °C, T_m = 165 °C) or poly(vinylidene fluoride) (PVDF). The polymers show complete miscibility at all blending ratios with a single glass transition temperature, while crystallization (exclusive of PVAc) is partially maintained. The T_g's of the blends are employed as the critical temperature for triggering the shape recovery while the crystalline phases serve as physical crosslinking sites for elastic deformation above T_g, but below T_m.

The preferred blends are formed from poly vinyl acetate(PVAC) and poly(lactic acid) (PLA) or poly(vinylidene fluoride) (PVDF). However, examples of other suitable blends include the pair PVDF/PMMA and ternary blends of PVDF/PMMA/PVAc. The PMMA and the combination of PMMA/PVAc serve the same role as PVAc in the blends as have been previously described. An advantage of adding PMMA is that the critical temperature can be increased arbitrarily to about 80 °C and the room temperature modulus can also be increased. The PVDF may be substituted by poly(vinylidene chloride) (PVDC), by copolymers of poly(vinylidene chloride/ply(vinyl chloride), or by any "C", polymer vide supra.

It has further been found that blending poly(vinyl chloride) with poly(butyl acrylate) or poly (butyl methacrylate) (PVC/PBA) has certain advantages. In the PVDF/PVAc case, PVAc simultaneously lowers the crystallinity of PVDF while increasing T_g. PVC may serve the same role as PVDF, but it already has a low degree of crystallinity, but a relatively high T_g (∼ 80 °C). Thus in this embodiment, the second component (PBA) serves only the role of decreasing T_g. This can also be achieved with smal molecule plasticizers, most notably dioctylphthalate (DOP), but is preferred to use a biocompatible polymeric plasticizer for intended implantable applications. The range of PBA compositions is 10-40%, with 20% being the most advantageous, yielding a T_g ∼ 40 °C.

Melt blending of PLA/PVAc and PVDF/PVAc of varying blend ratios was performed in a 30 ml Brabender mixer. The mixer was equilibrated at T = 180 °C for 5 minutes after which the mixer blade rotation was adjusted to 25 RPM and the premixed polymers pellets added to the chamber over the course of 1 minute. The polymers were mixed for 10 minute to ensure good dispersion. Nitrogen was purged through the chamber to mitigate potential oxidative degradation during mixing. After mixing, the blend was taken out of the chamber, cooled to room temperature, and then pressed between heated platens of a Carver press at 180 °C for 5 minutes under a load of 8 metric tons. A spacer was used to control the thickness of the film and rapid cooling to room temperature was carried out. The films thus formed were used for the subsequent thermal and mechanical characterization.

The TGA results demonstrated that both PLA and PVAc are stable for T &lt; 300 °C. Above this temperature PLA degrades completely (no char yield), while the PVAc degrades to yield an intermediate char yield of 25 wt% for 375 &lt; T &lt; 425 °C but complete degradation above 450 °C. Blend processing and thermal and dynamic mechanical analyses (DSC and DMA) were performed below 250 °C, to completely avoid degradation.

The crystallization behavior of semicrystalline PLA was investigated via DSC. The PLA samples were first heat pressed at 180 °C for 10 minutes and then quenched to room temperature with water cooling. One sample was directly analyzed by DSC, while another was first annealed at 110 °C (=1/2(T_g + T_m)) for 1 hour to reach an equilibrium level of crystallinity. <figref idref="f0023">Fig. 29</figref> shows a comparison of thermal behavior for these two samples. It was observed that quenching the PLA melt results in a low degree of crystallinity and virtually no recrystallization on heating, both indicating slow crystallization. Annealing at 110 °C for 1 hour results in significant crystallization evidenced by a large melting endotherm at T = 155 °C. The melting temperature did not shift dramatically due to annealing, but the endotherm shape did change. Complementary WAXD experiments yielded the same conclusions.

The crystallization behavior selected of polymer blends was also analyzed. All of the samples were heat pressed at 180 °C for 10 minutes and then annealed at 110 °C for 1 hour before thermal analysis, providing a standard condition for extensive crystallization. <figref idref="f0024">Fig. 30</figref> summarizes the first DSC heating trace of the samples measured after annealing. The results indicate that PVAc itself is amorphous (though with large physical aging content) but that incorporation of PLA leads to crystallization in proportion to the PLA wt-%. Also, the peak endotherm positions (melting transitions) shift slightly to higher temperatures with increasing PLA content. Quenching these samples to T = 20 °C and reheating to 200 °C again showed clearly that single T_g's are observed and that crystallization can be largely suppressed. Importantly for shape memory, the single glass transition events were not broadened in the blends relative to the pure components, suggesting that the amorphous phase was quite homogeneous for all of the blends. The observed T_g values are plotted in <figref idref="f0025">Fig. 31</figref> along with the best fit with the Fox equation, showing slight positive deviation. This leads to a conclusion that strong interactions between the two polymers that reduces free volume of the polymer blends and hence, increased glass transition temperature relative to the Fox equation prediction has occurred.

In order to elucidate the effect of PVAc on the degree of crystallinity and the crystal structures, the crystalline diffraction patterns were observed via wide-angle x-ray diffraction. The results indicate that the PVAc phase has only an amorphous halo, thus being totally amorphous, while the PLA exhibits three very strong diffraction peaks at 2θ = 22.3°, 25.0°and 28.6°, corresponding to d-spacings of 5.92, 5.29, and 4.64 A°, respectively. Upon addition of PVAc, all of the peak intensities were depressed, but the peak positions remained essentially unchanged. Consistent with the DSC results, the degree of crystallinity increases in proportion to PLA addition. From the peak width at half height, it was found that the crystalline lamellae size did not decrease, although the degree of crystallinity decreased, with increasing PVAc content. This means that the decrease in crystallinity and depression of the melting transitions are not due to a change of crystal size, but rather may be due to a thinning of the lamellae thickness or to a decrease of the crystal concentration.

