DE102016110410A1 - Stent and manufacturing process - Google Patents

Abstract

The present invention relates to a stent (10) for implantation in a, in particular neurovascular, blood vessel having a substantially tubular, self - expandable lattice structure (11), which is convertible from a radially fully expanded expansion diameter to a radially fully compressed compression diameter, wherein the Lattice structure (11) exerts a hysteresis with a compression pressure RRF and an expansion pressure COF, where for the compression pressure RRF and the expansion pressure COF in a diameter range from x1 to x2: [(RRF (x1) - RRF (x2)) - ( COF (x1) - COF (x2))] / [COF (x1) - COF (x2)] ≥ Awhen A = 10% and x2> x1.

Description

The invention relates to a stent for implantation in a, in particular neurovascular, blood vessel. Furthermore, the invention relates to a manufacturing method for such a stent.

Stents are well known and commonly used to treat stenoses in blood vessels. The purpose of the stent is to dilate the narrowed in the area of the stenosis blood vessel, so as to provide adequate blood flow. The placement of the stent is usually done by feeding the stent through a catheter to the treatment site. During delivery through the catheter, the stent having a tubular lattice structure is in a radially compressed state. At the treatment site, the stent is released from the catheter. Self-expanding stents expand spontaneously upon discharge from the catheter and exert a radial force on the vessel wall. Thus, the flow-through cross section of the blood vessel is maintained.

In order to expand a stenosis and thus largely restore the original, flow-through cross-sectional diameter of a blood vessel, a balloon catheter is often used in practice. The balloon catheter comprises an inflatable balloon which is guided into the region of the stenosis and widened there. This will dilate the stenosis. The expansion of the stenosis by the balloon is for medical reasons usually not up to the complete original vessel diameter, i. to the diameter of the blood vessel in healthy condition. A partial dilation is preferred so that the vascular tissue is not overstretched and thus injured.

An injury can lead to the removal of endothelial tissue. As a result, blood flowing through the blood vessel may come into contact with subendothelial tissue. The subendothelial tissue is highly thrombogenic and increases the risk of blood clots.

Nevertheless, in the medium and long term, a complete dilation of the blood vessel in the area of the stenosis up to the original vessel diameter is desired. However, dilation should be slow so that the tissue of the blood vessel can adapt to dilation, thus reducing the risk of injury. It is therefore desirable to have a stent which causes a continuous dilation of the blood vessel continuously and slowly, in particular over a period of several days.

For the operation of a stent in the blood vessel thus a radial force is required, which ensures that the stent exerts a pressure radially outward. This radial force is also problematic for the delivery of the stent, as the stent urges against the catheter inner wall. This increases the frictional forces when pushing the stent through the catheter.

This is problematic because a user has to apply a relatively high force to overcome the stiction between the stent and the catheter, which makes exact placement of the stent in the blood vessel difficult.

It is also known to preload self-expanding stents directly on a catheter tip, so that the stent does not have to be pushed through the entire catheter. Although this solves the problem of increased static friction during advancement of the stent through the catheter. At the same time, however, it is accepted that the catheter tip is stiffened by the stent placed therein. This, in turn, hinders navigation through curved sections of blood vessels that are common, especially in neurovascular blood vessels.

The object of the invention is to provide a stent which can easily be passed through a catheter and at the same time effectively dilates a stenosis. It is another object of the invention to provide a manufacturing method for such a stent.

According to the invention this object is achieved with regard to the stent by the subject-matter of patent claim 1 and with regard to the manufacturing method by the subject-matter of patent claim 18.

Specifically, the invention is based on the idea of specifying a stent for implantation in a, in particular neurovascular, blood vessel with a substantially tubular, self-expandable lattice structure. The grid structure can be converted from a radially completely expanded expansion diameter to a radially completely compressed compression diameter. In this case, the radial pressure exerted by the grid structure forms a hysteresis with a compression pressure RRF and an expansion pressure COF. For the compression pressure RRF and the expansion pressure COF applies in a diameter range from x ₁ to x ₂ : [(RRF (x ₁ ) - RRF (x ₂ )) - (COF (x ₁ ) - COF (x ₂ ))] / [COF (x ₁ ) - COF (x ₂ )] ≥ A

Furthermore, A = 10% and x ₂ > x ₁ .

Self-expanding stents have the property of being different when transitioning from the compressed state to the expanded state Have radial forces, as in the transition from the expanded state to the compressed state. In addition, the radial force or the pressure exerted radially by the grid structure differs for each cross-sectional diameter of the grid structure.

It has been found that a self-expanding grating structure provides a greater resistance to a radial compression action than it exerts when it radially expands. In other words, the compression pressure required to compress the self-expanding stent is greater than the expansion pressure that the mesh structure automatically exerts upon expansion.

