Patent Application
20130317596
1. Field of the Invention
This invention relates to bioresorbable polymer scaffolds and methods of treatment of coronary lesions with bioresorbable polymer scaffolds
2. Description of the State of the Art
This invention relates generally to methods of treatment with radially expandable endoprostheses, that are adapted to be implanted in a bodily lumen. An “endoprosthesis” corresponds to an artificial device that is placed inside the body. A “lumen” refers to a cavity of a tubular organ such as a blood vessel. A stent is an example of such an endoprosthesis. Stents are generally cylindrically shaped devices that function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of a bodily passage or orifice. In such treatments, stents reinforce body vessels 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.
Stents are typically composed of a scaffold or scaffolding that includes a pattern or network of interconnecting structural elements or struts, formed from wires, tubes, or sheets of material rolled into a cylindrical shape. This scaffold gets its name because it physically holds open and, if desired, expands the wall of a passageway in a patient. Typically, stents are capable of being compressed or crimped onto a catheter so that they can be delivered to and deployed at a treatment site.
Delivery includes inserting the stent through small lumens using a catheter and transporting it to the treatment site. Deployment includes expanding the stent to a larger diameter once it is at the desired location. Mechanical intervention with stents has reduced the rate of restenosis as compared to balloon angioplasty.
Stents are used not only for mechanical intervention but also as vehicles for providing biological therapy. Biological therapy uses medicated stents to locally administer a therapeutic substance. The therapeutic substance can also mitigate an adverse biological response to the presence of the stent. A medicated stent may be fabricated by coating the surface of either a metallic or polymeric scaffolding with a polymeric carrier that includes an active or bioactive agent or drug. Polymeric scaffolding may also serve as a carrier of an active agent or drug.
The stent must be able to satisfy a number of mechanical requirements. The stent must have sufficient radial strength so that it is capable of withstanding the structural loads, namely radial compressive forces imposed on the stent as it supports the walls of a vessel. Radial strength, which is the ability of a stent to resist radial compressive forces, relates to a stent's radial yield strength and radial stiffness around a circumferential direction of the stent. A stent'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.
Once expanded, the stent must adequately provide lumen support during a time required for treatment in spite of the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. In addition, the stent must possess sufficient flexibility with a certain resistance to fracture.
Stents made from biostable or non-degradable materials, such as metals that do not corrode or have minimal corrosion during a patient's lifetime, have become the standard of care for percutaneous coronary intervention (PCI) as well as in peripheral applications, such as the superficial femoral artery (SFA). Such stents have been shown to be capable of preventing early and later recoil and restenosis.
In order to effect healing of a diseased blood vessel, the presence of the stent is necessary only for a limited period of time, as the artery undergoes physiological remodeling over time after deployment. The development of a bioabsorbable stent or scaffold could obviate the permanent metal implant in vessel, allow late expansive luminal and vessel remodeling, and leave only healed native vessel tissue after the full resorption of the scaffold. Stents fabricated from bioresorbable, biodegradable, bioabsorbable, and/or bioerodable materials such as bioabsorbable polymers can be designed to completely absorb only after or some time after the clinical need for them has ended. Consequently, a fully bioabsorbable stent can reduce or eliminate the risk of potential long-term complications and of late thrombosis, facilitate non-invasive diagnostic MRI/CT imaging, allow restoration of normal vasomotion, provide the potential for plaque regression.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, wherein the segment comprises a scaffolded segment between a proximal and a distal end of the scaffold, a proximal segment proximally adjacent to the proximal end of the scaffold, and a distal segment distally adjacent to the distal end of the scaffold, wherein the proximal segment exhibits constrictive remodeling between baseline and two years after the deployment, wherein the constrictive remodeling comprises a decrease in a cross-sectional area of the proximal segment.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, wherein the segment comprises a scaffolded segment between a proximal and a distal end of the scaffold, a proximal segment proximally adjacent to the proximal end of the scaffold, and a distal segment distally adjacent to the distal end of the scaffold, wherein a content of fibrotic and fibrofatty (FF) tissue increases at the distal segment between baseline and two years after the deployment.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, wherein the segment comprises a scaffolded segment between a proximal and a distal end of the scaffold, a proximal segment proximally adjacent to the proximal end of the scaffold, and a distal segment distally adjacent to the distal end of the scaffold, and wherein at baseline there is a difference in a compliance of the scaffolded segment between the proximal segment and the distal segment.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, the polymer scaffold expanding during deployment which expands the segment to a target diameter, wherein vasomotion of the segment of the artery reappears after deployment due to the replacement of the polymer by de novo formation of malleable tissue comprising proteoglycan, wherein two years after deployment the scaffold area or volume has decreased by less than 10%.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, the polymer scaffold expanding during deployment which expands the segment to a target diameter, wherein a neointimal area increases and a mean scaffold area increase between baseline and 1 year and between one year and three years after baseline.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, the polymer scaffold expanding during deployment which expands the segment to a target, wherein a total plaque area in the segment increases between baseline and one year and then decreases between one and three years after deployment.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, the polymer scaffold expanding during deployment which expands the segment to a target diameter, and wherein a dense calcium percent and a hyperechogenic area of the segment decreases between baseline and 1 year and between one year and three years.
Embodiments of the invention include a method of treating vascular disease in a patient comprising: deploying a bioabsorbable polymer scaffold composed of a plurality of struts at a segment of an artery of a patient, the polymer scaffold expanding during deployment which expands the segment to a target diameter, wherein 3 years after deployment, the segment comprises: return of vasomotion to the segment; enlargement of the scaffold area and mean lumen area between baseline and three years; an increase of neointima in the segment between baseline and three years; and a reduction of plaque area between baseline and three years.
