JP3904598B2 - Method for manufacturing a system for delivering a stent to a body - Google Patents

Description

他の態様は、下記の請求の範囲内のものである。 Field of Invention
The present invention relates to an endoprosthesis stent that is placed within a body lumen that exhibits physiological movements such as peristaltic movement.
Background of the Invention
A medical stent is a tubular endoprosthesis that is placed in the body to perform functions such as opening the body cavity, eg, a passage blocked by a tumor. Generally, a stent is carried into the body by a catheter that holds the stent in a miniaturized form so that the stent is displaced to the desired site. When the desired site is reached, the stent expands to engage the lumen wall. This expansion mechanism, for example, inflates the stent carried out radially, such as inflating a balloon carried by the catheter and inelastically deforming and securing the stent in contact with the lumen wall in a predetermined expanded state. It may include expanding in the direction. The dilatation balloon is then deflated and the catheter is removed.
In another method, the stent is formed of a highly elastic material that self-expands after being miniaturized. Upon introduction into the body, the stent is restrained in a miniaturized state. When the stent is delivered to the desired site of implantation, this constraint is removed and the stent self-expands due to its own internal elastic recovery force.
Esophageal strictures cause obstructive dysphagia, resulting in debilitating malnutrition. To date, the theoretical benefits of placing a plastic stent to restore the patient's ability to perform include technical difficulties in mounting, morbidity and mortality associated with this method, and prosthetic devices This is offset by the poor performance over the long term. In particular, conventional stents improperly transmit peristaltic wave forces and distortions, for example, stents can enter the stomach, puncture the esophagus, or rupture the aorta.
Summary of the Invention
In a first aspect, the present invention relates to a method of providing reinforcement in the lumen of a peristaltic organ. The stent is formed by knitting filaments into a knitted loop, the pattern of the loop shifting from a relaxed state, with each row of loops axially and independent of the rows on either side Selected to do. The local extension and shortening allowed by this shift allows the stent to adapt to the peristalsis of the organ without moving within the organ.
A preferred embodiment of the stent is characterized by the following. The lumen to be treated is the esophagus. The expansion coefficient (factor) ε at which the stent can be locally expanded by the shift is in a relation represented by the relational expression: ε = 1.0 / cos θ with respect to the angle θ at which the lumen can tilt inward. The stent is knitted with a metal wire in a self-expandable manner so that the elastic recovery force of the metal wire allows the stent to expand outward relative to the body lumen wall. The stent is knitted with a nitinol wire having a diameter of about 0.15 mm. The stent has a contracted cross section in an unconstrained state. This contraction may have a valve.
The stent according to the present invention has the following advantages. This stent exerts a certain gentle radial force on the lumen wall, but this force maintains the lumen patency and actively resists compression by tumors and the like. The inherent flexibility of the knitted stent adapts to the peristalsis and transmits the peristaltic wave to the lumen, but does not change or shift the overall length. This reduces complications and improves long-term stability, patency, and patient comfort. The force exerted by the stent on the lumen is sufficient to squeeze the organ capillaries, thus preventing growth into the lumen. The stent can be delivered through a bulky delivery system that is smaller than a standard endoscope. The small diameter of the delivery system eliminates the need for pre-dilatation of the stenosis, making it easier to implant and being penetrated by patients with a tortuous esophageal anatomy and plastic stents. Even patients with easy stenosis can be placed.
In a second aspect, the invention features a stent for providing reinforcement to a selected region of a selected body lumen. The stent consists of two cylindrical mesh layers and a semi-permeable compliant membrane sandwiched between them.
Preferred embodiments of the present invention are characterized by the following. The two mesh layers may be flexible filament knits, which can be formed so that the stent can adapt to peristalsis of the body lumen. This membrane consists of expanded polytetrafluoroethylene.
The present invention or a preferred embodiment thereof is characterized by the following advantages. The semipermeable membrane prevents the cellular ingrowth of the stent. The force exerted on the lumen by the stent is sufficient to compress the organ's capillaries, thus preventing growth into the lumen.
In a third aspect, the present invention provides a method of manufacturing a delivery system for a resilient tubular device such that the tubular device is inserted into the body with a substantially reduced diameter. This method uses a confinement block having a hole and a groove leading to the hole. The tubular device is sandwiched and inserted into the hole and groove. Two mandrels are inserted into this hole, one inside the tubular device and the other outside. The mandrels rotate relative to each other until the tubular device is fully wound and confined within the hole with a reduced diameter, winding the tubular device itself. The tubular device is removed from the hole while constrained to the reduced diameter.
A preferred embodiment of the manufacturing method has the following characteristics. The removal step can be performed by pushing the tubular device from the end of the hole and restraining it as it exits the tubular device. This constraint may be done by wrapping a wire around the tubular device. The groove may be tangent to the hole in the confinement block. The tubular device may be a stent knitted with elastic filaments. One of the mandrels may be part of a delivery system used to deliver the stent.
The novel method of manufacturing a stent has the following advantages. Some conventional methods require several operators to simultaneously hold and hold the stent bounce and injure the operator's fingers. The new method can be performed easily by one operator. The stent delivery system produced by this method is more uniform in both the distribution of stress within one stent and the variation between stents than the stent produced by the conventional method, so that Prevents distortion of the stent and allows the physician to more accurately place the stent on the patient. The stent delivery system produced according to the present invention has a small profile, thus minimizing trauma to the patient during implantation.
In a fourth aspect, the present invention relates to a method for manufacturing a wire medical device. The method includes the following steps: bending the elastic wire into a regular pattern to define a tube wall that generally has a substantially constant outer diameter and geometry and extends to a desired axial length; and Applying a mechanically deforming force to the tube to conform to the device diameter or geometry, forming the tube to form devices with different desired diameters and dimensions, while the tube Heating and then cooling the tube to maintain the desired diameter or geometry when the mechanical deformation force is removed while maintaining the deformation force.
A preferred embodiment of this production method includes the following features. The mechanical deformation force is applied by confining the tube within a die cavity of a smaller diameter than the tube and / or by stretching the tube over a mandrel having a larger diameter than the tube. The deformed shape defines a stent that expands to a larger diameter at one or both ends than at the rest of the stent. This deformed portion of the tube can extend over at least about 10% of its length.