The invention proposes to set the grating structure such that, in a predetermined diameter range, the difference between the increase in the compression pressure and the increase in the expansion pressure is at least 0.1 times the increase of the expansion pressure.

In other words, it is intended to contrast the increase of the pressure, which manifests itself in the compression of the lattice structure, with the reduction of the pressure, which manifests itself in the expansion of the lattice structure. The expansion pressure decreases upon expansion of the lattice structure, the decrease being less than the increase in compression pressure in the compression of the lattice structure.

Surprisingly, it has been shown that when adjusting the compression pressure and the expansion pressure according to the above-mentioned formula, a particularly good compromise is achieved between the ability to deliver the stent through a catheter and the therapeutic effectiveness of the stent for dilating a stenosis. It is believed that the therapeutic efficacy is improved because the stenosis attempts to compress the stent, with the stent counteracting this compression with the relatively higher compression pressure. When the stent is delivered, the lattice structure automatically pushes outward, so that the expander pressure acts on the inner surface of the catheter, which is lower than the compression pressure.

In the invention it is provided that the compression pressure RRF is not constantly greater than the expansion pressure COF, but varies depending on the actual cross-sectional diameter of the grid structure. The smaller the cross-sectional diameter of the lattice structure, the higher the difference between the compression pressure RRF and the expansion pressure COF.

As a result, the grid structure exerts a relatively low but constant force, namely the COF expansion pressure, on a stenosis. The expansion pressure decreases over time as the grid structure slowly expands, increasing the cross-sectional diameter of the grid structure. In this way it is ensured that the stenosis is dilated continuously and vascular tissue-friendly over a long period, in particular of several days.

The different course of the compression pressure RRF and the expansion pressure COF thus has particular advantages with respect to the behavior of the stent in the blood vessel. A high difference between the compression pressure RRF and the expansion pressure COF leads to a relatively rigid behavior of the stent. The stent thus resists compression by the blood vessel, but acts only moderately on the blood vessel due to the relatively lower expansion pressure itself. Thus, the stent prevents on the one hand a further narrowing of the blood vessel and on the other hand simultaneously causes a moderate and slow expansion. In the area of a stenosis, this is particularly advantageous because too rapid expansion of the stenosis can lead to injuries of the vascular tissue.

This advantage of the invention is particularly evident in the treatment of neurovascular blood vessels in the intracranial region. Such blood vessels have relatively many vascular discharges, so-called perforators. In the treatment of stenoses by stents, these perforators emerging from the main vessel are particularly critical areas. In particular, there is a risk that parts of the vascular area affected by a stenosis will break up and be pushed into the perforators. This can lead to further vascular occlusions.

By the lattice structure continuously acting on the stenosis due to the low expansion pressure COF, the blood vessel is slowly expanded in this area. This avoids rupture of the affected by the stenosis vascular area and thus reduces the risk of further vascular occlusions.

Although the invention sets high value on a comparatively high compression pressure RRF of the lattice structure with simultaneously a relatively low or moderate expansion pressure COF. However, a small difference between the compression pressure RRF and the expansion pressure COF promotes compliance of the stent in the blood vessel. Such a small difference is achieved with comparatively larger cross-sectional diameters of the lattice structure, since the compression pressure RRF and the expansion pressure COF converge with increasing cross-sectional diameter of the lattice structure.

Particularly in the pulsatile cycle of the blood flow, a small difference between the compression pressure RRF and the expansion pressure COF is advantageous because the force of the stent is similar in systole (stent expansion) and in diastole (stent compression). The vascular wall of the blood vessel is thus subjected to a similar pressure during expansion in systole and in compression in diastole. This pressure adds up to the dynamic pressure of the bloodstream, which varies between systole and diastole. Thus, the smaller the difference between the compression pressure RRF and the expansion pressure COF, the less the difference in the load on the vessel wall between systole and diastole. The physiological, pulstile stretching of the vessel is so little affected.

The aforementioned behavior is particularly advantageous in peripheral areas of the stent, which have direct contact with the vessel wall in the implanted state. The good compliance behavior of the stent prevents the pulse wave from being interrupted at the edge of the stent, which can otherwise lead to vascular injuries and restenoses.

Particularly good results can be achieved if the value A is one of the following values in the above formula. In particular, it is preferred if A = 15%, in particular A = 20%, in particular A = 25%, in particular A = 30%, in particular A = 35%, in particular A = 40%.

The advantages described above are particularly evident in the diameter range between 1.5 mm and 2.5 mm. In this respect, it is preferred if x ₂ = 2.5 mm and x ₁ = 1.5 mm.