All publications patents, and patent applications mentioned in this specification are herein incorporated by reference to the same extent as if each individual publication, patents, or patent application was specifically and individually indicated to be incorporated by reference, and as if each said individual publication, patents or patent application was fully set forth, including any figures, herein.
Various embodiments of the present invention include treatment of coronary artery disease with bioresorbable polymer stents. The bioresorbable stents can include a support structure in the form of a scaffold made of a material that is bioresorbable, for example, a bioresorbable polymer such as a lactide-based polymer. The scaffold is designed to completely erode away from an implant site after treatment of an artery is completed. The scaffold can further include a drug, such as an antiproliferative or anti-inflammatory agents. A polymer coating disposed over the scaffold can include the drug which is released from the coating after implantation of the stent. The polymer of the coating is also bioresorbable.
The present invention is applicable to, but is not limited to, self-expandable stents, balloon-expandable stents, stent-grafts, and generally tubular medical devices in the treatment of artery disease. The present invention is further applicable to various stent designs including wire structures, and woven mesh structures.
Self expandable or self expanding stents include a bioabsorbable polymer scaffold that expands to the target diameter upon removal of an external constraint. The self expanding scaffold returns to a baseline configuration (diameter) when an external constraint is removed. This external constraint could be applied with a sheath that is oriented over a compressed scaffold. The sheath is applied to the scaffold after the scaffold has been compressed by a crimping process. After the stent is positioned at the implant site, the sheath may be retracted by a mechanism that is available at the end of the catheter system and is operable by the physician. The self expanding bioabsorbable scaffold property is achieved by imposing only elastic deformation to the scaffold during the manufacturing step that compresses the scaffold into the sheath.
The bioabsorbable scaffold may also be expanded by a balloon. In this embodiment the scaffold is plastically deformed during the manufacturing process to tightly compress the scaffold onto a balloon on a catheter system. The scaffold is deployed at the treatment site by inflation of the balloon. The balloon will induce areas of plastic stress in the bioabsorbable material to cause the scaffold to achieve and maintain the appropriate diameter on deployment.
A stent scaffold can include a plurality of cylindrical rings connected or coupled with linking elements. For example, the rings may have an undulating sinusoidal structure. When deployed in a section of a vessel, the cylindrical rings are load bearing and support the vessel wall at an expanded diameter or a diameter range due to cyclical forces in the vessel. Load bearing refers to the supporting of the load imposed by radial inwardly directed forces. Structural elements, such as the linking elements or struts, are generally non-load bearing, serving to maintain connectivity between the rings. For example, a stent may include a scaffold composed of a pattern or network of interconnecting structural elements or struts.
The structural pattern in
The treatment methods disclosed herein can apply to bioresorbable scaffolds for both coronary and peripheral treatment. Bioresorbable polymer scaffolds for coronary artery treatment can have a length between 12 to 18 mm. Such coronary scaffolds may be laser cut from polymer tubes with a diameter between 2.5 mm to 4.5 mm and with a thickness/width of 140-160 microns.
The coronary scaffold may be configured for being deployed by a non-compliant or semi-compliant balloon from about a 1.1 to 1.5 mm diameter (e.g., 1.35 mm) crimped profile. Exemplary balloon sizes include 2.5 mm, 3.0 mm, 3.5 mm, and 4.0 mm, where the balloon size refers to a nominal inflated diameter of the balloon. The scaffold may be deployed to a diameter of between 2.5 mm and 5 mm, 2.5 to 4.5 mm, or any value between and including the endpoints. The pressure of the balloon to deploy the scaffold may be 12 to 20 psi. Embodiments of the invention include the scaffold in a crimped diameter over and in contact with a deflated catheter balloon.
The intended deployment diameter may correspond to, but is not limited to, the nominal deployment diameter of a catheter balloon which is configured to expand the scaffold. The balloon pressure and the diameter to which the balloon inflates and expands the scaffold may vary from deployment to deployment. For example, the balloon may expand the scaffold in a range between the nominal inflated diameter to the nominal inflated diameter plus 0.5 mm, e.g., a 3.0 mm balloon may expand a scaffold between 3 and 3.5 mm. In any case, the inflated diameter at deployment is less than the rated burst diameter of the balloon.
A scaffold may be laser cut from a tube (i.e., a pre-cut tube) that is less than an intended deployment diameter. In this case, the pre-cut tube diameter may be 0.7 to 1 times the intended deployment diameter or any value or range in between and including the endpoints.
Compared with bare metal stents, drug-eluting stents (DES) that are not bioresorbable have been shown to be safe and to result in greater absolute reductions in target lesion revascularization (TLR) and target vessel revascularization. A DES refers to a stent including a support structure (e.g., scaffold) and also includes a drug eluting coating over the support structure. The coating can include a polymer and a drug. The polymer functions as a drug reservoir for delivery of the drug to a vessel. The polymer can be non-biodegradable or bioresorbable.—The DES that are not bioresorbable include a metal support structure with a drug eluting coating
The ABSORB Bioresorbable everolimus eluting vascular scaffold (ABSORB BVS) of Abbott Vascular Inc. of Santa Clara, Calif. was recently developed to provide an approach to treating coronary artery lesions with transient vessel support and drug delivery. Preclinical evaluation in an animal model demonstrated substantial polymer degradation at 2-years post ABSORB BVS implantation, with complete disappearance of the BVS strut “footprint” in the vessel wall within a 3-4 year period. The first generation BVS (BVS revision 1.0) was tested in the ABSORB cohort A trial and demonstrated promising results with a low event clinical rate at up to 4 years follow up (EuroIntervention 2012; 7:1060-1061). The device was however limited by a slightly higher acute recoil compared to conventional metallic platform stents. The ABSORB Cohort A 5 year follow-up clinical results are shown in Table 1 below.
