JP6765248B2 - Manufacturing method of tubular medical device transfer device - Google Patents

Description

The present invention relates to a method for manufacturing a device for transporting a tubular medical device into a body.

In recent years, minimally invasive treatment techniques have been developed in which tubular medical devices are transported and placed in lesions in the body. In this technique, a transport device for transporting and arranging a tubular medical device to a lesion through a lumen in the body is used. The transport device includes a tube, and the tubular medical device is transported to the lesion through a lumen in the body while the tubular medical device is held in the tube. After transport, the tubular medical device can be placed in the lesion by releasing the tubular medical device from the lumen of the transport device.

As the tubular medical device, for example, a stent, a stent graft, an occlusion device, an injection catheter, a prosthesis valve and the like are used. Of these, stents are generally tubular medical devices that treat a variety of diseases caused by the narrowing or obstruction of blood vessels or other in vivo lumens. More specifically, it is a tubular medical device capable of expanding the lumen of a lesion and maintaining the lumen diameter at a predetermined size by arranging a stent in a lesion such as a stenosis site or an occlusion site. Stents are classified into balloon-expandable stents, which are expanded by a balloon on which the stent is mounted, and self-expandable stents, which expand themselves by removing an external member that suppresses the expansion of the stent.

For example, Patent Documents 1 to 4 are known as a method of loading a stent into the lumen of a transport device.

Of these, Patent Document 1 provides a step of providing a loading device having a passage gradually narrowing from a larger first diameter to a smaller second diameter as a method of loading a stent into a catheter. A step of making a stent wrapped in a sleeve by arranging a flexible sleeve over the stent, and on the sleeve so that one end of the sleeve can be gripped from the smaller diametrically end of the passage. A step of placing the wrapped stent in the tapered portion of the passage in the loading device and a step of placing the catheter at or within the smaller diametrical end of the passage. A method comprising pulling the sleeve through the smaller diametrically end of the passage so that the stent is reduced in diameter and moved into the catheter is described. According to this method, the frictional (shearing) force acting on the wall of the stent during the process of loading the stent into the catheter can be minimized.

Patent Document 2 has an inlet having a relatively large cross-sectional area and an outlet having a relatively small cross-sectional area as a method of compressing the self-expandable stent in the radial direction and loading the stent into the supply sleeve. A step of providing a compressor in which the region between the inlet and the outlet is configured to gradually define a smaller cross-sectional area, a step of supporting the open end of the sleeve in alignment with the outlet, and the said. While the relative movement between the stent and the compression device and the sleeve is performed to advance the stent through the compression device into the sleeve and through the sleeve, the relative movement is performed. Describes a method comprising detecting a force applied to the stent and terminating the relative movement when the force exceeds a predetermined level. According to this method, the step of mounting the stent in the feed sheath or sleeve can be more automated, and strict quality control and process improvement can be facilitated.

Patent Document 3 describes a method of loading a naturally expandable stent into a feed sheath in which the stent is held in a radialally compressed pre-deployment form, i) compressing the stent from a relaxed outer diameter d0. A step of radially compressing to the outer diameter d1 and a step of providing a loading sheath having an inner diameter d2 such that ii) d0> d2> d1 and iii) moving the stent in parallel with the sheath. A step of allowing the stent to be received into the lumen of the loading sheath, a step of providing a feed sheath having an inner diameter d3 such that iv) d0> d3> d2, and v) the loading holding the stent. A method comprising a step of advancing the sheath into the lumen of the feed sheath and vi) deploying the stent into the lumen of the feed sheath is described. According to this method, it is possible to improve the stent feeding system which is compatible with the simplicity of design and the reliability of operation and can realize a smaller passing contour outer shape.

Patent Document 4 describes a method of loading a stent into a delivery tube. In this method, the stent is loaded into the opening of the crimping device, the stent has an initial diameter, and the stent is crimped through the crimping device so that the diameter of the stent is subtracted from the initial diameter. The stent is ejected from the opening of the crimp device into the passage of the delivery tube while vibrating the diameter of the opening of the delivery tube, and the passage of the delivery tube has a diameter smaller than the initial diameter of the stent. There is. According to this method, the force required to push or pull the compressed stent into the delivery tube can be reduced.