In general, it is advantageous in the invention if the relatively larger cross-sectional diameter x2 corresponds to the expansion diameter of the grid structure minus 0.5 mm. The relatively smaller cross-sectional diameter x ₁ is preferably smaller by 1 mm than the relatively larger cross-sectional diameter x ₂ . Mathematically expressed preferably: x ₂ = d _exp - 0.5 mm and x ₁ = x ₂ - 1 mm, where d _exp denotes the expansion diameter of the lattice structure.

For a stent with an expansion diameter of 4.0 mm this results in: x ₂ = 4.0 mm - 0.5 mm = 3.5 mm x ₁ = 3.5mm - 1.0mm = 2.5mm

The lattice structure of the stent may have an implantation diameter and an expansion diameter. The implantation diameter is also referred to as Intended Use diameter. This diameter indicates the preferred value for the diameter of the blood vessel in which the stent is to be inserted. For the compromise between compression pressure and expansion pressure described above, it has proved to be advantageous if the implantation diameter of the lattice structure is less than 3.5 mm, in particular between 3.4 mm and 1.0 mm.

The expansion diameter of the lattice structure is present when the lattice structure is fully expanded. The expansion diameter thus designates the unenergized state of the lattice structure. Preferably, the expansion diameter of the grid structure in the present invention is more than 2.5 mm and at most 4.0 mm. In particular, it can be provided that the expansion diameter is at least 3.0 mm and at most 3.5 mm. Concretely, the expansion diameter of the lattice structure can be 3.0 mm.

The grid structure preferably has a length between 8 mm and 35 mm, in particular between 8 mm and 20 mm, in particular between 10 mm and 15 mm.

In this context, it should be noted that the radial pressure described here acts on the environment via the outer contact surface of the lattice structure. The outer contact surface of the lattice structure is determined by the outer surfaces of the structural elements, in particular webs or wires, of the lattice structure. The sum of the outer surfaces of all structural elements of the lattice structure essentially forms the outer contact surface of the lattice structure with the environment, for example a vessel wall.

The radial pressure exerted on the outer contact surface of the grid structure can be measured with a device which essentially has a type of iris diaphragm. The stent is inserted into the iris diaphragm, with the iris diaphragm coupled to a pressure cell or strain gauge. The iris diaphragm absorbs the pressure of the lattice structure over the entire circumference of the lattice structure, so that the corresponding radial pressure can be measured with the pressure cell connected to the iris diaphragm.

The outer contact surface of the stent is preferably between 7.5 mm ² and 16.75 mm ² . For a stent with an implantation diameter of 3.0 mm and a length between 15 mm and 35 mm, the contact area is preferably between 8, 2 mm ² and 18.7 mm ² . A stent with an implantation diameter of 4.0 mm and a length between 15 mm and 35 mm may have a contact area between 7.5 mm ² and 16.7 mm ² .

For a stent with an implantation diameter between 4.5 mm and 5.5 mm and a length between 15 mm and 35 mm, the outer contact surface is preferably between 9.5 mm ² and 24.6 mm ² .

In particular, for the treatment of stenoses in neurovascular blood vessels, such as blood vessels in the intracranial region of the head, it has proved to be advantageous if the compression pressure RRF at an expansion diameter between 1.0 mm and 3.0 mm, in particular between 1.5 mm and 2 , 5 mm, at least 30 mmHg. It is also possible that the compression pressure RRF at the aforementioned expansion diameters is at least 40 mmHg, in particular at least 50 mmHg, in particular at least 60 mmHg.

The compression pressure RRF may, in particular for the previously described expansion diameter, be at most 150 mmHg, in particular at most 120 mmHg, in particular at most 100 mmHg, in particular at most 80 mmHg.

In the case of self-expansion, the lattice structure of the stent exerts an expansion pressure COF which, with an expansion diameter between 1.0 mm and 3.5 mm, in particular between 1.5 mm and 2.5 mm, can be at least 10 mmHg and at most 70 mmHg. It is preferred if the expansion pressure COF at the aforementioned diameter values is between 20 mmHg and 60 mmHg, in particular between 30 mmHg and 50 mmHg.

The lattice structure of the stent may comprise or consist of a nickel-titanium alloy, in particular nitinol. Such nickel-titanium alloys are suitable for the production of self-expanding stents and have corresponding material properties, which together with further optimization measures leads to the effects desired according to the invention.

In order to provide the desired effect by means of the material properties of the lattice structure, namely the improved ability to be supplied while at the same time having good therapeutic effectiveness of the stent, the nickel-titanium alloy can have between 50 atom% and 60 atom% nickel. It is particularly preferred if the nickel-titanium alloy has between 55 atomic% and 56 atomic%, in particular concretely 50.8 atomic%, of nickel.

The grid structure is formed in one piece from webs in particularly preferred embodiments. For example, the lattice structure may be made in one piece by laser cutting a tube or by sandwiching by a deposition method, in particular a vacuum deposition method, for example, a physical vacuum deposition (PVD) method or a sputtering method. However, it is conceivable that the grid structure is formed by a mesh of wires.