Improvements in design were therefore introduced in the second generation BVS (BVS revision 1.1), notably an enhanced mechanical strength, more durable support to the vessel wall, a reduced maximum circular unsupported surface area and a more uniform strut distribution and drug delivery. The performance of the next generation BVS revision 1.1 was subsequently investigated in the ABSORB Cohort B Trial which reported excellent clinical results at 1 and 2 year follow-up (J Am Coll Cardiol. 2011; 58: B66).
The polymer backbone is made of poly(L-lactide). The diameter of the scaffold is 3 mm and the length is 18 mm. The struts have a width of about 165 microns and thickness of about 152 microns. The coating is a mixture of poly(DL-lactide) and everolimus with a 1:1 ratio of polymer to drug. The coating is about 2 to 2.5 microns in thickness. The drug dose density is 100 μg/cm2, which is the drug mass per scaffold surface area. The surface area of the scaffold is 160 mm2, so the target drug dose is about 160 μg. The surface area of the scaffold per unit scaffold length is about 8.9 mm2/mm.
In a preferred embodiment a scaffold for coronary applications has the stent pattern described in U.S. application Ser. No. 12/447,758 (US 2010/0004735) to Yang & Jow, et al. Other examples of stent patterns suitable for PLLA are found in US 2008/0275537.
The stent pattern 700 includes various struts 702 oriented in different directions and gaps 703 between the struts. Each gap 703 and the struts 702 immediately surrounding the gap 703 define a closed cell 704. At the proximal and distal ends of the stent, a strut 706 includes depressions, blind holes, or through holes adapted to hold a radiopaque marker that allows the position of the stent inside of a patient to be determined.
One of the cells 704 is shown with cross-hatch lines to illustrate the shape and size of the cells. In the illustrated embodiment, all the cells 704 have the same size and shape. In other embodiments, the cells 704 may vary in shape and size.
Still referring to
Edge Effects
The vascular response of the segments adjacent to the proximal and distal edges of the ABSORB Everolimus-Eluting Bioresorbable Vascular Scaffold were investigated at 6 Months and 1 year follow-up. JACC Cardiovasc Interv. 2012 June; 5(6):656-65. Results are based on a virtual histology intravascular ultrasound study.
The study sought to investigate in vivo the vascular response at the proximal and distal edges of the ABSORB everolimus-eluting bioresorbable vascular scaffold (BVS). The edge vascular response after implantation of the BVS has not been previously investigated.
The adjacent (5-mm) proximal and distal vessel segments to the implanted ABSORB BVS were investigated at either 6 months (B1) or 1 year (B2) with virtual histology intravascular ultrasound (VH-IVUS) imaging. At the 5-mm proximal edge, the only significant change was modest constrictive remodeling at 6 months. The constrictive remodeling is demonstrated by a decrease in the vessel cross sectional area.
The change in vessel cross-sectional area at 6 months from deployment is −1.80% [−3.18; 1.30], p<0.05). There was a tendency for the constrictive remodeling to regress or decrease after 6 months, since at 1 year the change vessel cross-sectional area since deployment is −1.53% [−7.74; 2.48], p=0.06).
The relative change of the fibrotic and fibrofatty (FF) tissue areas at the proximal segment were not statistically significant at either time point. At the 5-mm distal edge, a significant increase in the FF tissue of 43.32% [−19.90; 244.28], (p<0.05) 1-year post-implantation was evident. The increase may be at least 40%. The changes in dense calcium need to be interpreted with caution since the polymeric struts are detected as “pseudo” dense calcium structures with the VH-IVUS imaging modality.
The vascular response up to 1 year after implantation of the ABSORB BVS demonstrated some degree of proximal edge constrictive remodeling that tends to regress at 1 year. Some degree of proximal edge and distal edge plaque compositional changes were observed with increase of the fibrofatty tissue component at 1-year. The distal edge increases in fibro-fatty tissue resulting in nonsignificant plaque progression with adaptive expansive remodeling. This morphological and tissue composition behavior appears to not significantly differ from the behavior of metallic drug-eluting stents at the same observational time points. The constrictive remodeling at the proximal edge tends to regress at 1-year. This biological behavior is similar to that observed with the metallic devices at the same follow-up points.
Tables 4A and 4B provide the proximal edge vascular response in terms of percent change in vessel cross-sectional area (CSA), lumen CSA, and plaque CSA. Tables 5A and 5B provide the distal edge vascular response.
.90]
.80; 8.65]
.0
9]
.60; 8.41]
6; 7.25]
; 8.30]
; 0.21]
indicates data missing or illegible when filed
]
indicates data missing or illegible when filed
Table 6 shows ABSORB Cohort B trial results up to 2 years follow-up. Table 6 shows no scaffold thrombosis out to 2 years and only 2 additional TLR events between 1 year and 2 years, and MACE rate of 8.9% (3 non-Q wave MI, 6 ID TLR) at 2 years which is comparable to Xience V.