In a fifth aspect, the present invention relates to a method of manufacturing a wire medical device and includes the following steps: Bending the elastic wire into a regular pattern to have a desired outer diameter and geometry; Defining a generally tubular medical device wall extending to a desired axial length, the pattern including portions where the wires overlap and contact; and heating the tube for a predetermined time to stress the overlapping contacts Removes and then stresses the overlapping contact formed when the tube cools and bends, improving the device's ability to respond to internal movement when subjected to physiological movements of the body lumen Let
A preferred embodiment of this production method is characterized by the following. The tube is formed by knitting a wire. The knitting machine has a knitting head formed from a durable, low friction polymer material such as delrin or nylon. The portion of the knitting needle that is in contact with the wire during knitting contains a durable, low friction polymeric material. The wire is selected from the group consisting of a nickel-titanium alloy and molybdenum or barium containing high modulus stainless steel. The tube is heated at about 400-500 ° C. for about 20-30 minutes.
In a sixth aspect, the present invention relates to a medical device for use in a body lumen. The medical device includes a tube formed by a flexible wire that is bent into a regular pattern to form the wall of the tube over a desired length, which pattern is the physiological of the body lumen. It is configured to allow relative movement of adjacent portions of the wire when undergoing movement. The tube has an outer diameter or geometric shape that varies along its length, which shape is selected to enhance the function of the device in the lumen in which it is used.
A preferable aspect in the sixth aspect includes the following features. The shape selected complements the inner wall of the lumen in which the device is used. The shape of the tube is to have a larger diameter part of its length in order to contact a correspondingly larger diameter lumen part and to contact a corresponding smaller diameter lumen part. The rest of its length is shaped with a smaller diameter. The tube shape is flared to a larger diameter at one or both ends, with a smaller diameter portion adjacent to the flared end, and the smaller diameter portion selected to fit the lumen diameter And the flare serves to secure the device to the lumen. The flare extends to a diameter that is about 10-25% larger than the small diameter portion, and the length is about 15-25% of the length of the tube. The tube has a shape with a second flare formed by a wire at the end of the pattern. A regular bending pattern is formed by knitting.
In a seventh aspect, the present invention relates to a valve device for implantation into a lumen of a body organ. This device is formed by a flexible wire that is bent into a regular pattern to form the wall of the tube over the desired length, which pattern is subject to physiological movement of the body lumen. , Formed to allow relative movement of adjacent portions of the wire. A portion of the tube has a very small diameter, providing a material that is substantially impermeable to bodily fluids. This device can prevent the flow of bodily fluid through the reduced diameter portion until the body fluid pressure is sufficient to elastically expand the reduced diameter portion and flow through it. Some of them relax to their reduced diameter when the body fluid pressure subsequently decreases.
In a preferred embodiment, this valve is formed as a valve for the urethra. Other advantages and characteristics of the invention will be apparent from the following description of the preferred embodiments and from the claims.
[Brief description of the drawings]
1 and 1c are plan views of a stent according to the present invention.
FIG. 1a is a front end view of the stent.
1b, 1d-1h, 4a, 6, and 11g are detailed views of the knitted loop of the knitted stent.
2, 2a, 3, and 3a-3e are cross-sectional views of the body showing the effect and operation of the stent in the esophagus.
FIG. 4 is a cross-sectional view of the peristaltic organ.
5, 5a and 5b are schematic views of another embodiment of a stent.
FIG. 5c is a partially separated view of another embodiment.
6a, 7, 7a, 7c-7j, 7l, 7m, 7p-7s, 10a-10c, 11, and 11a-11f are a time sequence of steps and a perspective view of the tool in a method for manufacturing a stent delivery system. It is.
FIG. 6b is a perspective view of another embodiment.
7b, 7k, 7n, and 7o are cross-sectional views during the manufacturing method of the delivery system.
FIG. 7t is a cross-sectional view of the delivery system.
FIG. 7u is a cutaway perspective view of the delivery system.
8 and 8a-8e are time series of cross-sectional views of the esophagus showing stent delivery.
9, 9a, and 9b are time series of cutaway views of another delivery method.
FIG. 10 is a partially cutaway perspective view of a die and a stent formed in the die.
Description of preferred embodiments
Referring to FIGS. 1 and 1a, a stent 100 according to a preferred embodiment is formed from a knit cylinder of length L and diameter D. By knitting, a series of loosely entangled loops (eg, adjacent loops 132 and 134 in FIG. 1b) are formed that slide together. This sliding or shifting allows the stent to adapt to organ movement without moving axially within the organ. This adaptation can be done by simply bending the stent filament.
When a stent is locally compressed radially, it maintains its axial working length L by locally expanding or contracting by shifting the row of loops relative to each other. FIG. 1c shows a region 130 of the stent that is not under radial compression in which the adjacent loops 132 and 134 are in a relaxed shape with overlapping, the heads of these loops being separated by a short distance s. In the case of the esophagus, large pieces of food widen the esophagus. At the first moment of expansion, the wall may deflect by an angle θ, but the organ diameter does not change much. As in area 140 in FIG. 1c, the local length of the wall extends by a factor 1 / cos θ. As shown in FIG. 1e, the separation of the head is the loop length l.₁The stent loop train shifts axially due to the elastic deformation of the loop wire. In the maximum inflated region 150, the length of each part of the esophagus returns to its resting length, but the diameter has increased. The stent knit loop can be expanded to accommodate this extension, as shown in FIG. Considering any peristaltic organ again, this organ contracts (c in FIG. 1c) and compresses one region. In the region 160 where the wall is at the deflection angle θ but the diameter is essentially equal to the resting diameter, the organ wall length extends by a factor 1 / cos θ, and the loop is essentially equal to the resting width. The same, but with the relative shift in the axial direction of the rows of loops, the length l shown in FIG.₁Pass through the state of extending to 再 び again. In the maximum compression region 170, the organ wall is at its resting length, but the circumference is significantly reduced. In this region, the stent loop deforms into the shape of FIG. 1g, where the loop length is s but the width is compressed. Eventually, as the peristalsis relaxes, the organ wall returns to the resting length of region 190 in FIG. 1c, and the resting circumference, and the stent loop returns to the overlapping resting shape of FIG. 1d.