In order to allow delivery of the stent into small blood vessels, in particular by means of a correspondingly thin catheter, the lattice structure can be compressible to a compression diameter of 0.4 mm.

Moreover, it is preferred if the grid structure has closed cells, in particular a closed-cell design. The closed cells facilitate delivery of the stent through a catheter. In particular, the webs of closed cells mutually support each other so that the stent can be easily moved as a unit through a passage lumen of the catheter.

A closed-cell design distinguishes between an open-cell design, which can lead to distortion of the lattice structure when released. In particular, in an open-cell design freestanding bar connectors can lead to additional irritation or injury to the tissue. By contrast, a closed-cell design has advantages in that it has no freestanding bar connectors and is usually not distorted when released from a catheter.

Further, a grid structure with a closed cell design may be withdrawn back into the catheter after partial discharge from the catheter. This may be helpful to correct mispositioning of the stent.

Advantageously, it is provided in the invention that a central portion of the lattice structure in the implanted state causes a sufficient expansion of a stenosis and edge regions of the lattice structure generate a radial pressure sufficient to hold the lattice structure in the blood vessel without causing injury to the vascular tissue.

The grid structure is usually placed in the blood vessel so that a central portion of the grid structure covers the stenosis. The edge sections or edge regions of the lattice structure, on the other hand, are located in healthy vessel sections which adjoin the stenosis. In the area of the healthy vessel sections, ie in the peripheral areas, the lattice structure is in the implanted state a larger cross-sectional diameter than in the middle section, whose cross-sectional diameter is limited by the stenosis.

Because of the relationship between the compression pressure RRF and the expansion pressure COF described in detail hereinabove with respect to the cross-sectional diameter, the grid structure is more stable in the central portion than in the peripheral portions. Therefore, the lattice structure may be subject to deformation of the blood vessel by the pulsatile blood flow, i. a pulse wave, follow better in the edge regions than in the middle section.

Specifically, it may be provided that the edge regions of the lattice structure avoid vascular injury or reduce the effect thereof that the amplitude of the pulse wave deformation of the blood vessel in the edge regions by at most 10%, in particular at most 20%, in particular at most 30%, in particular at most 40%, in particular at most 50%, based on the state without stent inserted is reduced.

At the same time, the radial pressure of the middle section of the grid structure causes the stenosis to expand. This expansion is preferably very slow to prevent separation of stenosis particles. Otherwise, detached stenosis particles could be flushed into smaller blood vessels and cause occlusion there.

The distention of the stenosis through the central portion of the lattice structure is preferably so slow that for dilation of the diameter of the blood vessel to 110% of the original diameter value, i. before implantation of the stent, a period of at least 12 hours is required. The dilation of the blood vessel to a diameter value which is increased by 10% compared to the original diameter may also be at least one day, in particular at least 2 days, in particular at least 3 days, in particular at least 1 week, in particular at least 2 weeks, in particular at least 1 month. last for.

It is also possible, particularly when using a lattice structure which provides an increased COF expansion pressure, that the stenosis is distended to a diameter exceeding the original diameter by 20% within the aforementioned periods. In any case, it is envisaged that distension of the stenosis up to a diameter of 110% of the original diameter will take no longer than 6 months.

In particular, the middle section of the lattice structure can expand the stenosis so slowly that branching perforator vessels in the area of the stenosis are not closed by the expansion of the stenosis. In particular, an expansion of 10% or 20% of the original diameter may occur in the aforementioned periods. This ensures that the perforator vessels have sufficient time for adaptation to the changed cross-sectional diameter of the main vessel.

The adaptation of the Perforatrengefäße to the changing cross-sectional diameter of the main vessel is referred to as "remodeling". An adequate remodeling ensures a sufficient blood flow into the perforator vessels, which is adapted to the (energy) needs of the following tissue. The remodeling takes several hours, sometimes several days or weeks. When a stenosis spreads too quickly, the tissue of the stenosis, called plaque tissue, slowly bulges into the perforator vessels and may result in a flow restriction or occlusion. The slow expansion of the stenosis avoids this.

A secondary aspect of the invention relates to a stent for implantation in a, in particular neurovascular, blood vessel having a substantially tubular, self-expandable grid structure, which can be converted from a radially fully expanded expansion diameter to a radially fully compressed compression diameter. The stent may in particular be formed according to the preceding embodiments. It is provided in the context of the independent aspect of the invention that the grid structure has at least one peripheral cell with a first length and at least one central cell with a second length. The ratio between the first length and the second length is preferably between 0.5 and 2.5, in particular between 1.0 and 2.0, in particular between 1.25 and 1.75. In a specific embodiment it can be provided that the ratio between the first length and the second length is 1.5.