Compliance
Vascular compliance changes in the coronary vessel wall after bioresorbable vascular implantation in the treated and adjacent segments. Implantation of a metallic prosthesis creates local stiffness with a subsequent mismatch in compliance between the scaffolded and the immediate adjacent segments.
A total of 83 patients from the ABSORB trials underwent palpography investigations (30 and 53 patients from ABSORB Cohort A and B, respectively) to measure the compliance of the scaffolded and adjacent segments at various time points (from pre-implantation up to 24 months). The mean of the maximum strain values in all cross sections was calculated per segment by utilizing the Rotterdam Classification (ROC) score and expressed as ROC/mm.
After scaffold implantation mismatch in compliance was evident in patients with paired analyses between the scaffolded and adjacent segments (proximal: 0.23[0.12-0.34], scaffold: 0.12[0.07-0.19], distal: 0.15[0.05-0.26], p=0.042). Thus, mismatch is greater between the scaffolded segment and the proximal segment and the scaffolded segment and the distal segment. The former may be at least 90% or 90 to 100% and the latter may be at least 10% or 10 to 40%. This reported compliance mismatch disappeared at short and mid-term follow-up (6 and 12 months).
The conclusions of the results are that the ABSORB scaffold decreases vascular compliance at the site of scaffold implantation. A compliance mismatch is present immediately post-implantation and in contrast to metallic stents disappears in the mid-term likely leading to a normalization of the rheological behavior of the scaffolded and adjacent segments. The Cohort A and B scaffolds have also been shown to exhibit low late loss and exhibit low restenosis. The BVS scaffolds provide these favorable clinical outcomes in spite of the thicker/wider struts of these scaffolds (approx. 150 microns) compared to metal stents, e.g., Xience V and Taxus Express.
It is believed that favorable clinical outcomes thus far for patients are due to synergy between various unique aspects of the BVS scaffolds:
In general, the treatment with bioabsorbable polymer stents has a number of advantages over permanent implants: (i) The stent disappears from the treated site resulting in reduction or elimination of late stent thrombosis; (ii) disappearance of the stent facilitates repeat treatments (surgical or percutaneous) to the same site; (iii) disappearance of the stent allows restoration of vasomotion at the treatment site (the presence of a rigid permanent metal stent restricts vasomotion); (iv) the bioabsorbability results in freedom from side-branch obstruction by struts; (v) the disappearance results in freedom from strut fracture and ensuing restenosis. Some of these advantages may be relevant to improving clinical outcomes for non-diabetic and diabetic patients.
In the short term and over the long term, a bioresorbable scaffold has the advantage of being less traumatic to the vessel wall. Since the bioresorbable scaffold degrades with time and eventually disappears, trauma associated with the presence of a scaffold decreases with time and eventually disappears. Resorption of a bioresorbable scaffold which restores vasomotion of the vessel wall may reduce long term thrombotic risk.
The thrombogenic potential has been evaluated based on platelet adhesion to the BVS cohort B scaffold deployed ex vivo. Platelets are indispensable initiators of thrombosis and their adhesion to intravascular devices is the critical step in the thrombus formation. In a study of platelet adhesion, metallic coronary stents (BMS Multilink Vision and Xience V) and BVS scaffolds were deployed in a Chandler Loop perfused with freshly prepare porcine platelet rich plasma (PRP) instead of whole blood. The extent of platelet adhesion is determined by measuring the LDH activity extracted from the adherent platelets which is directly proportional to the number of platelets. Such properties may be of particular benefit in diabetic vascular disease. Thrombogenicity based on the adhesion of platelets was consistently the highest for the BMS Multi-Link Vision followed by the Xience V stents and followed by the BVS scaffolds.
Thicker scaffold struts with a higher total dose of drug may be beneficial in reducing incidence of smooth muscle cell proliferation. The thicker struts in the BVS scaffold, about 150 to 165 microns, results in a total dose of everolimus that is almost two fold higher than XIENCE V.
One aspect is the use of a polymer, in particular a bioresorbable polymer, for the scaffold. A polymer scaffold may be less traumatic to a vasculature. Polymers are softer, less stiff or have a lower modulus than metals. Thus, the presence of a softer, more flexible implant may be less traumatic to a soft, flexible vessel segment than a metal implant. For example, aliphatic bioresorbable polymers have tensile moduli generally less than 7 GPa and in the range of 2 to 7 GPa (US2009/0182415). Poly(L-lactide) has a tensile modulus of about 3 GPa.
Metals used to make a stent and their approximate moduli include stainless steel 316L (143 GPa), tantalum (186 GPa), Nitinol or nickel-titanium alloy (83 Gpa), and cobalt chromium alloys (243 Gpa). These moduli are significantly higher than aliphatic polymers. The strengths of these metals are also significantly higher than the polymers as well. As a result, a bioresorbable polymeric scaffold has thicker struts to help compensate for the difference in the material properties to provide a radial stiffness and radial strength this sufficient to provide patency.
Also, the mismatch of the properties of a polymer scaffold and a vessel segment is lower than for a metallic scaffold. This mismatch can be expressed formally in terms of compliance mismatch between the scaffold and the vessel segment at the implant site. The compliance of a material, which is the inverse of stiffness or modulus of a material, refers to the strain of an elastic body expressed as a function of the force producing the strain. The compliance of a scaffold or radial compliance of the scaffold can likewise be defined as the inverse of the radial stiffness of the scaffold. The radial stiffness of the bioresorbable scaffold is lower than a metallic scaffold, so the radial compliance of the bioresorbable scaffold is higher than a metallic scaffold. The compliance mismatch of a polymer scaffold is lower than a metallic stent.
