Patent Application
20120280432
The technical field relates to a method for manufacturing a bioabsorbable stent, which is to be implanted in the blood vessel.
Various medical situations require the use of an endoprosthesis to support a constricted vessel and maintain an open passageway through the vessel. Artery diseases such as atherosclerosis or myocardial infarction are usually treated with the percutaneous transluminal angioplasty (PTA) through the use of a balloon catheter. It involves passing a small, balloon-tipped catheter percutaneously into a vessel and up to the region of obstruction. The balloon is then inflated to dilate the area of obstruction. However, restenosis or reclosure of the dilated vessel are usually occurred due to the thrombosis following the angioplasty. A stent is a radially expandable endoprosthesis, which is adapted to be implanted in the bodily lumen, are used to prevent these situations.
Stents restrict restenosis or reclosure of blood vessels following the angioplasty by scaffolding intimal tissue flaps that have separated from deeper arterial layers, controlling early elastic recoil, optimizing vessel caliber, and preventing subsequent constrictive remodeling that were major limitations of angioplasty. Stents have also been implanted in urinary tracts and bile ducts and other bodily lumen. Delivery and implantation of a stent is accomplished by disposing the stent about a distal portion of the catheter, percutaneously inserting the distal portion of the catheter in a vessel, advancing the catheter in the lumen to a desired location, expanding the stent and removing the catheter from the lumen. In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter and expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn. In the case of a self-expanding stent, the stent may be held in place on the catheter via a retractable sheath. When the stent is in a desired bodily location, the sheath may be withdrawn allowing the stent to self-expand. The stent stays in the artery permanently, holds it open, and improves blood flow to the organ. Within a few weeks of the time the stent was placed, the inside lining of the artery (the endothelium) grows over the surface of the stent.
The development of bare metal stents (BMS) has been a major advance in the treatment of obstructive artery disease since the introduction of PTA. BMS is a mesh-like tube of thin wire usually made of 316L stainless steel or cobalt chromium alloy. However, metal is hydrophilic and tends to form the thrombus. Neointimal hyperplasia occurring within the BMS leading to in-stent restenosis is a main obstacle in the long-term success of PTA. The recent development of drug-eluting stents (DES) further contributes a major breakthrough to the interventional cardiology. DES is a metal stent placed into obstructive blood vessels that slowly releases a drug to block cell proliferation. The neointimal growth response to stenting that contributes to restenosis can be largely abolished by coating stents with antiproliferative drugs. DES has showed a remarkable reduction in in-stent restenosis and target vessel revascularization when compared with BMS. Despite the high success rate of DES, there is a low incidence of late stent thrombosis, which is probably because the antiproliferative drugs delay the growth of healthy endothelium over stent struts and their durable polymer coating. Furthermore, implanting a BMS or DES which will remain permanently in the human body, possess various long-term potential problems.
It has been a clinical consensus that the vessel scaffolding and drug delivery is only needed during the vascular healing period after stenting, and permanent scaffolding is needless after the cease of acute recoil and constrictive remodeling processes [Circulation 102 (2000) 371]. The recent introduction of the fully bioabsorbable stents that aim to fulfill these purposes will hold potential advantages. Unlike permanent BMS and DES, which are both afflicted with long-term risks, bioabsorbable stents that once dissolved will leave behind only the healed natural vessel, no need of eventual surgical removal, and late stent thrombosis will no longer be a concern. Further advantages of bioabsorbable stents over permanent stents include improved lesion imaging with computed tomography or magnetic resonance, facilitation of repeat surgical treatment to the same site, restoration of vasomotion, and freedom from side-branch obstruction by struts and from strut fracture-induced restenosis. Because bioabsorbable stents are less rigid than metal stents, they are more suitable for complex anatomy such as in superficial femoral and tibial arteries, where stent crush and fracture may occur due to flexion or extension articulations.
The development of bioabsorbable stents goes back to the mid-1980s, when pioneering work was done by Stack et al. [American J. Cardiovascular 62(1988)3F]. Since then, a number of international research groups have reported on various bioabsorbable stent designs, and some have gone through preclinical to clinical evaluation. Bioabsorbable stents are generally tubular and have been made of many materials, including bioabsorbable polymer, iron or magnesium based alloy. The polymer PLLA (poly-L-lactic acid) is used as a bioabsorbable coating of permanent metallic stents but can also be used to manufacture complete stents. The PLLA stents undergoes hydrolysis, resulting in lactic acid, and finally metabolism into carbon dioxide and water within 2-3 years after implantation. The bioabsorbable iron or magnesium based stents degrades within the body over a 2- to 3-month timeframe, forming inorganic salts containing calcium, chloride, oxide, sulfates and phosphates. The structure of bioabsorbable stents comprises pattern or network of interconnecting structural elements referred to as struts. A number of techniques have been suggested for the fabrication of stents from tubes, wires, or sheets of material rolled into a cylindrical shape.