Special Table 2002-525167 Special Table 2003-530143 Special Table 2012-500075 Special Table 2014-523294

In order to further develop the minimally invasive treatment, it is important to further reduce the diameter of the tube for holding the tubular medical device provided in the tubular medical device transport device to improve operability and passability. On the other hand, in order to dilate the lumen of the lesion and maintain the lumen diameter, the tubular medical device is required to have a large radial dilation force (hereinafter, may be referred to as radial force). However, increasing the radial force of the tubular medical device increases the radial force of the tubular medical device that has been reduced in diameter for loading into the lumen of the tubular medical device carrier. Therefore, it becomes difficult to load the tubular medical device into the lumen of the transport device, which may cause damage to the tubular medical device and the transport device.

The present invention has been made by paying attention to the above circumstances, and an object of the present invention is to reduce the diameter of a tubular medical device transport device in order to load it into the lumen provided in the transport device. It is an object of the present invention to provide a method for manufacturing a tubular medical device transporting device, which can easily load the provided tubular medical device into the lumen of the transporting device.

Method of manufacturing a tubular medical device conveying apparatus according to the present invention that has solved the above problems is a method for manufacturing a device for conveying the tubular medical device within the body, the device comprises a tube having an inner diameter D _4, The point is that the following steps (1) to (4) are performed in this order. However, in the following (1) to (4), D ₁ > D ₄ > D ₃ > D ₂ .
(1) A step of bringing a tubular medical device having an outer diameter of D ₁ to a temperature T ₁ in which an austenite phase is present.
(2) A preloading morphogenesis step in which an external force is applied to the tubular medical device to reduce the outer diameter from D ₁ to D ₂ at a temperature T ₂ lower than the temperature T ₁ .
(3) Loading that weakens the external force and increases the outer diameter of the tubular medical device from D ₂ to D ₃ at a temperature T ₃ lower than the austenite phase transformation end temperature A _{f of} the alloy constituting the tubular medical device. Morphogenesis step.
(4) at the temperature T _4, the tubular medical device having an outer diameter D _3, the step of loading into the lumen of the inner diameter D _4.

The difference between the outer diameter D ₂ and the outer diameter D ₃ of the tubular medical device is preferably 0.01 to 0.50 mm.

The temperatures T _1, it is preferable that the a temperature higher than the austenitic phase transformation start temperature A _s of the alloy. The temperature T ₂ is preferably a temperature lower than the R phase transformation start temperature R _{s of the} alloy. The temperature T ₂ is preferably a temperature lower than the R phase transformation end temperature R _{f of the} alloy. Wherein the temperature T ₃ is preferably a temperature lower than the R 'phase transformation finish temperature R _f' of the alloy. It is preferable that the temperatures T ₁ , T ₂ , T ₃ , and T ₄ satisfy the relationship of T ₁ > T ₂ > T ₃ = T ₄ .

Radial force F _{3 at} the outer diameter D ₃ in the middle of reducing the outer diameter of the tubular medical device from D ₁ to D ₂ in the step (2), and the tubular by increasing the diameter in the step (3). the outer diameter of the medical device 'is, F _3> F _3' radial force F ₃ at the time of the D ₃ preferably satisfy the relationship.

The tubular medical device is preferably a self-expanding stent.

According to the present invention, prior to loading a tubular medical device into the cavity of a transport device, the outer diameter of the tubular medical device is once reduced to a smaller diameter than the outer diameter at the time of loading after appropriately controlling the temperature. However, since this is expanded to the outer diameter at the time of loading, the loading load when loading the tubular medical device into the lumen of the transport device can be reduced. As a result, the tubular medical device can be easily loaded into the lumen of the transport device, so that the tubular medical device and the transport device can be prevented from being damaged.

FIG. 1 is a schematic view for explaining a method of manufacturing a tubular medical device transfer device according to the present invention. FIG. 2 is a graph showing the change in calorific value when the nickel-titanium alloy exhibits a one-step phase transformation. FIG. 3 is a graph showing the change in calorific value when the nickel-titanium alloy exhibits a two-step phase transformation. FIG. 4 is a perspective view showing an example of a self-expandable stent. FIG. 5 is a schematic diagram for explaining an OTW type self-expanding stent-carrying catheter structure. FIG. 6 is a schematic view for explaining a method of manufacturing a conventional tubular medical device transfer device. FIG. 7 is a development view showing an expanded state of the self-expandable stent used in the examples and comparative examples of the present invention. FIG. 8 is a partially enlarged view showing a substantially corrugated component that can be extended in the circumferential direction of FIG. 7. FIG. 9 is a partially enlarged view showing a connecting portion connecting adjacent substantially waveform components of FIG. 7.