For organs that can contract to an almost closed state, as the lumen is compressed radially, the circumference shortens, so that a portion of the filament length that contributes to the circumference of the stationary stent is reduced. Released and contributes to the length, so the loop has a length l as shown in FIG.₂Can stretch.
s to l₂Elongation to occurs only by elastic bending deformation and slipping of the rows of loops against other rows without significant elongation of the stent's own filaments. The ratio of the maximum local length to the relaxed local length, l₂/ S depends on the loop shape and the elastic limit of the filament material.
Referring again to FIG. 1c, local stretch in the region of tilt, compression or radial extension does not substantially affect loops in nearby regions that are not subjected to radial compression, and these loops are They can stretch, contract, and expand in response to their own local movement of the organ. In this way, the stent maintains its overall working length L even when locally expanded or compressed. When the radial compression is released, the adjacent loop in the compression region slides back into the relaxed overlap state separated by the distance s, so that the stent is brought back to its original resting diameter D by the elasticity of the filament without changing the overall length. Inflate and return. The stent, on average, is about one loop size in the local area and maintains point-to-point contact with the organ so that the stent maintains position in the organ and does not move with peristalsis.
A further characteristic of this stent is that the condition characterized by the fact that all adjacent loops are separated by a distance s is a stable balance between the elastic recovery force of the wire and the compression force of the esophagus. Therefore, the stent is automatically adjusted to the full length L regardless of the initial shape of the loop and the length of the stent. For example, if the loop is in the stretched state as in FIG. 1e or 1h, when shrunk, the loop adjacent to the compression region is pulled inward in the axial direction to the relaxed shape of FIG. Pull the end (120 in FIG. 1) and the distal end 122 inward. When compression is released, the ends of the stent 120 and 122 are pulled inward, and the loops in the compression region also shorten as the adjacent loops slide inward and adjust for the reduced overall length. Once the stent has settled in this equilibrium, the overall length and position within the lumen is stable.
These features are made possible by the sliding movement of adjacent filament loops in the preferred stents of the present invention, which sliding movement reduces the overlap of loops as described, filament elasticity, and loops. This is possible depending on the shape. The sliding motion allows the stent to extend or shorten locally in an axial direction substantially independent of other remote portions of the stent. The elasticity of the stent filaments allows the stent and lumen to return to their desired open state without inelastic deformation when the compressive force is removed. Since the elastic recovery force seeks its minimal stress, the relaxed state of FIG. 1b, the loop is shaped so that the stent has the desired diameter and overall length. This minimization occurs when an adjacent loop touches its widest point, eg, point 136 between loops 132 and 134.
These features are particularly useful for passages in the body that perform luminal wall movements such as peristaltic movements of the esophagus as physiological functions. For example, referring to FIG. 2, the esophagus 200 is occluded by a tumor 202. In FIG. 2a, the lumen patency is restored after the stent 100 is inserted. Once implanted in the esophagus, the stent is held from a free diameter when outside the body to a static diameter slightly compressed from the esophagus. It is the elastic recovery force of the stent that resists this compressive force that holds the stent in place.
A stent for an organ such as the esophagus according to the present invention not only keeps the lumen open, but also allows the organ to maintain physiological movement. Furthermore, the stent adapts to this movement without peristaltic movement itself--it does not shift towards the stomach or change its overall length with each esophageal contraction. The operation of the elastic knitted stent is shown in FIGS. 3 and 3a-3e. A small piece of food 310 is pushed through the lumen 320 by a peristaltic wave 322 propagating down the esophagus. This wave is triggered by the circumferential contraction of the muscular tissue around the lumen, so that the wall extends radially inward. Before this wave reaches the stent portion, the stent is positioned as a length L between points T and B. As the wave reaches point T and the portion of the esophagus reinforced by stent 100, the stent follows radial contraction, as shown in FIGS. 3a-3e. As shown in FIG. 3e, after the peristaltic wave passes, the stent has not moved from points T and B and maintains its full length L. This adaptive and recovery characteristic eliminates the axial movement of the stent that occurs in a monolithic structure in which stress in one part is substantially transmitted to the other part, maintaining the stent in position within the lumen. can do.
With reference to FIGS. 4 and 4a, the shape of the knit loop of the stent is determined based on the degree of radial motion due to the peristaltic motion and the resulting axial extension. In general, the stent loop length at the extended position l is:
l = εs (1)
Where s is the axial length of the body lumen portion, for which two adjacent loops of the stent extend, and ε is the stent where the body tube It is a factor that must be stretched corresponding to the local stretching of the cavity wall. It can be seen that the maximum local extension occurs at the wall portion of the maximum angle θ from the rest position. At the maximum limit, the entire peristaltic wave has a wall at an angle θ, and as a result, it can be approximated as a part of a lumen of a stationary length a extending to a triangular wave having a hypotenuse b of a right triangle. Therefore,
b / a = 1 / cosθ (2)
It is. This ratio b / a is an extension factor that the loop must extend from its relaxed length s to its extended length 1 in order to accommodate the extension of the lumen wall at the slope of the peristaltic wave. Therefore,
b / a = 1 / cos θ = l / s = ε (3)
It is. In order to maintain the point-to-point contact with the lumen wall as a whole and to keep the stent as a whole along the lumen axis, in the region of peristaltic compression, The loop head can be slid locally at a distance of (l-s) / 2.
Regardless of the amount of deflection c, even in the extreme case where the organ can be compressed until it completely closes (when deflection c in FIG. 4a is equal to the lumen radius 2 / D), the stent is reduced to 1 / cos θ. It can be seen that the stent accommodates the lumen wall extension if it is possible to have equal local extension ε.