Feasibilities of bioabsorbable stents have already been established. There are a number of requirements that must be satisfied by bioabsorbable stents, when they keep the blood vessel open for a specified period of time. Polymers tend to have lower strength than metals based on same mass basis. Therefore, polymeric stents typically have less circumferential strength and radial rigidity than metallic stents of the same or similar dimensions. Especially in the bending portions of the stent that are bent during crimping and expansion of the stent. As the structural element, the stent needs to possess sufficient radial strength to against radial compressive forces imposed on the stent. Once expanded, the stent must adequately maintain its size and shape throughout its service life, radial compressive forces tend to cause a stent to recoil inward. Generally, it is desirable to minimize recoil. Also, the stents should be sufficiently rigid to avoid stent deformity, despite the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. In addition, the stents should have sufficient toughness or resistance to fracture from stress arising from crimping, expansion, and cyclic loading. Furthermore, the stent must possess sufficient flexibility to allow for crimping, expansion, and cyclic loading. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous vascular path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure.
It has been reported that the strength and rigidity of the polymer tube can be increased by expanding the tube wall radially and/or axially so as to orient the polymer molecules of the tube.
Above mentioned bioabsorbable stents are fabricated from expanded polymer tubes, which are made by reheating and expanding polymer tubes. However, polymer tubes that made from polymer resins or pellets must be processed by casting molding, injection molding, or extrusion molding in the first step of a two-stage process. The polymer tube is allowed to cool to the room temperature and is possible stored at temperature below −20° C. for later use in the case of PLLA. Polymer tubes are subsequently reheated and stretch blown through a second step of blow molding process into expanded polymer tubes. When PLLA resins are processed in elevated temperatures, PLLA resins are known to undergo thermal degradation, which can impact the mechanical properties of the resulting stents.
The thermal degradation of PLLA, which leads to the formation of lactide monomers, is related to the process temperature and the residence time in the extruder and hot mold. The rate of molecular weight loss of PLLA at 60° C. is more than 100 times greater than that PLLA at 40° C. [Progress in Polymer Science 2008(33) 820]. By and large, thermal degradation of PLLA resin can be attributed to: (a) hydrolysis by trace amounts of water, (b) zipper-like depolymerization, (c) oxidative, random main-chain scission by oxygen in air, (d) intermolecular transesterification to monomer and oligomeric esters, and (e) intramolecular transesterification resulting in formation of monomer and oligomer lactides of low molecular weight [Progress Material Sciences 2002(27)1123]. The moisture content of PLLA resin, temperature, and residence time of PLLA resin during the thermal processes are important contributors to molecular weight loss of PLLA [Apply Polymer Sciences 2001(79)2128]. These results highlighted the importance of minimizing the residence time and process temperature during the processing of PLLA resins. Furthermore, it is sometimes difficult to reheat the polymer tube uniformly to a suitable blowing temperature using a heater, which provides radiant energy to the outside of the polymer tube. A temperature gradient can exist from the outside wall to the inside wall of the polymer tube. It is possible to overheat the outside wall of the polymer tube when reheating the polymer tube to a suitable blowing temperature, which will affect the uniformity of wall thickness and mechanical properties of expanded polymer tubes after the blow-molding process.
One embodiment of the disclosure provides a method for manufacturing a bioabsorbable stent comprising providing a polymer resin; melting the polymer resin to form a molten hollow parison; cooling the molten hollow parison to form a hot hollow parison; elongating the hot hollow parison; expanding the hot hollow parison by feeding a compressed gas into the hot hollow parison to form a stent preform; and patterning the stent preform to form a bioabsorbable stent.
One embodiment of the disclosure provides a method for manufacturing a bioabsorbable stent comprising forming a molten hollow parison of a polymer resin from an annular die-head assembly; closing around the molten hollow parison by closing two halves of an opening tubular mold; shaping and partially cooling the molten hollow parison into a hot hollow parison; opening the tubular mold; closing around the hot hollow parison by closing two halves of an opening stretch-blowing mold; axially elongating the hot hollow parison by clamping one end of the hot hollow parison with a mandrel and moving inside the stretch-blowing mold; radially expanding the hot hollow parison by feeding a compressed gas into the hot hollow parison until the hot hollow parison conforms to an inside surface of the stretch-blowing mold to form an inflated hollow parison; cooling the inflated hollow parison to an ambient temperature to form a stent preform; releasing the stent preform from the stretch-blowing mold; and fabricating the stent preform into a bioabsorbable stent by impinging a specified pattern onto the stent preform with a pulsing laser cutting device.