The compliance of a stent, both nondegradable and resorbable, is necessarily much lower than the vessel segment in order for the scaffold to support the vessel at a deployed diameter with minimal periodic recoil due to inward radial forces from the vessel walls. Additionally, it results in better conformity (and less straightening) of the scaffolded segment to the overall curvature of the adjacent segments in the treated vessel. However, an additional aspect of a bioresorbable polymer scaffold that may contribute to favorable clinical outcomes is that the compliance mismatch decreases with time due to the degradation of the bioresorbable polymer. As the polymer of the scaffold degrades, mechanical properties of the polymer such as strength and stiffness decrease and compliance increases. As a result, the radial strength of the scaffold decreases with time and the compliance of the scaffold increases with time since these properties depend on the properties of the scaffold material.
In the long term, the compliance of a vessel segment with an implanted scaffold converges to that of the natural compliance of the vessel. The convergence of the compliance occurs gradually as the vessel segment heals. Since natural compliance of a vessel segment is eventually restored due to complete resorption of the scaffold, natural vasomotion of the vessel segment is also restored. Compliance mismatch in the treatment with metallic stents is permanent and has been identified as a contributor to the process of restenosis and potentially late adverse events.
Another aspect that may contribute to favorable clinical outcomes of bioresorbable scaffolds is a higher drug loading or target dose of the bioresorbable scaffold. From above, the BVS scaffold in the ABSORB Cohort A and B trials is 18 mm long and has a drug dose density of 100 μg/cm2 and a target drug dose of about 160 μg. The target drug dose per unit scaffold length of the ABSORB Cohort B trial scaffold is about 8.9 μg/mm. The delivery of the target dose to the vessel can occur over a period of about 2 to 3 months after implantation.
The drug dose density of the XIENCE V® stent (http://www.accessdata.fda.gov/cdrh_docs/pdf11/P110019b.pdf) and TAXUS Express® (American Heart Journal Volume 163, Number 2, p. 143-148) are both reported to be 100 μg/cm2. However, the BVS target dose and dose per unit length is larger due to the wider and thicker struts compared to these stents: XIENCE V® (91 mm×81 mm) and TAXUS Express® (91 mm×132 mm).
BMS and metallic DES stents typically have strut widths and thicknesses much less than the BVS stent (Interventional Cardiology, Vol. 6, Issue 2, pp. 143-147). The larger strut width and strut thickness, or equivalently, larger surface area of the BVS scaffold may also contribute to favorable clinical outcomes of diabetic patients. The larger strut width and strut thickness or surface area of a bioresorbable scaffold contributes by providing a higher target dose due to the higher surface area of contact with the vessel walls.
The 3 year results of the ABSORB Cohort B Trial are further provided summarized in
The clinical results for Cohort B Groups 1 and 2 are shown in Table 7. There is no scaffold thrombosis by ARC or protocol.
The results include serial image acquisition at baseline, 1 year, and 3 years including events: OCT Optional 19 patients, IVUS-GS Mandatory 45 patients, IVUS-VH Mandatory 38 patients, IVUS-Echogenicity, derived from GS 29 patients, and angiography mandatory 51 patients.
In the following months (from 6 to 12 months) it has been shown that physiological and pharmacological vasomotion reappears confirming the fact that the mechanical stiffness of the polymer is progressively replaced by de novo formation of malleable tissue such as proteoglycan.
At 2 years it has been demonstrated that the scaffold device despite its malleable and deformable structure did not undergo any reduction in area or volume. In contrast, a late enlargement of the scaffold was documented, probably due to the intraluminal expansive force of the systolic/diastolic wall stress. This late enlargement of the scaffold compensates for the intraluminal growth of neointimal tissue.
The ultimate expectation of the bioabsorbable stent intervention is the occurrence of late lumen enlargement, associated with wall thinning, without expansive remodeling.
At 3-year follow-up of the Cohort B showed: stable late loss, return of vasomotion to the scaffolded segment, enlargement of scaffold area as well as mean lumen area despite persisting increase of neointima, reduction of plaque area, and bioresorption slower than the first generation of ABSORB (1.0).
3 year follow-up results show improvements in blood vessel movement, area inside the vessel, and reduction of plaque where the scaffold was placed.
At three years, the rate of major adverse cardiovascular events (MACE) in 101 patients was 10 percent, similar to a comparative set of data with a best-in-class drug eluting stent at three years. MACE is a combined endpoint that includes heart attacks, deaths for heart related causes or re-blockages of the blood vessel resulting in symptoms requiring the need for additional procedures at the original site of scaffold implantation.
In a subset of 46 patients, pictures inside the blood vessel using state-of-the art imaging techniques showed improvements in vessel motion and an average increase of 7.3 percent between one and three years in the area within the blood vessel, allowing more blood to flow through the vessel as the body requires, a finding unique to Absorb and not typically observed with metallic stents that cage the vessel. There was also a decrease of plaque inside the vessel between two and three years. Plaque is made up of fat, cholesterol, calcium and other deposits that accumulate on the inner wall of the artery in patients with coronary heart disease and can slow or stop blood flow to the heart.