The amount of force exerted by the stent against the lumen wall is selected to be greater than the blood pressure in a typical tumor capillary, thereby preventing the tumor from growing further into the esophageal lumen. This force is determined by the elastic modulus of the filament, by the loop shape and knit density (loops per unit of axial length) and by the lumen and stent diameter. For example, by selecting a stiff material for the filaments and knitting the stent with more and smaller loops per unit length (reducing the s and changing the shape of the loop to maintain the l / s ratio) Or by knitting the stent to a larger resting diameter D, the stent design can exert a greater force on the lumen wall. This radial force is bounded at the point where the loop reaches the relaxed shape of FIG. 1b and the diameter of the stent reaches diameter D (this force is zero at the contact 136 and the force acting on the lumen itself is zero). become). Thus, if the stent is to maintain its position in the lumen, the stent diameter D must be slightly larger than the lumen diameter.
Referring again to FIGS. 1, 1a and 1b, a particular embodiment for use as an esophageal stent is knitted with a Nitinol wire having a diameter of about 0.15 mm and has a diameter D of about 18 mm, but also has a diameter of 14 to 25 mm. Can be used. Proximal end 120 is flared to 20 mm to ensure fixation to the esophageal wall. By manufacturing a stent with a total length of 5 to 15 cm, it is possible to select a stent that meets the needs of the patient. The relaxation loop length s is about 0.80 to 0.85 mm, and the maximum loop length l obtained without significantly distorting the loop is about 1.05 to 1.15 mm. This elongation factor of about 1.4 is close to the square root of 2 and allows for a maximum angle θ of about 45 °. When the loop is in its resting state, the height P of the apex from the apex of the loop is about 2.2 mm.
Examples of filament materials include shape memory metals such as nitinol, tantalum steel, stainless steel or other elastic metals, or plastics such as polyester, polypropylene, or carbon fibers. To select the filament to have a sufficiently high elastic limit and expand the stent, the delivery system can be made completely dependent on this elasticity rather than, for example, the balloon 820 of FIG. 8e. The filament can be formed from a two-component metal wire system that exhibits desirable physical properties such as high radiopacity along with desirable mechanical properties such as ultimate elasticity. Composite medical wires are described in detail in US patent application Ser. No. 07 / 86,253 entitled “Medical Wire” by Kevin R. Heath. The contents described therein are included in the present invention as a reference. The stent is a knit of two or more filaments.
As shown in FIG. 4a, the stent is knitted with a single filament. This figure shows only the front half of the stent. Ordinary screws have only one ridge from head to tip, but the loops represented in this figure as separate rows are actually a series wound around a single helix. In another aspect of the stent, multiple filaments or other knits can be used as long as the knit structure stretches a single row without shifting two adjacent rows. The last loop of wire is segment 440. To prevent the stent from unraveling, the last three loops of the stent (two of 450 and 452 are shown) are applied with urethane as shown in FIG. 4b. This application also covers the pointed end of the filament.
It can be seen that the stent is applicable to malignant or benign occlusions in many other organs. A stent for treating bile duct obstruction may be about 8-10 mm in diameter and 4-8 cm in length, for example when treating liver sclerosis or bleeding. The stent for the ureter may be about 6-10 mm in diameter and about 2-10 cm in length. The urethral stent may be about 10-20 mm in diameter and about 2-6 cm in length. The stent for the prostate urethra may be about 10-20 mm in diameter and about 2-6 cm in length. The colonic stent may be about 10-20 mm in diameter and about 4-10 cm in length. The stent for the hemodialysis shunt may be about 6-8 mm in diameter and about 2-6 cm in length. The stent for the portal vessel may be about 8-14 mm in diameter and about 4-8 cm in length. The stent for the trachea and bronchus can be about 8-25 mm in diameter and 1-8 cm in length. Stents for occlusion of the gastric outlet may be about 8-20 mm in diameter and 1-25 cm in length. Peristaltic stents can also be shaped for aortic aneurysms or dissection (preferably interweaving filament material with a skin covering such as dacron) and for the treatment of superior vena cava syndrome and venous restraint You can also. The present invention is also useful for lumens where compression is caused by some external force, for example, blood vessels that are compressed by muscle contraction, limb movements or pressure caused by objects present outside the body.
Figures 5, 5a, and 5b represent another form of stent. (These figures only represent the shape of the stent, as the knitted loops in these perspective views are confusing, so these figures represent only the shape of the stent.) Can be shaped to include. This indentation causes the stent to conform to the anatomy of a natural sphincter structure, such as the pylorus or heart. Such a deflated stent allows its trachea to close, for example to prevent reflux. For example, this indentation can be shaped at one end for use in the rectum of the anus, or the total duct of the papilla.
This indentation can be conically shaped for use in the sphincter organ, as shown in FIG. FIG. 5a shows a stent incorporating flattening in which the circumference in the flattened region is reduced to keep its width constant. The latter embodiment can be used for occlusion by two lips such as vocal cords. In either case, the loop in the region of contraction is shaped such that its free state is similar to one of the compressed shapes (eg FIG. 1g), so that the contraction can be opened to a stationary diameter D. It is desirable to do so.
As shown in FIG. 5b, the blank space can incorporate a valve to completely close the stented organ, such as a reinforced lip 520, for example. This valve can be opened and closed by the muscles normally surrounding the point of contraction or by manual control extending outside the body. This allows the stent to be used, for example, across an aortic valve or as a substitute for a urinary sphincter. It is desirable to reinforce the point of contraction, for example with a hard wire, particularly in connection with the flattened contraction of FIG. 5a. It is also desirable to attach a watertight membrane to the valved stent in the form shown in FIG. 5c and described below.
A stent can be constructed to exert a force that varies along its length, for example, by changing the density of a wire gauge or knit. In the case of the above-mentioned crimped stent, it is desirable to make the stent particularly flexible in the area of the crimp.
Some tumors are very invasive and soon grow into the stent. As shown in FIG. 5c, an elastic semipermeable membrane 530 with a porosity of less than 50 microns and a very low modulus can be sandwiched between two knit layers to produce a stent. This membrane is preferably foamed polytetrafluoroethylene (Teflon) or latex. Inner layer 532 is essentially the same as a single layer stent and provides most of the elastic force to the lumen. The outer knitted layer 534, which serves to retain the membrane, is generally composed of a thin wire, such as 0.07 mm in diameter, or a low modulus material such as polypropylene or polyethylene. The outer knitted layer is slightly shorter than the inner layer.