One embodiment of the disclosure provides a method for manufacturing a bioabsorbable stent, comprising: forming a molten hollow parison of a programmed wall thickness of a polymer resin from an annular die-head assembly; closing around the molten hollow parison by closing two halves of an opening tubular mold; shaping and partially cooling the molten hollow parison of the programmed wall thickness into a hot hollow parison of a programmed wall thickness; opening the tubular mold; closing around the hot hollow parison of the programmed wall thickness by closing two halves of an opening stretch-blowing mold; axially elongating the hot hollow parison of the programmed wall thickness by clamping one end of the hot hollow parison of the programmed wall thickness with a mandrel and moving inside the stretch-blowing mold; radially expanding the hot hollow parison of the programmed wall thickness by feeding a compressed gas into the hot hollow parison of the programmed wall thickness until the hot hollow parison conforms to an inside surface of the stretch-blowing mold to form an inflated hollow parison of a programmed wall thickness; cooling the inflated hollow parison of the programmed wall thickness to an ambient temperature to form a stent preform of a programmed wall thickness; releasing the stent preform of the programmed wall thickness from the stretch-blowing mold; and fabricating the stent preform into a bioabsorbable stent by impinging a specified pattern onto the stent preform of the programmed wall thickness with a pulsing laser cutting device.
A detailed description is given in the following embodiments with reference to the accompanying drawings.
The disclosure can be more fully understood by reading the subsequent detailed description and examples with references made to the accompanying drawing, wherein:
In the following detailed description, for purposes of explanation, numerous specific details are set forth in order to provide a thorough understanding of the disclosed embodiments. It will be apparent, however, that one or more embodiments may be practiced without these specific details. In other instances, well-known structures and devices are schematically shown in order to simplify the drawing.
It is also to be understood that the terminology used herein is for the purpose of describing particular embodiments, and is not intended to be limiting, as the scope of the present invention will be defined by the appended claims and equivalents thereof.
The singular forms “a,” “an,” and “the” include plural referents unless the context clearly dictates otherwise.
The term “coronary arteries” as used herein refers arteries that branch off the aorta to supply the heart muscle with oxygenated blood.
The term “endoprosthesis” as used herein refers to an artificial device that is placed inside the body of human or animal.
The term “lumen” as used herein refers to a cavity of a tubular organ such as a blood vessel, urinary tracts or bile ducts.
The term “hollow parison” as used herein refers to a hollow tubular polymeric mass before it is shaped into its final form. The term “molten hollow parison” as used herein refers to a hollow parison of molten polymer resin extruded from a die head assembly. The term “hot hollow parison” as used herein refers to a hollow parison partially cooled down to the stretch-blowing temperature of the polymer resin which made of it.
The term “peripheral arteries” as used herein refers to blood vessels outside the heart and brain.
The term “radial strength” as used herein refers the external pressure that a stent is able to withstand without incurring clinically significant damage. The necessary radial strength for most vascular applications is 0.8-1.2 bar [J. Chem. Technol. Biotechnol. 2010(85)744].
The term “resins” as used herein refers to designate any polymer that is a basic material for plastics.
The term “restenosis” as used herein refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated as by balloon angioplasty or valvuloplasty.
The term “stenosis” as used herein refers to a narrowing or constriction of the diameter of a bodily passage or orifice.
The term “stent preform” as used herein refers to a tubular material that has undergone preliminary engineering processes before being laser cutting or chemical etching into the stent structure.
The term “thermal degradation” as used herein refers to deterioration of the material by heat, characterized by molecular scission.
The term “stretch-blowing temperature” as used herein refers to the temperature at which a thermoplastic polymer is undergoing an expanding deformation. The stretch-blowing temperature of a thermoplastic polymer is usually in the range between the melting temperature and the glass transition temperature of the thermoplastic polymer.
The term “glass transition temperature” as used herein refers to the temperature at which a polymer changes from (or to) a viscous or rubbery condition (or from) a hard and relatively brittle one.