The length of a cell is usually determined between two connection points or intersection points of the webs, which are arranged in the longitudinal direction of the lattice structure immediately adjacent to each other. The length thus runs parallel to the longitudinal axis of the lattice structure. Edge cells are cells which are located at the longitudinal axial edge of the lattice structure, thus closing off the lattice structure longitudinally.

The setting of different cell lengths between at least one edge cell and at least one middle cell has several advantages. On the one hand, this achieves the result that comparatively low radial forces are established in the edge area, which improves the ability to feed the stent through a catheter. On the other hand, it can be provided that each second peripheral cell in the circumferential direction has the first length. In contrast, the edge cell arranged in each case can comprise a different length. Thus, it is possible that the edge-side tips of the edge cells are offset from one another, which provides for an improved compressibility of the grid structure.

It is particularly preferred if the marginal cell with the first length has an extension which carries an x-ray marker. Particular advantages arise in particular when alternate in the circumferential direction of the grid structure each edge cells with an extension and edge cells without an extension. Since the x-ray marker, which is arranged on the extension, increases the wall thickness and web width of the lattice structure, an improved compression can be achieved by staggered arrangement of the edge-side tips of the edge cells. This in turn contributes to the fact that the stent can be easily passed through a catheter to the treatment site despite sufficiently high therapy efficiency.

The edge cell of the first length and the center cell of the second length may be arranged in the longitudinal direction of the grid structure immediately adjacent to each other. Alternatively or additionally, it is possible for the edge cell to be coupled to the first cell and the middle cell for the second length to be coupled together by a common cell connector.

Another subsidiary aspect of the invention relates to a manufacturing method for a stent described above. In the production method according to the invention, the lattice structure is preferably deposited on a substrate by a physical deposition method, in particular a sputtering method, and then removed from the substrate.

The lattice structure can be deposited both in a tubular form, as well as deposited first planar and then converted into a tubular form. In the first planar formation of the lattice structure is preferably provided that the shaping of the final tube shape at elevated temperatures between 450 ° C and 680 ° C takes place.

It can preferably be provided that the lattice structure is formed by depositing a supporting material and a radiopaque material other than the supporting material. In other words, the lattice structure may be formed of two different materials, with a base layer and a surface layer being provided, and the surface layer having radiopaque properties.

It is also conceivable that the radiopaque material is applied in regions or at points on the supporting material, in particular so as not to affect the mechanical properties of the supporting material. Preferably, however, the X-ray-visible material is applied distributed along the lattice structure such that at least the longitudinal-axial ends of the lattice structure are easily recognizable under X-ray control.

In particular, the supporting material may comprise or consist of a nickel-titanium alloy. The nickel-titanium alloy may contain a predetermined amount of nickel as described above. The radiopaque material preferably comprises or consists of tantalum and / or platinum and / or gold. Such materials are particularly suitable for the coating of a nickel-titanium alloy to improve the radiopaque properties.

The invention will be explained below with reference to an embodiment with reference to the accompanying schematic diagrams. Show:

1 FIG. 4: an unfolded view of a lattice structure of the stent according to the invention according to a preferred embodiment; FIG.

2 FIG. 4 is a graph showing the compression pressure and the expansion pressure in the stent according to FIG 1 ; and

3 a tabular overview of different peripheral surfaces in the stent according to 1 ,

1 shows a stent in an unfolded view 10 or its lattice structure 11 , The grid structure 11 is integrally formed and has a plurality of webs 13 on. The bridges 13 limit cells 12 , Specifically, it is envisaged that the lattice structure 11 is designed as a closed grid structure (closed cell design). In particular, each limit four bars 13 a cell 12 , The cells 12 have essentially a dragon-shaped or diamond-shaped basic geometry.

The individual bars are on cell connectors 15 connected with each other. The cell connectors 15 are preferably aligned in the longitudinal direction of the stent (in the image plane from left to right) in alignment with each other.

In 1 It can be seen that at the axial longitudinal ends of the lattice structure 11 boundary cells 12a are arranged, which have a first length l ₁ . The marginal cells 12a are through a cell connector 15 with middle cells 12b coupled. The middle cells 12b have a second length l ₂ .

Specifically, it is envisaged that the middle cells 12b each with four equally long bars 13 are limited. The marginal cells 12a On the other hand, webs of different lengths can be used 13 exhibit. In particular, it can be seen that at the axial longitudinal ends of the lattice structure 11 several border cells 12a are arranged adjacent to each other in the circumferential direction, each second peripheral cell 12a two V-shaped mutually arranged, longer webs 13 having. The longer bridges 13 lead into an extension 14 which is parallel to the longitudinal axis of the lattice structure 11 over the marginal cells 12a extends beyond. The extension 14 is intended to record an x-ray marker. For example, an X-ray marker sleeve on the extension 14 attached and fixed by crimping.