The storage modulus of the polymer blends was also measured using DMTA, first investigating the effects of annealing on the storage modulus. Below their glass transition temperatures, T_g, both samples exhibit similar high storage moduli (3 GPa), as well as similar softening points. When heated above T_g, the storage modulus of thermally quenched samples decreases sharply to about 2 MPa; however, further increasing the temperature induces a modulus increase attributed to recrystallization of the samples at higher temperatures. This also proved that the sample is not in an equilibrium state and that its mechanical properties in the rubbery region depend on thermal history. To reach equilibrium, the sample was annealed at 110 °C for 1 hour as previously described for DSC analyses. The storage modulus above T_g shifts to about 200 MPa until melting, the increase being due to an increase of the degree of crystallinity on annealing to tune the rubbery modulus at equilibrium state,. PLA was blended in different proportions to PVAc and annealed as above. Storage moduli for such blends were measured and the results are plotted in <figref idref="f0026">Fig. 32</figref>. It can be seen that, below T_g, all of the samples show similar large moduli while above T_g the moduli decrease to a plateau whose magnitude depends on crystallinity and thus PLA content. This trend is in accordance with that of DSC and XRD, and can be explained by the fact that the increase of storage moduli came from the physical crosslinking formed by crystals and the filler effect of the high modulus crystalline phase.

Stress-free shape memory tests were carried out in hot water at 65 °C, with an annealed sample composed of 30 % PLA. The results show that the sample features quick and complete shape memory behavior: the sample recovers to the original shape (straight bar) within 10 seconds, with most of the recovery being accomplished within the first several seconds.

The same characterizations were carried out on the blends of PVDF and PVAc as above disclosed. The TGA and DSC results show that PVDF is also thermally stable up to 300 °C, and the mixtures form only one glass transition, the values fall between the Tgs of the two homopolymers and changes with changing composition. At the same time, the melting points and the degrees of crystallinity were depressed with the incorporation of amorphous PVAc.

The storage moduli of the blends, which give the rigidity of the materials, were also measured. The results are similar to those of the PLA/PVAc blends, the PVDF/PVAc blends being very rigid below the critical temperatures (T_g), and featuring a sharp modulus changes at the Tg to a plateau modulus ranging from several MPa to tens of MPa, depending on the degree of crystallinity of the blends. These plateau moduli can be tuned by adjusting the degree of crystallinity of the blend, that is, adjust the blend composition.

The polymer in any of the above stents may be bioabsorbable or non-bioabsorbable. Bioabsorbable polymers include, for example, polyurethanes and polyurethane copolymers such as those described above with the general formula (directly below), where X/Y is, for example, 1 to 20, n is, for example, 2 to 1000, and the total degree of polymerization m is, for example, 2 to 100 <chemistry id="chem0004" num="0004"><img file="EP1560612B1_D0004.tif" /></chemistry>

The bioabsorbability of the polymers is enhanced by copolymerization of polyurethane and POSS with suitable monomers. Examples of suitable monomers include caprolactone, ethyleneglycol, ethylene oxide, lactic acid, and glycolic acid. The copolymers from these monomers can hydrolyze and cleave the polymer linkage.

Other embodiments of stents can also be formed to include materials described above. In some embodiments, an implantable medical stent may be delivered into the body using a catheter. The stent can be delivered in a small diameter form and then expanded at a treatment site by triggering a shape change (for example, by heat application) caused by the shape memory properties of the polymer. The stent can also be expanded by a mechanical expander such as an inflatable balloon of the type used on an angioplasty catheter.

In some embodiments, the stent is sized (e.g., an expanded inner diameter of about 2 mm to about 20 mm) and configured for use in the vascular system, particularly the coronary arteries, and implanted after or simultaneously with an angioplasty procedure to maintain an open lumen and reduce restenosis. Vascular stents are described in <patcit id="pcit0008" dnum="US418023P" dnum-type="L"><text>U.S. Provisional Application Number 60/418,023</text></patcit>. For example, a stent for coronary use can have an initial diameter of about 2 mm, an expanded diameter of about 4 mm, and a wall thickness of about 0.005 mm to 0.1 mm. Other exemplary applications include neuro, carotid, peripheral, and vasculature lumens. The vascular stent can be bioabsorbable or non-bioabsorbable.

In other embodiments, a stent, e.g., a bioabsorbable or a non-bioabsorbable stent, is constructed for use in nonvascular lumens, such as the esophagus, ureteral, biliary, or prostate.

In other embodiments, the stent is conductive to allow an electrical current to pass through the stent, for example, to deliver electricity to an area of the body or to trigger, for example, a physical change in the stent, for example, a change in the diameter of the stent.

In still other embodiments, the stent, for example, of <figref idref="f0006">Figs. 4</figref>, <figref idref="f0021">23</figref> and <figref idref="f0022">26</figref>, is made porous by, for example, adding a chemical foaming agent to the polymer from which the stent is made during the production the plastic strand. In some implementations, the stent is porous and includes a medicament. The initial porosity of the stent can be reduced, for example, by the application of heat and pressure before deployment in the body. Upon deployment of the stent in the body, the porosity is increased by a triggering event, for example, the application of heat to the stent at the desired site of treatment.