The present inventors have made extensive studies to provide a method for manufacturing a tubular medical device transport device by easily loading the tubular medical device into the lumen of the transport device. As a result, prior to loading the tubular medical device into the cavity of the transport device, the temperature is appropriately controlled, and then the outer diameter of the tubular medical device is once reduced to a smaller diameter than the outer diameter at the time of loading. The present invention has been completed by finding that the loading load when loading a tubular medical device into the cavity of a transport device can be reduced by expanding the diameter to the outer diameter at the time of loading.

The method for manufacturing a tubular medical device transporting device according to the present invention is a method for manufacturing a device for transporting a tubular medical device into the body, and the device includes a tube having an inner diameter D _{4 and} is described in (1) to (4) below. ) Is characterized in that the process is performed in this order. However, in the following (1) to (4), D ₁ > D ₄ > D ₃ > D ₂ .
(1) A step of bringing a tubular medical device having an outer diameter of D ₁ to a temperature T ₁ in which an austenite phase is present.
(2) A preloading morphogenesis step in which an external force is applied to the tubular medical device to reduce the outer diameter from D ₁ to D ₂ at a temperature T ₂ lower than the temperature T ₁ .
(3) Loading that weakens the external force and increases the outer diameter of the tubular medical device from D ₂ to D ₃ at a temperature T ₃ lower than the austenite phase transformation end temperature A _{f of} the alloy constituting the tubular medical device. Morphogenesis step.
(4) at the temperature T _4, the tubular medical device having an outer diameter D _3, the step of loading into the lumen of the inner diameter D _4.

Hereinafter, a method for manufacturing a tubular medical device transporting device according to the present invention will be described with reference to FIG. 1, but the present invention is not limited thereto.

FIG. 1 is a schematic view showing a manufacturing procedure of a tubular medical device conveying device according to the present invention. In FIG. 1, 11 is a tubular medical device, 22 is an outer shaft, 31 is a tubular medical device contracting device, and 32 is an extrusion device. Shows the rod.

[Step (1)]
In the step (1), the tubular medical device having an outer diameter D ₁ is brought to a temperature T ₁ where the austenite phase is present. The step (1) will be described with reference to FIG. 1 (i).

In the step (1), as shown in FIG. 1 (i), the tubular medical device 11 is arranged in the tubular medical device contracting device 31. At this time, no external force is applied from the tubular medical device contracting device 31 to the tubular medical device 11.

The tubular medical device contraction device 31 is a device capable of uniformly applying an external force from the radial direction of the tubular medical device 11 to reduce the diameter of the tubular medical device 11, and a known device can be used. By using the tubular medical device contraction device 31, the radial force of the tubular medical device (expansion force in the radial direction of the tubular medical device) at a certain contraction diameter can be measured.

The tubular medical device contracting device 31 is described, for example, by Machine Solutions Inc. It is commercially available from (hereinafter, sometimes referred to as MSI), Blockwise Engineering LLC (hereinafter, sometimes referred to as Blockwise) and the like.

The outer diameter of the tubular medical device 11 is D _1, which is the reference outer diameter of the tubular medical device 11.

In the step (1), it is important that the tubular medical device 11 having an outer diameter D ₁ has a temperature T ₁ (° C.) in which the austenite phase is present. By setting the temperature T _{1 at} which the austenite phase exists, the effect of phase transformation due to the temperature change in the subsequent step can be surely obtained.

The tubular medical device 11 may be made of, for example, a material containing an alloy, and as the alloy, for example, a shape memory alloy can be used.

The shape memory alloy has shape memory properties and superelastic properties, and is also excellent in workability, and a nickel-titanium alloy can be used more preferably.

Among the nickel-titanium alloys, a nickel-titanium alloy containing about 50 to 60% by mass of nickel and the balance being titanium can be further preferably used.

Shape recovery by nickel-titanium alloys is generally thought to be based on a phase transformation between the martensite phase and the austenite phase (matrix phase), which is caused by temperature or stress dependence. The martensite phase refers to a crystal phase that appears at a low temperature, whereas the crystal phase that appears at a high temperature is called an austenite phase.

Here, the phase transformation of the nickel-titanium alloy will be described in detail with reference to FIGS. 2 and 3.