The stent is knitted on a conventional knitting machine very similar to the knitting machine used to knit stockings. During the knitting process, the wire is deformed beyond its elastic limit. Referring to FIG. 6, it is preferable to make the knit with a “upper loop” 610 that is different from the “lower loop” 612 in some types of knitting machines or for stents of a certain diameter. In certain applications, for example in the case of the aorta, it is important to make the loop uniform so that the stent exerts a uniform pressure along the lumen wall. During the knitting process, the wire is under tension so that the loop is in a tensioned shape similar to FIG. 1e or similar to FIG. 1f or 1h depending on the shape and configuration parameters of the knitting machine itself.
The knitting machine forms a “rope” with a long knit loop. This rope is cut to a length somewhat longer than the final length of the stent. This extra length allows for the shortening of the stent that occurs when the loop is shortened from the knitting machine to the resting state of FIG. As shown in FIG. 6a, after knitting, a stent is attached to the mandrel 620 for annealing to remove strain induced by plastic deformation due to knitting and give the wire greater elasticity. This mandrel is in the free shape of the stent, is 18 mm in diameter and has a 20 mm flare 622 at one end.
In order to obtain the contracted embodiment of FIGS. 5 and 5a, the mandrel has a constriction formed therein and an external constraint is applied to the stent so that the annealed shape is the shape shown in those figures. Applied. When the stent is attached to the mandrel, the operator shortens the overall length so that the loop is in the relaxed, shortened state of FIG. The stent is annealed at about 450 ° C. for about 15 minutes.
After annealing, the stent is cut to a final length and urethane is dropped on the three loops at each end of each stent to prevent unwinding (450 in FIG. 4a or FIG. 6).
Alternatively, the stent may be an interlock pre-formed sine ring knit, two of which are shown in FIG. 6b.
The stent is placed in a delivery catheter as shown in FIGS. 7 and 7a-7u. The center of the delivery catheter is a carrier tube 700 as shown in FIG. The carrier is a flexible tube of Pebax, a polyether / polyamide-12 resin from Atochimie with the desired flexibility / stiffness properties, 2.5 mm in diameter and about 80 cm in length. This carrier has several radiopaque O-rings 704, 706 attached up to 20 cm distal. A preferred radiopaque material is tantalum.
Several tools are used in the preferred method of attaching the stent to the carrier tube 700: a containment block, two mandrels, and a pusher. The containment block 710 shown in FIGS. 7a-c is cylindrical and is about 20 cm, somewhat longer than the stent itself, and is made of a hard plastic with low friction properties, preferably Delrin or nylon. This block has an 8 mm diameter hole 712 and a 1 mm wide slot 714 that cuts from the outer top of the block and meets the hole 712 tangentially. The slot and hole may have a guide passage 718 formed to facilitate the next step. The block may also have a flat 716 scored at the bottom so that the block is secured in a vise. The first mandrel 720 shown in FIG. 7d is a simple rod about 30 cm long and about 3 mm in diameter. The second mandrel 722 shown in FIG. 7e has a diameter of about 3 mm, a shaft that is longer than the containment block, two handles 724 each about 10 mm in diameter with a central hole that friction fits at the end of the mandrel shaft, and It has a slot 726 that is wide to accommodate the carrier tube. Both mandrels have round ends so as not to catch the stent loop. A third tool is the pusher 728 shown in FIG. 7f, with a shaft 729 having a diameter slightly smaller than 8 mm and a hole 730 that is somewhat larger than the outer diameter of the carrier 700. This hole may be the full length of the pusher or have a rear hole 732 as shown in FIG. 7f. The fourth tool seen in FIG. 7q is a flexible copper wire with a silicone sheath 760 (silastic) on top. The sheath wire is about 50 cm long and the sheath is about 1-2 mm in diameter.
Referring to FIG. 7g, the containment block 710 is secured to the vise 750. The operator squeezes the stent 100 flat and puts it into the slot 714, preferably starting from the corner 752 and finally into the hole 712. Referring to FIG. 7h, the stent is placed in the containment block such that the proximal end 120 of the stent extends from the end of the containment block. The first mandrel 720 is inserted into the distal end of the carrier 700 and then the mandrel and carrier tube are passed through the center of the stent. As shown in FIG. 7i, the operator slides the stent to the center of the containment block. Referring to FIG. 7j, the stent is slipped back so that the open distal end again extends from the end of the containment block. One handle of the second mandrel is removed and the second mandrel shaft 722 is inserted into the hole in the containment block, but outside the stent. Referring to FIG. 7 k, the first mandrel 720 is inside the carrier tube 700, which is located inside the stent 100. The lower portion of the stent and the second mandrel 722 are present inside the hole 712 of the containment block. Referring to FIG. 71, the stent is slipped back to the center of the containment block. Several back-and-forth slides distribute the knit loop evenly over the length of the stent. The second handle of the second mandrel is attached to the shaft of the second mandrel and the slot 726 of the handle is engaged with the carrier tube 700 and / or the first mandrel 720. An operator can center the O-rings 704, 706 in the center of the stent so that the carrier is axially positioned almost exactly within the stent. Referring to FIG. 7m, the operator twists the handle, rotates the two mandrels around each other, and wraps the stent around the two mandrels. FIG. 7n shows the placement of the stent and two mandrels after about half a turn. The operator continues to wind until the stent is completely wound into the hole in the containment block. The operator removes the handle from the second mandrel and removes the second mandrel from the containment block. As shown in FIG. 7o, the containment block keeps the stent in a wound configuration.
Referring to FIG. 7p, the operator attaches a pusher 728 to the proximal end of the carrier, with the shaft 729 at the distal end. The pusher is used to slowly push the stent out of the hole in the containment block.