Currently, the most widely used bioabsorbable polymers have been polyglycolide (PGA), polylactide (PLA) and their copolymers. These bioabsorbable polymers are thermoplastic, linear, partially crystalline or totally amorphous polymers with a definitive melting temperature (Tm) and glass transition (Tg) region. PGA of high molecular weight is a hard, tough, crystalline polymer melting at about 224-228° C. with a Tg of 36° C. PLA is a pale polymer with a melting temperature at about 175-185° C., with a Tg of 55-57° C. Commercial PLA are copolymers of poly(1-lactic acid) (PLLA) and poly(d, 1-lactic acid) (PDLLA), while 1-isomer constitutes the main fraction. Depending on the composition of the 1- and d, 1-enantiomers, PLA polymer with higher 1-enantiomer content, tends to have higher melting temperature and higher glass transition temperature.
The materials used for bioabsorbable stents must provide certain essential safety-related mechanical properties, which include high initial strength, appropriate initial modulus and acceptable in vivo biodegradation rate. The high initial strength is required because the bioabsorbable stent must resist mechanical stresses during a surgical procedure and it must carry external and physiological loads during the healing stage of blood vessel. The appropriate modulus means that the bioabsorbable stent must not be too stiff and too flexible for the special purpose where it is used. The stent should possess ductile behavior so that it does not fracture with a brittle mechanism. The in vivo biodegradation rate of bioabsorbable polymer is essential for controlling the strength and modulus of bioabsorbable stent retaining in blood vessel. The loss of strength and modulus in vivo must be in coordinate with the healing of blood vessel. However, the mechanical properties of most bioabsorbable polymers are weak than the permanent metallic stent materials, such as stainless steel or cobalt-chromium alloy.
The mechanical properties of bioabsorbable polymers can be increased by reinforcing processes such as stretch molding, blow molding or stretch blow-molding. The polymeric articles made by the stretch blow-molding process give higher strength and modulus compared to those made by compress molding or injection molding process. The biaxial molecular orientation induced during the stretch blow-molding process will increase the strength and modulus of the resulting polymeric article. The polymeric articles made by the stretch blow-molding process is reinforced with oriented polymeric chains, fibrils, fibers, extended chain crystals and shish-kebab crystals, which have the same chemical composition as the polymer. In addition, the crystallites produced during strain-induced crystallization also reduce the aging effect since they can act as the physical crosslink to stabilize the amorphous phase, thereby reducing its brittleness.
Embodiments of the disclosure relate to a method and an apparatus for manufacturing a bioabsorbable stent from a thermoplastic bioabsorbable polymer resin. In particular, the embodiments of the disclosure relates to a method and an apparatus for fabricating a stent preform from a polymer resin in one-stage process with reduced thermal exposure time to the polymer. A perspective flowchart of a method in this disclosure is depicted in
In one embodiment of this disclosure, a method and an apparatus for producing a bioabsorbable stent are depicted in
In another embodiment of this disclosure, a method and an apparatus for producing a bioabsorbable stent are depicted in
In contrast to the bioabsorbable stents that fabricate expanded polymer tubes by two-stage process in two separate apparatuses, the embodiments of the disclosure provides a cost-effective method for manufacturing bioabsorbable stents that fabricate expanded polymer tubes by one-stage process and one apparatus. In the two-stage process, polymer resins are first processed into polymer tubes by casting molding, injection molding, or extrusion molding processes in one apparatus, then, polymer tubes are reheated and expanded by blow-molding process in another apparatus. Because the overall exposure time of PLLA in elevated temperature during one-stage process is less than that of two-stage process, the thermal degradation of expanded PLLA tubes made from one-stage process are likely smaller than that of expanded PLLA tubes made from two-stage process. In addition, it is possible to overheat the outside wall of the polymer tube when reheating the polymer tube to a suitable blowing temperature in two-stage process. In one-stage process of the embodiments of the disclosure, an expanded polymer tube is fabricated from a hot parison by a blow-molding process. The hot parison is made from molten polymer resins by an extruder and is uniformly cooled down to a suitable blow-molding temperature. These results will affect mechanical properties of expanded PLLA tubes, which in turn, will influence mechanical properties of final bioabsorbable stents. The embodiments of the disclosure is to provide a method for manufacturing a bioabsorbable stent by fabricating an expanded polymer tube from polymer resins in one-stage process and an apparatus for doing the same.
It will be apparent to those skilled in the art that various modifications and variations can be made to the disclosed embodiments. It is intended that the specification and examples be considered as exemplary only, with a true scope of the disclosure being indicated by the following claims and their equivalents.
This application claims the benefit of U.S. Provisional Application No. 61/483,447, filed on May 6, 2011, the entirety of which is incorporated by reference herein.
| Number | Date | Country | |
|---|---|---|---|
| 61483447 | May 2011 | US |