Nickel-titanium alloys may exhibit a one-step phase transformation or a two-step phase transformation. The one-step phase transformation is a reversible phase transformation between the austenite phase (B2 structure) and the martensite phase (B19'structure). The two-step phase transformation is a phase transformation that produces an R phase between the austenite phase (B2 structure) and the martensite phase (B19'structure). That is, the nickel-titanium alloy depends on the temperature and stress between the austenite phase (B2 structure) that is stable on the high temperature side and the martensite phase (B19'structure) that is stable on the low temperature side and has a monoclinic structure. It exhibits normal or reverse phase transformation, but a crystalline phase of rhombic structure may appear prior to the transformation of the austenite phase to the martensite phase or the transformation of the martensite phase to the austenite phase. , This crystalline phase is called the R phase (Rhombohedral phase).

The austenite phase is a phase with relatively high strength. The martensite phase is a phase that can be easily deformed by an external force and returns to its original state even if it is deformed by being strained up to about 8%.

The strain introduced into the martensite phase in the alloy to cause the shape change is restored when the reverse phase transformation to the austenite phase is complete, which restores the material to its original shape. Forward-phase transformation and reverse-phase transformation are caused by stress loading or unloading (superelastic effect), temperature change (shape memory effect), or a combination thereof. In addition, in this specification, the term "shape memory alloy" can be used interchangeably with the term "superelastic alloy" in order to indicate a material suitable for the method of the present invention.

Next, the change in calorific value when the nickel-titanium alloy undergoes a one-step phase transformation will be described with reference to FIG.

FIG. 2 shows the results of differential scanning calorimetry (DSC) of a nickel-titanium alloy showing a one-step phase transformation. The horizontal axis shows the temperature and the vertical axis shows the calorific value. The upper graph of FIG. 2 shows the result of the calorific value change when cooling, and the lower graph shows the result of the calorific value change when heating. ..

First, when the nickel-titanium alloy exhibiting a one-step phase transformation is cooled from the temperature at which it is in the austenite phase to the temperature at which it is in the martensite phase, the change in calorific value is shown in the upper graph of FIG.

In FIG. 2, M _s is the martensitic phase transformation initiation temperature, upon cooling, it shows the phase transformation starts temperature from austenite phase to the martensite phase. M _f is the temperature at which the martensitic phase transformation ends, and indicates the temperature at which the phase transformation from the austenite phase to the martensitic phase ends during cooling. M _p is the martensitic phase transformation peak temperature, and indicates the exothermic peak temperature of the martensitic phase transformation curve.

In the temperature range above the martensitic phase transformation initiation temperature ( _Ms ), only the austenite phase is present. In the temperature range below the martensitic phase transformation start temperature (M _s ) and above the martensitic phase transformation end temperature (M _f ), the austenite phase and the martensitic phase are mixed. In the temperature range below the martensitic phase transformation end temperature (M _f ), only the martensitic phase is present.

Next, when the nickel-titanium alloy exhibiting a one-step phase transformation is heated from the temperature at which it is in the martensite phase to the temperature at which it is in the austenite phase, the change in calorific value is shown in the lower graph of FIG.

In FIG. 2, A _s is an austenite phase transformation start temperature, upon heating, indicating the temperature at which the phase transformation to austenite phase (reverse transformation) begins martensite phase. A _f is the temperature at which the austenite phase transformation ends, and indicates the temperature at which the phase transformation (reverse transformation) from the martensite phase to the austenite phase ends during heating. _Ap is the austenite phase transformation peak temperature, and indicates the endothermic peak temperature of the austenite phase transformation curve.

The lower temperature range than the austenitic phase transformation starting temperature (A _s), only the martensitic phase is present. Austenite phase transformation start temperature (A _s) above, in the austenite phase transformation finish temperature (A _f) temperature range below, a state of austenite phase and martensite phase are mixed. In the temperature range above the austenite phase transformation end temperature (A _f ), only the austenite phase is present.

Each transformation temperature is usually _{M s> M p> M f} , is a _{A f> A p> A s} , not necessarily limited thereto.

Next, the change in calorific value when the nickel-titanium alloy undergoes a two-step phase transformation will be described with reference to FIG.

FIG. 3 shows the results of differential scanning calorimetry (DSC) of a nickel-titanium alloy showing a two-step phase transformation. The horizontal axis shows the temperature and the vertical axis shows the amount of heat. The upper graph of FIG. 3 shows the result of the change in the amount of heat when cooled, and the lower graph shows the result of the change in the amount of heat when heated. ..