Referring to FIG. 7q, the pusher is used to push the stent about 1 cm from the containment block. The operator makes any final adjustments necessary to place the radiopaque O-ring in the center of the stent. The operator winds the copper wire and silicone sheath 760 around the exposed distal end 122 of the stent several times with a spacing of about 1 mm for each rotation and places the coated wire bit into the slot of the containment block. . Referring to FIG. 7r, the operator uses a pusher to gradually feed the stent out of the containment block and wrap the covered wire around the stent to limit the diameter to about 8 mm. The operator maintains the wrapping at a substantially uniform 1 mm spacing.
After the stent is fully constrained within the copper wire 760, the pusher is returned to the proximal end of the carrier tube, the carrier tube is pushed out of the hole in the containment block, and the stent and the silastic / copper wire packet are immersed in USP grade dissolved gelatin. And harden the gelatin. The copper wire can then be unwound; the silicone sheath acts as a release surface, the gelatin peels from the wire, and settles on the stent as a 1 mm wide “threaded” strip, 770 in FIG. Confine.
Referring to FIGS. 7t and 7u, the stent delivery system catheter 799 is completed by adding a nosepiece 772 to the distal end of the carrier 700 and covering the entire assembly with a cylindrical sheath 774. Both the carrier and the sheath are essentially rigid in the axial direction so that they can be used to push and pull to position the catheter, and the handles 782 and 784 are squeezed together to make the sheath stent Can be withdrawn from. Also shown in FIG. 7t is a radiopaque marker 704, 706 and in FIG. 7u is a scale line 778 on the sheath, both of which are used to guide positioning during implantation. The pair inside the marker indicates the length when the stent is fully expanded with a diameter of 18 mm, and the outer pair indicates the length of the stent when compressed to a diameter of 8 mm. The guide line 778 is threaded through the central hole 776 of the carrier during implantation.
The stent implantation process is shown in FIGS. 8 and 8a-8e. Referring to FIG. 8, an endoscope 810 is used to identify the proximal end 812 of the stenosis 814. A guide line 778 advances through the stenosis. In FIG. 8a, an 8 cm long balloon 820 is advanced on the guide line to inflate to a diameter of 12 mm and widen the stenosis to 12 mm. After examining the stenosis with an endoscope and fluoroscopy, a gelatin-coated stent that is 4-6 cm longer than the stenosis is selected. The delivery system 799 is passed over the guide line and advanced until the inner radiopaque marker 704 is 2-3 cm farther from the distal end 832 of the stenosis.
Referring to FIG. 8c, compressing handles 782 and 784 (see FIG. 7u) together retracts the outer sheath and begins to deploy the stent. As the gelatin begins to dissolve immediately, the stent expands due to its own elastic recovery. The stent filament material, Nitinol, is selected so that the elastic limit is not exceeded, even with the considerable deformation required to compact the stent into a delivery system. Referring to FIG. 8d, the catheter 799 can be removed after the proximal and distal ends 120, 122 of the stent have spread and firmly attached to the esophageal wall. Referring to FIG. 8e, in some patients, a 12 mm diameter balloon 820 is inflated within the stent to allow the occlusion to be opened to the desired degree of opening, to secure the stent to the esophageal wall and for endoscopy. Ensure proper esophageal lumen size. Peristaltic contraction of the esophagus allows the stent to “settle” into its most relaxed shape.
Referring to FIG. 9, in another delivery system, from an elastic filament material selected to produce an internal recovery force by compacting that causes the stent to return to a static diameter after compression suppression is removed, 100 is formed. The stent may be compressed onto a catheter 900 that includes a sleeve 902 that retains the stent in a relatively compact state. This compaction is generally accomplished by rolling up the stent using two mandrels as shown in FIGS. 7g-7o. In other cases, the stent can be placed coaxially with the catheter. A catheter is placed in the lumen of the tumor 202 region. In FIG. 9a, the sleeve can be removed from around the stent, for example by pulling in the axial direction of arrow 910, so that the stent 100 can expand radially by release of internal recovery force. As shown in FIG. 9b, the axial force exerted by the stent is sufficient to widen the lumen 200 by pushing the tumor 202 outward, in some cases compressing the occlusion towards the lumen wall. Enough to do. The catheter can then be removed.
One aspect of the present invention is to form a stent having a shape or profile along the length of the stent that is adapted for a particular application in a particular lumen of the body. This profile can be used after weaving a stent of constant diameter and geometry, mechanically deforming the stent to a desired shape with a different diameter or geometry, and then after the mechanical deformation force is removed. This is accomplished by heat treating the stent so that its shape can be maintained. The deformation force applied prior to heat treatment is usually less than the force required to plastically deform the stent wire. However, plastic deformation before heat treatment may be used.
The shape of the stent is selected to have a variable diameter, such as a flare at one end that helps secure the stent in a lumen with inherent physiological lumen wall motion such as peristalsis can do. The transition between the large diameter flare and the small diameter portion will become increasingly tapered. In esophageal stents, this flare is usually located upstream of a small diameter section. The large diameter of the taper and flare provides a smooth transition to a small diameter section and reduces the likelihood that a food piece will be able to grip the stent. This flare is preferably of a substantial length that extends into a number of rows of knit loops. This flare is, for example, about 5-25% of the total length of the stent and may extend to a substantial width, for example, a diameter that is 5 to 35% larger than the diameter of the stent body. This flare may be, for example, a non-uniform diameter similar to a trumpet bell.
Stents can also be shaped to capture the changing diameter of body lumens. For example, a stent for a bronchial tube includes a 15 mm diameter portion located in the trachea and a small diameter portion e.g. 11 mm extending into the bronchus (branch). The stent has a tapered transition region approximately 1 cm long between the two portions. For use in the colon, the stent can have flares at both ends to secure the stent to the lumen wall at both ends. Other important lumens include the bile duct, prostatic urethra, and vasculature, in which the axial extension movement of, for example, the wall of the coronary artery or aorta is described above with respect to the radial folds in peristalsis. It is accommodated by the movement of the knit loop in the same way.