In 1 is clearly the difference in length between the first length l ₁ and the second length l ₂ characterized. The first length l _{1 of} the boundary cell 12a and the second length l _{2 of} the middle cell 12b be between the in the longitudinal direction of the lattice structure 11 immediately adjacent and aligned cell connectors 15 measured.

Because only every second boundary cell 12a is extended longitudinally axially and in an extension 14 opens, remains through the staggered cell tips 16 a free space for the x-ray markers. In this way, the crimpability of the lattice structure 11 improved. Since the X-ray markers can "dodge" into the spaces between two extended marginal cells, the marginal areas of the grid structure push 11 in the compressed state not disproportionately against an inner peripheral surface of a catheter.

2 shows a diagram in which the cross-sectional diameter of the stent 10 or the grid structure 11 and that of the lattice structure 11 applied radial pressure are compared. Exemplary here are the values for a stent 10 shown, whose implantation diameter is 3.0 mm. The length of the stent 10 or the grid structure 11 is preferably 35 mm.

Shown in 2 the change in radial pressure during expansion of the stent 10 from a cross-sectional diameter of 2 mm to a cross-sectional diameter of 2.5 mm.

The radial pressure in the expansion is called the expansion pressure COF. It can be seen that the expansion pressure COF decreases as the lattice structure expands. The idealized, linear course of the expansion pressure COF between the diameter values 2.0 mm and 2.5 mm has a comparatively flat slope.

Compression of the stent 10 a radial pressure, referred to as the compression pressure RRF. The compression pressure RRF is in 2 also shown in the diameter range between 2 mm and 2.5 mm. It can be seen that the radial pressure of the lattice structure 11 when compressing the stent 10 increases greatly from a cross-sectional diameter of 2.5 mm to a cross-sectional diameter of 2.0 mm. The slope of the idealized linear change in the compression pressure RRF is comparatively large.

It is also possible that the slope of the curves of and / or the expansion pressure COF is not linear. For example, the inclination of the course of the compression pressure RRF (RRF curve) or the slope of the course of the expansion pressure COF (COF curve) may increase with decreasing cross-sectional diameters. In the other curve, the inclination can decrease with decreasing diameter. The difference in the slope of the RRF curve and the COF curve can thus result in relatively small or relatively large cross-sectional diameters of the lattice structure 11 stand out more clearly.

It is also possible that the slope of both curves reduces or increases with decreasing cross-sectional diameters. It is also possible that one curve is nearly linear and the other curve is a slope change, i. a curved course, having. By appropriate adjustment of the slope of the two curves curves, various variations are conceivable, which are advantageous for use in different stenoses.

In particular shows 2 clearly that the expansion pressure COF is basically smaller than the compression pressure RRF. At the same time, it can be seen that the course of the compression pressure RRF between two diameter values has a significantly higher gradient than the profile of the expansion pressure between two diameter values. This is used in the stent according to the invention, on the one hand to achieve good feedability through a catheter. Within the catheter, in particular, the expansion pressure COF is effective. Since this is relatively small, the stent 10 be pushed well through the catheter.

At the treatment site, especially in the area of a stenosis, the stent is compressed by the stenosis. The pressure the stent is holding is the compression pressure RRF, which is relatively high. Thus, the stent is particularly effective for keeping open a blood vessel used. The advantage of the increased compression pressure RRF with a relatively smaller cross-sectional diameter comes into play in particular when the blood vessel to be treated is prone to restenosis. The re-occlusion of the blood vessel leads to a compression of the stent 10 , which counteracts with now increasing compression pressure RRF.

The diagram according to 2 makes something very clear. Because the expansion pressure and the compression pressure are above the current diameter of the stent 10 or the grid structure 11 change, different radial pressures are also in different areas of the stent 10 effective. This is particularly advantageous for a gentle therapy of a stenosis. This is explained below:
In the area of a stenosis, the stent is possibly in a middle section, in particular in the area of the middle cells 12a , more compressed than in its outer portions or edge regions, in particular in the region of the peripheral cells 12a , This already results from it, since the stent 10 is usually used in a length which is greater than the length of the affected by the stenosis blood vessel section. This oversizing of the stent 10 causes the marginal cells 12a outside the stenosis, whereas the middle cells 12b get in direct contact with the stenosis.

As in 2 is clearly recognizable, acts at smaller diameter of the stent 10 a relatively higher compression pressure RRF contrary to a compression of the stent. The middle section of the stent 10 , which is located in the region of the stenosis, is usually more compressed than the marginal areas. Thus, the more compressed stent in the area of the stenosis due to its regionally higher compression pressure RRF can well prevent the blood vessel from further tapering. The middle section of the stent 10 So it works effectively on the stenosis.