First, when the nickel-titanium alloy exhibiting a two-step phase transformation is cooled from the temperature at which it is in the austenite phase to the temperature at which it is in the martensite phase, the change in calorific value is shown in the upper graph of FIG.

In FIG. 3, R _s is the R phase transformation start temperature, and indicates the temperature at which the phase transformation from the austenite phase to the R phase begins during cooling. R _f is the R phase transformation end temperature, and indicates the temperature at which the phase transformation from the austenite phase to the R phase ends at the time of cooling. R _p is the R-phase transformation peak temperature, and indicates the exothermic peak temperature of the R-phase transformation curve. M _s is the martensitic phase transformation start temperature, and indicates the temperature at which the phase transformation from the R phase to the martensitic phase begins during cooling. M _f is the martensitic phase transformation end temperature, and indicates the temperature at which the phase transformation from the R phase to the martensitic phase ends at the time of cooling. M _p is the martensitic phase transformation peak temperature, and indicates the exothermic peak temperature of the martensitic phase transformation curve.

In the temperature range above the R-phase transformation initiation temperature (R _s ), only the austenite phase is present. In the temperature range below the R-phase transformation start temperature (R _s ) and above the R-phase transformation end temperature (R _f ), the austenite phase and the R phase are mixed. In the temperature range lower than the R phase transformation end temperature (R _f ) and higher than the martensitic phase transformation start temperature (M _s ), only the R phase is present. In the temperature range below the martensitic phase transformation start temperature (M _s ) and above the martensitic phase transformation end temperature (M _f ), the R phase and the martensitic phase are mixed. In the temperature range below the martensitic phase transformation end temperature (M _f ), only the martensitic phase is present.

Next, when the nickel-titanium alloy exhibiting a two-step phase transformation is heated from the temperature at which it is in the martensite phase to the temperature at which it is in the austenite phase, the change in calorific value is shown in the lower graph of FIG. In FIG. 3, R _{s'is the} R'phase transformation start temperature, and indicates the temperature at which the phase transformation from the martensite phase to the R phase begins during heating. R _{f'is the} R'phase transformation end temperature, and indicates the temperature at which the phase transformation from the martensite phase to the R phase ends at the time of heating. R _{p'is the} R'phase transformation peak temperature, and indicates the endothermic peak temperature of the R phase transformation curve. A _s is an austenite phase transformation start temperature, when heated, shows a phase transformation starts temperature to the austenite phase from R-phase. A _f is the temperature at which the austenite phase transformation ends, and indicates the temperature at which the phase transformation from the R phase to the austenite phase ends during heating. _Ap is the austenite phase transformation peak temperature, and indicates the endothermic peak temperature of the austenite phase transformation curve.

In the temperature range below the R'phase transformation initiation temperature (R _s '), only the martensite phase is present. In the temperature range above the R'phase transformation start temperature (R _s ') and below the R'phase transformation end temperature (R _f '), the martensite phase and the R phase are mixed. Higher than R 'phase transformation finish temperature (R _f'), at the lower temperature range than the austenitic phase transformation starting temperature (A _s), only the R phase is present. Austenite phase transformation start temperature (A _s) above, in the austenite phase transformation finish temperature (A _f) temperature range below, a state of R-phase and austenite phase are mixed. In the temperature range above the austenite phase transformation end temperature (A _f ), only the austenite phase is present.

Each transformation temperature is usually R _s > R _p > R _f > M _s > M _p > M _f , and A _f > _Ap > A _s > R _f '> R _p '> R _s '. , Not necessarily limited to these.

The differential scanning calorimetry may be performed in accordance with, for example, ASTM F2004-05 "Standard Test Method for Transmission Temperature of Nickel-Titanium Alloys by Thermal Analysis".

The phase transformation occurs when there is no stress load on the alloy, and the phase formed by the change in temperature is called a "heat-induced phase".

On the other hand, even when the temperature does not change, phase transformation occurs when stress is applied to the alloy. The phase formed by stress loading is called the "stress-induced phase".

As the nickel-titanium alloy, a commercially available product or a product manufactured by a known method can be used.

In particular, a shape memory alloy exhibiting a two-step phase transformation is preferably produced under the following conditions, for example.