Referring to FIG. 10, a stent can be formed by applying a confinement force to a uniform stent using a die 1000 at a desired location along its length. The die 1000 is made of a heat-resistant material such as a 1 mm thick 360 stainless steel pipe. The die 1000 includes an inner diameter portion 1002 that is essentially equal to the static outer diameter D of the desired stent 100 and has an expanded portion 1004 with an inner diameter equal to the outer diameter of the desired flare. The die has a portion 1003 with a gradual transition between a large diameter portion and a small diameter portion. The flared end 1006 of the stent extends beyond the die and is wrapped around the outside of the die and held in place with a retention wire 1008. The assembly is heat treated. After heat treatment, the wrapped end 1006 is cut so that the stent can be removed from the die. After removal, the stent maintains the shape of the die. (Alternatively, the die may have a contraction portion 1002 with a cross-section equal to the desired cross-section of the desired contraction, as shown in FIGS. 5 and 5a.)
An advantage of using a die for confining the stent is that the end loop of the stent can be formed to have the same diameter as the adjacent loop of the stent body. Generally, after heat treatment with a mandrel, the knit end loop 1070 extends outward due to residual stress, forming a short end loop flare, as shown in FIG. 10c. This short end loop flare is often beneficial, for example in the case of an esophageal stent, which helps to secure the stent and prevents food fragments from collecting in the end loop. The short end loop flare is most useful as the second flare at the end of the main fixed flare and at the upstream end of the stent. In other applications, this flare is not necessary and may damage the lumen wall tissue. The second flare is also particularly useful for a delivery system that attaches to a stent loop, as described, for example, in US patent application Ser. No. 08/065238, filed May 20, 1993, which is described therein. Is included in the present invention.
Referring to FIG. 10 a, in another aspect, one end of stent 101 is confined within die 1010 and the other end is stretched over mandrel 1012. After the heat treatment, one has a diameter that is larger than the diameter of the original knit and the other end that is smaller than the diameter of the original knit, maintaining a shape with a transition region therebetween.
Referring to FIG. 10b, in another embodiment, both ends of the stent are stretched over separate mandrels, 1014, 1015. After heat treatment, when the mandrel is removed from the end, the stent retains an enlarged diameter flare at both ends and has a diameter intermediate portion corresponding to the original knit diameter.
Various forms of stents can be constructed by using these techniques. The cross-sectional shape of the stent can be changed by changing the shape of the die or mandrel. This forming technique can also be used to manufacture medical devices other than stents. For example, in certain embodiments, a knitted tube can be made, for example, for a heart by shaping the stent so that one end has a diameter approximately equal to the diameter of the lumen and the other end has a neck down to a very small diameter. Can be formed into vascular valves. After shaping, the knit mold is covered with a blood impermeable polymer, such as silicone or urethane. The device is implanted in a blood vessel with the large diameter end anchored to the vessel and the small diameter end directed downstream. The small diameter end substantially blocks the blood flow until it has enough blood pressure to open it, and a large amount of blood flow until the pressure drops again and the end of the device relaxes again to its small diameter state. Through. The thickness and elasticity of the polymer may be selected to enhance small end opening of the device, but is not necessarily required. In another embodiment, a valve of this structure can be implanted in the prostate to treat incontinence. This valve prevents urine from flowing out until sufficient pressure is generated. The valve can be configured so that the valve opens by the patient relaxing muscles in the ureter that prevent urine flow.
Another important feature of the present invention is that the knitted medical device is heat treated to reduce the stress caused by the knitting process, thereby making the device, eg, stent, more elastic and long lasting. As shown in FIGS. 1d-1h, adjacent rows of knit loops have overlapping contact portions. During the knitting process to form such a structure, the metal filaments are bent, thereby introducing stresses that give some stiffness within the stent. At the overlap, this stiffness prevents movement of the row adjacent to one row of knit loops, thus reducing stent compliance and reducing the ability to adapt to the physiological movement of the lumen wall in which the stent is placed. Let Furthermore, this stiffness can increase the friction in the overlap region, thereby reducing the life of the stent. These concerns can be mitigated by a suitable heat treatment that reduces the stress created in the overlap region by knitting and increases elasticity. This stress can also be mitigated by a combined choice of knitting wire diameter and knit loop bend amount. In general, the smaller the knitting wire and the larger the loop size, the less the rigid stress of the resulting knit device. The heat treatment can also vary with these parameters.
The heat treatment not only reduces the stress in the overlap region, but also restores the tensile strength of the stent wire. The heat treatment conditions depend on the degree of work hardening that is induced during knitting and weakens the wire. The greater the bend, the greater the work hardening. For example, a wire having a tensile strength of 250,000-300,000 psi has a tensile strength of 70-90,000 psi after being knitted into an esophageal stent. After the heat treatment, the tensile strength is restored to, for example, 180 to 190,000 psi. The heat treatment that reduces the stress in the overlap region and restores the tensile strength can be performed simultaneously with the heat treatment for shaping the stent. A heat treatment that reduces the stress in the overlap region and restores the tensile strength is useful even when the stent is not shaped.
The performance of a knitted medical device can be improved by making it on a knitting machine structured to limit the friction of the knitting wire during knitting. Referring to FIG. 11, the knitted stent can be manufactured by a conventional circular knitting machine 1100 very similar to that used to knit stockings. The knitting machine includes a knitting head 1102 for guiding a series of needles 1104 extending or contracting axially by a rotating (in the direction of arrow 1106) platen 1108.
The portion of the knitting machine that contacts the knitting wire is made of a low friction durable polymer to reduce friction. Usually, the iron or steel knitting head 1102 is preferably made of a low-friction durable member such as nylon or delrin (polyacetate) for application in medical devices, as these are pulled into the knitting head. Reduces stent wire friction, scratching and scoring. This will be described in detail below. Needle head 1120 may also be formed or coated with such a polymer.