In contrast, in the edge sections of the stent, which are arranged outside the stenosis, it is expedient for the stent to be able to follow a movement of the vessel walls. Stenosis-free areas of a blood vessel are very flexible and continuously in motion because of the pulsating blood flow. In particular, the vessel diameter of a blood vessel changes constantly as a function of the pulse beat. In the area of a stenosis, this effect is significantly reduced because the stenosis leads to a stiffening of the blood vessel.

The edge portions of the stent follow a temporary expansion of the blood vessel, for example due to a pulse wave, with the expansion pressure COF, which is relatively low. The expansion pressure COF is in particular lower than in the middle section of the stent, which is more compressed because of the stenosis. At the same time, the compression pressure RRF is relatively lower in the edge sections of the stent than in the middle section. In particular, because of the different slope of the compression pressure RRF and the expansion pressure COF, the difference between these two pressures is smaller for larger cross-sectional diameters than for smaller cross-sectional diameters.

In other words, the larger the cross-sectional diameter in a section of the stent 10 or the grid structure 11 is, the more the compression pressure RRF and the expansion pressure COF are equal to each other. The edge sections of the stent 10 , which are arranged in healthy vessel sections, can follow the movement of the vessel wall due to the pulse waves of the blood well. In the middle section, which has a relatively smaller cross-sectional diameter because of the stenosis-induced constriction, the lattice structure is 11 however, stabilized by the higher difference between the compression pressure RRF and the expansion pressure COF.

Specifically, it will be out 2 and the dependence of the radial pressure of the lattice structure 11 from the current local cross-sectional diameter of the grid structure 11 clearly that the stent 10 has more flexibility in the edge portions than in the more compressed middle portion that supports the stenosis.

This effect is particularly effective when, for the difference between the slope of the curve of the compression pressure RRF and the curve of the expansion pressure COF between two predetermined diameter values x ₁ , x ₂ , this difference is at least 0.1 times the slope of the curve of the expansion pressure equivalent. In mathematical terms, the following applies: [(RRF (x ₁ ) - RRF (x ₂ )) - (COF (x ₁ ) - COF (x ₂ ))] / [COF (x ₁ ) - COF (x ₂ )] ≥ A where A = 10% and x ₂ > x ₁ .

3 shows by way of illustration a table with different area values for the outer peripheral surface of the stent 10 , To calculate the radial pressure in the unit mmHg, the force values recorded in the measuring device are converted via the cylindrical peripheral surface of the stent. The cylindrical peripheral surface or cylindrical surface is idealized.

It is assumed that the stent 10 is a solid cylinder with a diameter corresponding to the implantation diameter. The height of the cylinder corresponds to the length of the stent. This results in the cylindrical surface, which is used as ideal conversion factor (cylindrical surface). In fact, however, the pressure is effective over a contact area, which results from the number of webs. The contact area of the stent corresponds to the cylindrical area minus the cell openings.

For the stent according to 1 , whose radial pressure values in 2 are shown, the cylindrical surface is 288.4 mm ² and the contact surface 16.64 mm ² . The stent has a stent length of 35 mm and an implantation diameter of 3.0 mm.

LIST OF REFERENCE NUMBERS

10

 stent

11

 lattice structure

12

 cell

12a

 edge cell

12b

 central cell

13

 web

14

 extension

15

 cell connector

16

 cells tip

RRF

 compression pressure

COF

expansion pressure

₁

 first length

l ₂

 second length

Claims (24)