In order for the nickel-titanium alloy to exhibit an R-phase transformation, it is effective to have a nickel-rich composition. For example, it is produced by melting about 51 atomic% nickel and about 49 atomic% titanium. Is preferable.

Further, for example, the alloy may contain an alloying element in addition to nickel and titanium, and examples of the alloying element include iron element. As the iron element, for example, a trivalent or tetravalent element can be used.

It is also preferable that the alloy is cold-worked and then heat-treated at a temperature of about 400 ° C. to about 550 ° C.

By appropriately combining these, the martensitic phase transformation can be suppressed with respect to the R phase transformation, and the nickel-titanium alloy exhibits the R phase transformation.

The crystal phase of the nickel-titanium alloy can be identified by using an analysis method such as an X-ray diffraction method (XRD).

The phase transformation of the nickel-titanium alloy has been described above.

The temperatures T ₁ is preferably be a temperature higher than the austenitic phase transformation start temperature A _s of the alloy, and more preferably to a temperature higher than the austenitic phase transformation finish temperature A _f of the alloy. In particular, the austenite phase can be made into a single phase by setting the temperature higher than A _f .

[Step (2), Step (3)]
In the step (2), at a temperature T ₂ lower than the temperature T ₁ , an external force is applied to the tubular medical device to reduce the outer diameter from D ₁ to D ₂ to obtain the pre-loading form. The step (2) will be described with reference to (ii) of FIG.

In the step (2), as shown in FIG. 1 (ii), the tubular medical device contraction device 31 is used, and an external force is applied to the tubular medical device 11 from the radial direction to reduce the outer diameter from D ₁ to D ₂ . Let me.

The outer diameter D _2, in the step of which will be described later (4), a tubular medical device 11, it is important to reduce than the outer diameter D ₃ for loading the tube with an inner diameter D ₄ provided in the conveying device .. That is, in the step (2), the outer diameter of the tubular medical device 11 needs to be reduced so as to satisfy D ₁ > D ₂ and D ₁ > D ₃ > D ₂ .

As described above, the size of the outer diameter D ₂ when the tubular medical device 11 is reduced in diameter in the step (2) is not particularly limited as long as it is smaller than the outer diameter D _3, but the diameter is too reduced to be tubular. It is necessary that the medical device 11 does not buckle and is damaged.

In the step of (2), the outer diameter of the tubular medical device 11, the difference (D ₃ -D ₂₎ between the outer diameter D ₃ obtained while expanded in the step (3), for example, 0.01 It is preferable to reduce the diameter so as to be ~ 0.50 mm. The above difference is more preferably 0.03 mm or more, further preferably 0.05 mm or more, and particularly preferably 0.10 mm or more. Further, the above difference is more preferably 0.40 mm or less, further preferably 0.30 mm or less, and particularly preferably 0.20 mm or less.

The step (2) needs to be performed at a temperature T ₂ lower than the temperature T ₁ in the step (1). When the temperature T ₂ is the same as the temperature T ₁ or higher than the temperature T ₁ , the loading load when the tubular medical device 11 is loaded into the lumen provided in the tubular medical device transport device cannot be reduced. .. Further, by lowering the temperature T ₂ , the crystal phase of the alloy contains a large amount of the heat-induced R phase and / or martensite phase, and the tubular medical device 11 is reduced in diameter by applying an external force. Since more stress-induced martensite phases are expressed in the process, the reduced diameter load can be made smaller than in other temperature ranges. The temperature T ₂ is preferably lower than the R-phase transformation start temperature R _{s of the} alloy, and more preferably lower than the R-phase transformation end temperature R _{f of the} alloy.

In the step (3), the external force is weakened at a temperature T ₃ lower than the austenite phase transformation end temperature A _{f of} the alloy constituting the tubular medical device, and the outer diameter of the tubular medical device is changed from D ₂ to D ₃ . Increase the diameter. The step (3) will be described with reference to (iii) of FIG.

In the step (3), as shown in (iii) of FIG. 1, the external force applied from the tubular medical device contracting device 31 to the tubular medical device 11 is reduced, and the outer diameter of the tubular medical device 11 is changed from D ₂ to D ₃ Increase the diameter to.