The clinical data up to 3 years a showed an ID-MACE rate of 10.0% with no events of scaffold thrombosis. The late loss at 3 years was 0.32±042 mm. The IVUS grey scale results revealed scaffold and lumen enlargement between baseline and 3 years (6.29±0.91 vs. 7.08±1.55, p<0.0001 and 6.29±0.90 vs. 6.81±1.62, p=0.0155, respectively). The scaffold enlargement was confirmed by OCT (7.76±1.07 at baseline vs. 8.64±2.15 at 3 years, p=0.0446).
The IVUS-VH and the IVUS-derived echogenicity results show signs of bioresorption indicated by a significant reduction in dense calcium and in percent hyper-echogenic area, respectively, between baseline and 3 years.
Table 8 shows the VH results of dense calcium area percent at baseline, 1 year, and 3 years follow-up. The dense calcium area percent decreases between baseline and 1 year and between 1 year and 3 years.
*
The ABSORB EXTEND study is a single-arm trial evaluating Absorb in patients with more complex heart disease. Data from 450 patients enrolled in this trial showed that the rates of MACE at one year were slightly lower than a best-in-class DES. In an analysis of 119 patients with diabetes from the EXTEND trial, rates of MACE were the same in patients with and without diabetes, a promising finding as event rates are typically higher in patients with diabetes when compared to patients without diabetes.
The prevailing mechanism of degradation of many bioabsorbable polymers is chemical hydrolysis of the hydrolytically unstable backbone. In a bulk degrading polymer, the polymer is chemically degraded throughout the entire polymer volume. As the polymer degrades, the molecular weight decreases. The reduction in molecular weight results in changes in mechanical properties (e.g., strength) and stent properties. For example, the strength of the scaffold material and the radial strength of the scaffold are maintained for a period of time followed by a gradual or abrupt decrease. The decrease in radial strength is followed by a loss of mechanical integrity and then erosion or mass loss. Mechanical integrity loss is demonstrated by cracking and by fragmentation. Enzymatic attack and metabolization of the fragments occurs, resulting in a rapid loss of polymer mass.
The behavior of a bioabsorbable stent upon implantation can divided into three stages of behavior. In stage I, the stent provides mechanical support. The radial strength is maintained during this phase. Also during this time, chemical degradation occurs which decreases the molecular weight. In stage II, the scaffold experiences a loss in strength and mechanical integrity. In stage III, significant mass loss occurs after hydrolytic chain scission yields water-soluble low molecular weight species.
The scaffold in the first stage provides the clinical need of providing mechanical support to maintain patency or keep a vessel open at or near the deployment diameter. In some treatments, the patency provided by the scaffold allows the stented segment of the vessel to undergo positive remodeling at the increased deployed diameter. Remodeling refers generally to structural changes in the vessel wall that enhances its load-bearing ability so that the vessel wall in the stented section can maintain an increased diameter in the absence of the stent support. A period of patency is required in order to obtain permanent positive remodeling.
The manufacturing process of a bioabsorbable scaffold includes selection of a bioabsorbable polymer raw material or resin. Detailed discussion of the manufacturing process of a bioabsorbable stent can be found elsewhere, e.g., U.S. Patent Publication No. 20070283552. The fabrication methods of a bioabsorbable stent can include the following steps:
(1) forming a polymeric tube from a biodegradable polymer resin using extrusion,
(2) radially deforming the formed tube to increase radial strength,
(3) forming a stent scaffolding from the deformed tube by laser machining a stent pattern in the deformed tube with laser cutting, in exemplary embodiments, the strut thickness can be 100-200 microns, or more narrowly, 120-180, 130-170, or 140-160 microns,
(4) optionally forming a therapeutic coating over the scaffolding,
(5) crimping the stent over a delivery balloon, and
(6) sterilization with election-beam (E-beam) radiation.
Poly(L-lactide) (PLLA) is attractive as a stent material due to its relatively high strength and rigidity at human body temperature, about 37° C. Since it has a glass transition temperature between about 60 and 65° C. (Medical Plastics and Biomaterials Magazine, March 1998), it remains stiff and rigid at human body temperature. This property facilitates the ability of a PLLA stent scaffold to maintain a lumen at or near a deployed diameter without significant recoil (e.g., less than 10%). In general, the Tg of a semicrystalline polymer can depend on its morphology, and thus how it has been processed. Therefore, Tg refers to the Tg at its relevant state, e.g., Tg of a PLLA resin, extruded tube, expanded tube, and scaffold.
In general, a scaffold can be made of a bioresorbable aliphatic polyester. Additional exemplary biodegradable polymers for use with a bioabsorbable polymer scaffolding include poly(D-lactide) (PDLA), polymandelide (PM), polyglycolide (PGA), poly(L-lactide-co-D,L-lactide) (PLDLA), poly(D,L-lactide) (PDLLA), poly(D,L-lactide-co-glycolide) (PLGA) and poly(L-lactide-co-glycolide) (PLLGA). With respect to PLLGA, the stent scaffolding can be made from PLLGA with a mole % of GA between 5-15 mol %. The PLLGA can have a mole % of (LA:GA) of 85:15 (or a range of 82:18 to 88:12), 95:5 (or a range of 93:7 to 97:3), or commercially available PLLGA products identified as being 85:15 or 95:5 PLLGA. The examples provided above are not the only polymers that may be used. Many other examples can be provided, such as those found in Polymeric Biomaterials, second edition, edited by Severian Dumitriu; chapter 4.
Polymers that are more flexible or that have a lower modulus than those mentioned above may also be used. Exemplary lower modulus bioabsorbable polymers include, polycaprolactone (PCL), poly(trimethylene carbonate) (PTMC), polydioxanone (PDO), poly(4-hydroxy butyrate) (PHB), and poly(butylene succinate) (PBS), and blends and copolymers thereof.