Stent ( 10 ) for implantation in a, in particular neurovascular, blood vessel having a substantially tubular, self-expanding lattice structure ( 11 ), which is convertible from a radially fully expanded expansion diameter to a radially fully compressed compression diameter, wherein the of the grid structure ( 11 ) applies a hysteresis with a compression pressure RRF and an expansion pressure COF, and wherein for the compression pressure RRF and the expansion pressure COF in a diameter range of x ₁ to x ₂ : [(RRF (x ₁ ) - RRF (x ₂ )) - (COF (x ₁ ) - COF (x ₂ ))] / [COF (x ₁ ) - COF (x ₂ )] ≥ A where A = 10% and x ₂ > x ₁ . Stent ( 10 ) according to claim 1, characterized in that A = 15%, in particular A = 20%, in particular A = 25%, in particular A = 30%, in particular A = 35%, in particular A = 40%. Stent ( 10 ) according to claim 1 or 2, characterized in that x ₂ = 2.5 mm and x ₁ = 1.5 mm. Stent ( 10 ) according to one of the preceding claims, characterized in that an implantation diameter of the lattice structure ( 11 ) is less than 3.5 mm, in particular between 3.4 mm and 1.0 mm. Stent ( 10 ) according to claim 1, characterized in that the expansion diameter of the grid structure ( 11 ) is greater than 2.5 mm and is at most 4.0 mm, in particular at least 3.0 mm and at most 3.5 mm, in particular 3.0 mm. Stent ( 10 ) according to one of the preceding claims, characterized in that the grid structure ( 11 ) has a length between 8 mm and 35 mm, in particular between 8 mm and 20 mm, in particular between 10 mm and 15 mm. Stent ( 10 ) according to one of the preceding claims, characterized in that the compression pressure RRF at an expansion diameter between 1.0 mm and 3.0 mm, in particular between 1.5 mm and 2.5 mm, at least 30 mmHg, in particular at least 40 mmHg, in particular at least 50 mmHg, in particular at least 60 mmHg. Stent ( 10 ) according to one of the preceding claims, in particular according to claim 7, characterized in that the compression pressure RRF is at most 150 mmHg, in particular at most 120 mmHg, in particular at most 100 mmHg, in particular at most 80 mmHg. Stent ( 10 ) according to one of the preceding claims, characterized in that the grid structure ( 11 ) in the self-expansion an expansion pressure COF exerts, with an expansion diameter between 1.0 mm and 3.5 mm, in particular between 1.5 mm and 2.5 mm, at least 10 mmHg and at most 70 mmHg, in particular between 20 mmHg and 60 mmHg, in particular between 30 mmHg and 50 mmHg. Stent ( 10 ) according to one of the preceding claims, characterized in that the grid structure ( 11 ) comprises or consists of a nickel-titanium alloy, in particular nitinol. Stent ( 10 ) according to claim 10, characterized in that the nickel-titanium alloy between 50 atomic% and 60 atomic%, in particular between 55 atomic% and 56 atomic%, in particular 50.8 atomic%, nickel. Stent ( 10 ) according to one of the preceding claims, characterized in that the grid structure ( 11 ) in one piece from webs ( 13 ) or is formed by a network of wires. Stent ( 10 ) according to one of the preceding claims, characterized in that the grid structure ( 11 ) is compressible to a compression diameter of at most 0.5 mm, in particular at most 0.4 mm. Stent ( 10 ) according to one of the preceding claims, characterized in that the grid structure ( 11 ) closed cells ( 12 ), in particular a closed-cell design. Stent ( 10 ) according to one of the preceding claims, characterized in that a middle section of the grid structure ( 11 ) causes a sufficient expansion of a stenosis in the implanted state and edge regions of the lattice structure ( 11 ) generate a radial pressure sufficient to cause the lattice structure ( 11 ) in the blood vessel without causing injury to the vascular tissue. Stent ( 10 ) according to claim 15, characterized in that the central portion of the lattice structure so slowly expands the stenosis that in the region of the stenosis branching perforator vessels are not closed by the expansion of the stenosis. Stent ( 10 ) for implantation in a, in particular neurovascular, blood vessel having a substantially tubular, self-expanding lattice structure ( 11 ), which can be converted from a radially completely expanded expansion diameter to a radially completely compressed compression diameter, in particular stent ( 10 ) according to one of the preceding claims, wherein the grid structure ( 11 ) at least one boundary cell ( 12a ) having a first length l ₁ and at least one middle cell ( 12b ) having a second length l ₂ , and wherein a ratio between the first length l ₁ and the second length l _{2 is} between 0.5 and 2.5, in particular between 1.0 and 2.0, in particular between 1.25 and 1.75, in particular 1.5. Stent ( 10 ) according to claim 15, characterized in that the marginal cell ( 12a ) an extension ( 14 ) carrying an x-ray marker. Stent ( 10 ) according to claim 15 or 16, characterized in that the peripheral cell ( 12a ) and the middle cell ( 12b ) in the longitudinal direction of the grid structure ( 11 ) are arranged immediately adjacent to each other and / or by a common cell connector ( 15 ) are coupled together. Manufacturing process of a stent ( 10 ) according to one of the preceding claims, in which the grid structure ( 11 ) is deposited on a substrate by a physical deposition process, in particular a sputtering process, and subsequently removed from the substrate. Manufacturing method according to claim 18, characterized in that the lattice structure ( 11 ) is formed by depositing a supporting material and a radiopaque material other than the supporting material. Manufacturing method according to claim 19, characterized in that the supporting material comprises or consists of a nickel-titanium alloy, in particular according to claim 11, and in that the radiopaque material comprises or consists of tantalum and / or platinum and / or gold. Treatment system, in particular for the treatment of stenoses in intracranial blood vessels, with a balloon catheter and a stent ( 10 ) according to one of claims 1 to 18, wherein the balloon catheter comprises a catheter tube carrying an inflatable balloon, and wherein the catheter tube has a passage lumen in which the stent is longitudinally displaceable. Treatment system according to claim 21, characterized in that the passage lumen has an inner diameter of 0.4 mm.

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