After the outer diameter of the tubular medical device 11 reduced in diameter from D ₁ to D _2, by expanding the diameter in D ₂ greater than D _3, the radial force F ₃ of the tubular medical device 11 ', of the tubular medical device 11 The outer diameter can be made smaller than the radial force F ₃ of the tubular medical device 11 when the outer diameter is directly reduced from D ₁ to D ₃ . That is, when the outer diameter of the tubular medical device 11 is reduced from D ₁ to D ₃ , the radial force of the tubular medical device 11 is F ₃ , and the outer diameter of the tubular medical device 11 is smaller than D ₁ to D ₃ D ₂ Assuming that the radial force of the tubular medical device 11 is F ₂ , the radial force of the tubular medical device 11 becomes larger as the outer diameter becomes smaller, so that F ₂ > F ₃ . However, the radial force of the tubular medical device 11 when the outer diameter of the tubular medical device 11 whose outer diameter is reduced to D ₂ is expanded to D ₃ does not become the above F ₃ and is smaller than the above F _3. It becomes F ₃ '. Therefore, even if the outer diameter of the tubular medical device 11 is the same D ₃ , the outer diameter D ₃ is reduced to the outer diameter D ₃ and the outer diameter D ₂ is smaller than the outer diameter D _3. When the diameter is increased to, there is a difference of F ₃ − F ₃ ′ in the radial force. If the radial force can be reduced, the loading load when the tubular medical device 11 is loaded onto the outer shaft 22 in the step (4) described later can be reduced, so that the tubular medical device 11 can be easily manufactured.

The diameter expansion needs to be performed at a temperature T ₃ lower than the austenite phase transformation end temperature (A _f ) of the alloy constituting the tubular medical device 11. When the alloy constituting the tubular medical device 11 exhibits a two-step phase transformation, it is preferably performed at a temperature lower than the R'phase transformation end temperature (R _f '). By performing the above diameter expansion at a temperature lower than the austenite phase transformation end temperature (A _f ) or the R'phase transformation end temperature (R _f '), the crystal phase of the alloy can be obtained even when the outer diameter is relaxed to the maximum. Since it contains a large amount of R phase or martensite phase, the austenite phase is less likely to appear in the process of unloading and expanding the diameter of the tubular medical device 11, and the diameter reduction load can be made smaller.

Further, the temperature T ₃ is preferably lower than the temperature T ₂ . That is, the temperature is preferably T ₁ > T ₂ > T ₃ . By lowering the temperature for each step in this way, the consumption of energy required for cooling can be suppressed, and the diameter of the tubular medical device 11 can be efficiently reduced while the diameter reduction load is reduced as compared with the conventional case.

(3) In the process of the outer diameter of the tubular medical device 11, the difference between the outer diameter D ₂ in the steps described above (2) (D ₃ -D _2), for example, 0.01～0.50Mm It is preferable to increase the diameter so as to be.

[Step (4)]
(4) In the step, at a temperature T _4, the tubular medical device having an outer diameter D _3, loaded into the lumen of the inner diameter D _4. The step (4) will be described with reference to (iv) of FIG.

In the step of (4), as shown in (iv) in FIG. 1, the outer shaft 22 having a lumen with an inner diameter D ₄ on one side of the tubular medical device contraction device 31 disposed, tubular medical devices shrink device 31 Nomou From one side, the tubular medical device 11 is loaded into the cavity of the outer shaft 22 by the extrusion rod 32.

Radial force of the tubular medical device 11, as described in the process (3), since it is the F ₃ is less than F ₃ ', the frictional resistance load when extruding a tubular medical device 11 by the push rod 32 decreases However, it can be easily loaded into the cavity of the outer shaft 22.

The temperature T _{4 at} which the step (4) is performed is not particularly limited, but may be the same as the temperature T _{3 at} the step (3) above. By setting the temperature T ₄ and the temperature T ₃ to the same temperature, the reduced diameter load reduction effect obtained in the step (3) can be used as the friction load reduction effect when the tubular medical device 11 is loaded on the outer shaft 22. It can be applied and smooth work is possible.

The outer diameter of the extrusion rod 32 is preferably the same as the inner diameter of the tubular medical device 11 expanded to the outer diameter D ₃ , or larger than the inner diameter in the range of 0.20 mm or less, and the tubular medical device 11 It is preferable that the outer diameter is the same as D ₃ or smaller than the outer diameter in the range of 0.20 mm or less. By adjusting the outer diameter of the extrusion rod 32 to this range, the entire peripheral edge of the end portion of the tubular medical device 11 can be pushed, so that the tubular medical device 11 can be efficiently moved into the outer shaft 22 from the tubular medical device contracting device 31. Can be extruded.