In exemplary embodiments, higher modulus polymers such as PLLA or PLLGA may be blended with lower modulus polymers or copolymers with PLLA or PLGA. The blended lower modulus polymers result in a blend that has a higher fracture toughness than the high modulus polymer. Exemplary low modulus copolymers include poly(L-lactide)-b-polycaprolactone (PLLA-b-PCL) or poly(L-lactide)-co-polycaprolactone (PLLA-co-PCL). The composition of the blend can include 1-5 wt % of low modulus polymer.
The BVS scaffolds are coated with a polymer mixture that includes Everolimus, an antiproliferative agent. In general, the anti-proliferative agent can be a natural proteineous agent such as a cytotoxin or a synthetic molecule or other substances such as actinomycin D, or derivatives and analogs thereof (manufactured by Sigma-Aldrich 1001 West Saint Paul Avenue, Milwaukee, Wis. 53233; or COSMEGEN available from Merck) (synonyms of actinomycin D include dactinomycin, actinomycin IV, actinomycin actinomycin X1, and actinomycin C1), all taxoids such as taxols, docetaxel, and paclitaxel, paclitaxel derivatives, all olimus drugs such as macrolide antibiotics, rapamycin, everolimus, structural derivatives and functional analogues of rapamycin, structural derivatives and functional analogues of everolimus, FKBP-12 mediated mTOR inhibitors, biolimus, perfenidone, prodrugs thereof, co-drugs thereof, and combinations thereof. Representative rapamycin derivatives include 40-O-(3-hydroxy)propyl-rapamycin, 40-O-[2-(2-hydroxy)ethoxy]ethyl-rapamycin, or 40-O-tetrazole-rapamycin, 40-epi-(N-1-tetrazolyl)-rapamycin (ABT-578 manufactured by Abbott Laboratories, Abbott Park, Ill.), prodrugs thereof, co-drugs thereof, and combinations thereof.
An anti-inflammatory agent can be a steroidal anti-inflammatory agent, a nonsteroidal anti-inflammatory agent, or a combination thereof. In some embodiments, anti-inflammatory drugs include, but are not limited to, alclofenac, alclometasone dipropionate, algestone acetonide, alpha amylase, amcinafal, amcinafide, amfenac sodium, amiprilose hydrochloride, anakinra, anirolac, anitrazafen, apazone, balsalazide disodium, bendazac, benoxaprofen, benzydamine hydrochloride, bromelains, broperamole, budesonide, carprofen, cicloprofen, cintazone, cliprofen, clobetasol propionate, clobetasone butyrate, clopirac, cloticasone propionate, cormethasone acetate, cortodoxone, deflazacort, desonide, desoximetasone, dexamethasone dipropionate, diclofenac potassium, diclofenac sodium, diflorasone diacetate, diflumidone sodium, diflunisal, difluprednate, diftalone, dimethyl sulfoxide, drocinonide, endrysone, enlimomab, enolicam sodium, epirizole, etodolac, etofenamate, felbinac, fenamole, fenbufen, fenclofenac, fenclorac, fendosal, fenpipalone, fentiazac, flazalone, fluazacort, flufenamic acid, flumizole, flunisolide acetate, flunixin, flunixin meglumine, fluocortin butyl, fluorometholone acetate, fluquazone, flurbiprofen, fluretofen, fluticasone propionate, furaprofen, furobufen, halcinonide, halobetasol propionate, halopredone acetate, ibufenac, ibuprofen, ibuprofen aluminum, ibuprofen piconol, ilonidap, indomethacin, indomethacin sodium, indoprofen, indoxole, intrazole, isoflupredone acetate, isoxepac, isoxicam, ketoprofen, lofemizole hydrochloride, lomoxicam, loteprednol etabonate, meclofenamate sodium, meclofenamic acid, meclorisone dibutyrate, mefenamic acid, mesalamine, meseclazone, methylprednisolone suleptanate, morniflumate, nabumetone, naproxen, naproxen sodium, naproxol, nimazone, olsalazine sodium, orgotein, orpanoxin, oxaprozin, oxyphenbutazone, paranyline hydrochloride, pentosan polysulfate sodium, phenbutazone sodium glycerate, pirfenidone, piroxicam, piroxicam cinnamate, piroxicam olamine, pirprofen, prednazate, prifelone, prodolic acid, proquazone, proxazole, proxazole citrate, rimexolone, romazarit, salcolex, salnacedin, salsalate, sanguinarium chloride, seclazone, sermetacin, sudoxicam, sulindac, suprofen, talmetacin, talniflumate, talosalate, tebufelone, tenidap, tenidap sodium, tenoxicam, tesicam, tesimide, tetrydamine, tiopinac, tixocortol pivalate, tolmetin, tolmetin sodium, triclonide, triflumidate, zidometacin, zomepirac sodium, aspirin (acetylsalicylic acid), salicylic acid, corticosteroids, glucocorticoids, tacrolimus, pimecorlimus, prodrugs thereof, co-drugs thereof, and combinations thereof.