During the knitting operation, wire 1110 is fed from spool 1112 to the needle. A knitted stent 100 is manufactured around the polymer mandrel 1114 and pulled down in the direction of arrow 1116. With particular reference to FIGS. 11 a-11 d, each needle 1104 includes a needle head 1120 and a pivoting needle tongue 1122. During the needle up stroke (FIG. 11a), the head 1120 grabs the wire 1110 and the tongue 1122 is initially in the down position. In the down stroke (FIG. 11b), the tongue 1122 is deflected upward and engages a portion of the knitting head 1102, thus enclosing one strand within the head. The descending stroke (FIG. 11c) continues for a selected period in the knitting head, with the deformed wire 1110 passing its elastic limit and forming a loop 610 in the wire. In the up stroke (FIG. 11d), this strand deflects the tongue 1122 downward, thereby releasing the strand from the needle head 1120. As the stent 100 is pulled down (arrow 1116), the loop 610 is pulled and finally upwards. The platen 1108 rotates and this cycle is repeated. As shown in FIG. 11e, it is preferable to make the knit with a different form of “up-loop” 610 rather than “down-loop” 612 on certain knitting machines or for a stent of a certain diameter. In certain applications, for example in the case of the aorta, it is important that the loop is uniform so that the stent exerts a uniform pressure on the lumen wall. Since the wire is under tension during the knitting process, the loop is in a tensioned shape similar to FIG. 1e, or FIG. 1f or 1h, depending on the shape and configuration parameters of the knitting machine.
Referring to FIGS. 11 e and 11 f, many of the geometric parameters of the knitted stent depend on the shape of the knitting head 1102. The knitting head 1102 is generally a truncated cone, with a number of slots 1160 in the central through hole and side. The number of circumferential loops 610 of the stent depends on the number of needles 1104 that move through the slot 1160 of the knitting head. The overall diameter D of the stent is approximately equal to the diameter 1162 of the through hole of the knitting head. The width 1154 of the up loop 610 is related to the diameter of the needle 1104 and the width of the knitting head slot 1160: A narrow needle in a narrow slot narrows the loop as the strand bends around the narrow needle head at an acute angle. Form. The width 1156 of the down loop 612 is related to the land width 1164 between the knitting head slots 1160. For example, the amount of stent compression to compress the stent from the working stationary diameter to the diameter required to enter the small delivery system is most strongly influenced by the dimension 1156, which is the width of the down loop 612. For example, by reducing the number of circumferential loops and increasing the dimension 1156, the stent can be made more compressible while maintaining its diameter. As shown in FIG. 11g, when such a stent is compressed, as shown in region 1170, the down loop 612 bends and the shoulders of the loop overlap.
The knitting machine produces “ropes” with long knit loops. The rope is cut somewhat longer than the final length of the stent. This excess length allows for the shortening of the stent that occurs when the loop as it exits the knitting machine is shortened to the resting state of FIG. 1b, and also allows some trimming. The knitted tube having a constant diameter is then stress relieved and shaped as described above.
Thus, according to the methods and suggestions above, in a preferred embodiment, this wire is 0.002-0.010 inches in diameter, eg from Shape Memory Applications, Sunnyvale, CA. Available, this is equal to -5 to + 10 ° C from 1/4 "nickel / titanium alloy wire available by Fort Wayne Metals, for example from Furukawa Electric Company, Japan A_fIt was stretched by. After stretching, the wire has a tensile strength of about 250,000 psi to 300,000 psi. Tubes knitted for esophageal stents are heat treated in a vacuum at 400 ° C. for 20 minutes, then cooled to 100 ° C. with a stream of nitrogen for 1 minute and then cooled to room temperature over 20 minutes. The heat treatment may be changed as described above. For example, in the case of a biliary stent, the maximum temperature is 450 ° C. For prostate stents, the maximum temperature is 500 ° C. for 30 minutes. Other materials include high modulus stainless steel alloys such as alloys containing molybdenum and cobalt.
As mentioned above, certain applications have preferred standards and dimensions. Preferred dimensions are shown in Table 1. For example, a preferred esophageal stent is formed of 0.006 inch diameter wire, has an overall diameter D of 18 mm, has a 20-22 mm flare 622 at one end, and has 16 loops around its circumference. This flare is preferably 20 mm in diameter and 1.5 cm in length. Bile duct stents have no flare on either end. The colonic stent has a diameter of about 10-20 mm, a length of about 4-10 cm, and flares at both ends. Vascular stents are a special embodiment. The wall of the blood vessel slightly expands and contracts with each heart beat, but does not substantially extend radially inward. When a blood vessel is on the limb, especially near the joint, the length and bend of the blood vessel undergo significant changes. A vascular stent would be about 6 mm in diameter. The radial force of the stent must be strong enough to keep the lumen open and not block blood flow. The wire diameter and composition are selected to have high fatigue resistance and meet the above requirements.

Other embodiments are within the scope of the following claims.

Claims (5)

A method of manufacturing a system for delivering an endoprosthetic device to a body by reducing the endoprosthetic device from a static diameter to a substantially reduced diameter comprising:
An endoprosthetic device is a stent for a body lumen that exhibits a physiological movement, such as a peristaltic movement,
The method is
Providing a confinement block having a hole at least as large as the reduced diameter, having a slot on a side of the confinement block, the slot terminating at one end in the hole;
Sandwiching the endoprosthetic device to form a flat portion of the endoprosthetic device;
Inserting a flat portion of the endoprosthetic device into the containment block, the flat portion of the endoprosthetic device entering the slot, and an adjacent portion of the endoprosthetic device entering the aperture;
Inserting a first mandrel into a portion of the endoprosthetic device within the hole;
Inserting a second mandrel into the hole but outside the endoprosthesis device;
Rolling the mandrel relative to each other to wind the endoprosthetic device until the endoprosthetic device is fully wound and has a reduced diameter in the hole;
Removing the endoprosthetic device from the hole while constraining the endoprosthetic device to its reduced diameter state;
Comprising a method. The method of claim 1, wherein the removing step slowly pushes the endoprosthetic device from the distal end of the hole and restrains the endoprosthetic device when the endoprosthetic device comes out. The method of claim 2, wherein the constraint comprises winding a wire around the endoprosthetic device. The method of claim 1, wherein the slot is tangent to the hole of the containment block. The method of claim 4, wherein the first mandrel includes an elongated delivery carrier for the endoprosthetic device.

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