These agents can also have anti-proliferative and/or anti-inflammatory properties or can have other properties such as antineoplastic, antiplatelet, anti-coagulant, anti-fibrin, antithrombonic, antimitotic, antibiotic, antiallergic, antioxidant as well as cystostatic agents. Examples of suitable therapeutic and prophylactic agents include synthetic inorganic and organic compounds, proteins and peptides, polysaccharides and other sugars, lipids, and DNA and RNA nucleic acid sequences having therapeutic, prophylactic or diagnostic activities. Nucleic acid sequences include genes, antisense molecules which bind to complementary DNA to inhibit transcription, and ribozymes. Some other examples of other bioactive agents include antibodies, receptor ligands, enzymes, adhesion peptides, blood clotting factors, inhibitors or clot dissolving agents such as streptokinase and tissue plasminogen activator, antigens for immunization, hormones and growth factors, oligonucleotides such as antisense oligonucleotides and ribozymes and retroviral vectors for use in gene therapy. Examples of antineoplastics and/or antimitotics include methotrexate, azathioprine, vincristine, vinblastine, fluorouracil, doxorubicin hydrochloride (e.g. Adriamycin® from Pharmacia & Upjohn, Peapack N.J.), and mitomycin (e.g. Mutamycin® from Bristol-Myers Squibb Co., Stamford, Conn.). Examples of such antiplatelets, anticoagulants, antifibrin, and antithrombins include sodium heparin, low molecular weight heparins, heparinoids, hirudin, argatroban, forskolin, vapiprost, prostacyclin and prostacyclin analogues, dextran, D-phe-pro-arg-chloromethylketone (synthetic antithrombin), dipyridamole, glycoprotein IIb/IIIa platelet membrane receptor antagonist antibody, recombinant hirudin, thrombin inhibitors such as Angiomax ä (Biogen, Inc., Cambridge, Mass.), calcium channel blockers (such as nifedipine), colchicine, fibroblast growth factor (FGF) antagonists, fish oil (omega 3-fatty acid), histamine antagonists, lovastatin (an inhibitor of HMG-CoA reductase, a cholesterol lowering drug, brand name Mevacor® from Merck & Co., Inc., Whitehouse Station, N.J.), monoclonal antibodies (such as those specific for Platelet-Derived Growth Factor (PDGF) receptors), nitroprusside, phosphodiesterase inhibitors, prostaglandin inhibitors, suramin, serotonin blockers, steroids, thioprotease inhibitors, triazolopyrimidine (a PDGF antagonist), nitric oxide or nitric oxide donors, super oxide dismutases, super oxide dismutase mimetic, 4-amino-2,2,6,6-tetramethylpiperidine-1-oxyl (4-amino-TEMPO), estradiol, anticancer agents, dietary supplements such as various vitamins, and a combination thereof. Examples of such cytostatic substance include angiopeptin, angiotensin converting enzyme inhibitors such as captopril (e.g. Capoten® and Capozide® from Bristol-Myers Squibb Co., Stamford, Conn.), cilazapril or lisinopril (e.g. Prinivil® and Prinzide® from Merck & Co., Inc., Whitehouse Station, N.J.). An example of an antiallergic agent is permirolast potassium. Other therapeutic substances or agents which may be appropriate include alpha-interferon, and genetically engineered epithelial cells. The foregoing substances are listed by way of example and are not meant to be limiting. Other active agents which are currently available or that may be developed in the future are equally applicable. The scaffold can exclude any of the drugs disclosed herein.
“Baseline” refers to a time immediately after deployment of a scaffold to a target diameter in a vessel or at a time after deployment long enough to make measurements on the newly deployed scaffold.
The “glass transition temperature,” 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, the Tg corresponds to the temperature where the onset of segmental motion in the chains of the polymer occurs. When an amorphous or semi-crystalline polymer is exposed to an increasing temperature, the coefficient of expansion and the heat capacity of the polymer both increase as the temperature is raised, indicating increased molecular motion. As the temperature is increased, the heat capacity increases. The increasing heat capacity corresponds to an increase in heat dissipation through movement. Tg of a given polymer can be dependent on the heating rate and can be influenced by the thermal history of the polymer as well as its degree of crystallinity. Furthermore, the chemical structure of the polymer heavily influences the glass transition by affecting mobility.
The Tg can be determined as the approximate midpoint of a temperature range over which the glass transition takes place. [ASTM D883-90]. The most frequently used definition of Tg uses the energy release on heating in differential scanning calorimetry (DSC). As used herein, the Tg refers to a glass transition temperature as measured by differential scanning calorimetry (DSC) at a 20° C./min heating rate.
“Stress” refers to force per unit area, as in the force acting through a small area within a plane. 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 applied that leads to expansion (increase in length). In addition, compressive stress is a normal component of stress applied to materials resulting in their compaction (decrease in length). Stress may result in deformation of a material, which refers to a change in length. “Expansion” or “compression” may be defined as the increase or decrease in length of a sample of material when the sample is subjected to stress.
“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.
“Strength” refers to the maximum stress along an axis which a material will withstand prior to fracture. The ultimate strength is calculated from the maximum load applied during the test divided by the original cross-sectional area.
“Modulus” may be defined as the ratio of a component of stress or force per unit area applied to a material divided by the strain along an axis of applied force that results from the applied force. The modulus typically is the initial slope of a stress—strain curve at low strain in the linear region.
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.
This application claims the benefit of U.S. Patent Application No. 61/615,185 filed Mar. 23, 2012, U.S. Patent Application No. 61/768,394 filed Feb. 22, 2013, and U.S. Patent Application No. 61/775,424 filed Mar. 8, 2013, all of which are incorporated by reference herein.
| Number | Date | Country | |
|---|---|---|---|
| 61775424 | Mar 2013 | US | |
| 61768394 | Feb 2013 | US | |
| 61615185 | Mar 2012 | US |