The material of the extrusion rod 32 is not particularly limited as long as it is a rigid material capable of pushing the tubular medical device 11 into the outer shaft 22. As such a material, for example, a metal material such as stainless steel, aluminum, iron, nickel, titanium, or an alloy thereof, or a resin material such as polyetheretherketone, polyamide, polyimide, or polycarbonate can be used.

The method of loading the tubular medical device 11 into the outer shaft 22 may be a known method other than pushing with the extrusion rod 32.

The inner diameter D ₄ of the outer shaft 22 is preferably larger than the outer diameter D ₃ of the tubular medical device 11 in order to facilitate loading of the tubular medical device 11. That is, it is preferable that D ₁ > D ₄ > D ₃ > D ₂ . When the outer diameter D ₃ of the tubular medical device 11 is equal to or larger than the inner diameter D ₄ of the outer shaft 22, it becomes difficult to load the outer shaft 22 into the lumen. However, if the diameter of the tubular medical device 11 is reduced too much, the tubular medical device 11 may buckle and be damaged. To load the tubular medical device 11 into the lumen of the outer shaft 22 without buckling, set the outer diameter D ₃ of the tubular medical device 11 to 0.30 mm or less from the inner diameter D ₄ of the outer shaft 22. It is preferably small, more preferably small in the range of 0.01 to 0.20 mm, and even more preferably small in the range of 0.02 to 0.15 mm.

From the viewpoint of feeding the tubular medical device 11 without buckling into the lumen of the outer shaft 22, the difference between the radial force F ₃ and the radial force F ₃ 'is preferably at least 10 N, more preferably more than 20 N, 30N or more is more preferable.

In the steps (1) to (4), it is preferable that the temperatures T ₁ , T ₂ , T ₃ , and T ₄ satisfy the relationship of T ₁ > T ₂ > T ₃ = T ₄ .

In the past, in order to reduce the loading load at the time of loading, the temperature was cooled to at least the temperature at which the martensite phase develops, that is, the temperature lower than the martensite phase transformation start temperature ( _Ms ), but this temperature is controlled. Requires, for example, a refrigerant such as liquid nitrogen, and the operation is complicated. However, according to the manufacturing method of the present invention, the radial force of the tubular medical device is reduced even in an easy-to-operate temperature range by combining the reduction and expansion of the diameter of the tubular medical device by using the tubular medical device contraction device. Therefore, the loading load at the time of loading can be reduced.

The tubular medical device 11 preferably exhibits superelastic properties at the body temperature of the human body, and specifically, the alloy composition may be designed so that the austenite phase transformation end temperature (A _f ) is 37 ° C. or lower. preferable. The austenite phase transformation end temperature (A _f ) is more preferably 35 ° C. or lower, and even more preferably 33 ° C. or lower. The lower limit of the austenite phase transformation end temperature (A _f ) is not particularly limited, but is preferably 15 ° C. or higher, more preferably 20 ° C. or higher, and even more preferably 22 ° C. or higher.

The tubular medical device 11 preferably has an alloy component designed so that the R'phase transformation end temperature (R _f ') is lower than the austenite phase transformation end temperature (A _f ). The R'phase transformation end temperature (R _f ') is preferably 10 to 37 ° C, more preferably 12 to 35 ° C, and even more preferably 15 to 30 ° C.

Next, (v) of FIG. 1 will be described. FIG. 1 (v) shows a state in which the tubular medical device 11 is loaded into the lumen of the outer shaft 22. When the tubular medical device 11 is loaded into the lumen of the outer shaft 22, the tubular medical device 11 expands in diameter and becomes equal to the inner diameter D ₄ of the outer shaft 22.

As described above, according to the manufacturing method of the present invention, it is possible to reduce the loading load when the tubular medical device is loaded into the lumen of the transport device. As a result, the tubular medical device can be easily loaded into the lumen of the transport device, so that the tubular medical device and the transport device can be prevented from being damaged.

Next, the tubular medical device used in the present invention will be described.

The medical device is a tubular body, and the shape thereof is preferably a cylindrical shape.

The type of the tubular medical device is not particularly limited, and examples thereof include a stent, a stent graft, an occlusion device, an injection catheter, and a prosthesis valve. In particular, a stent can be preferably used.

The form of the stent is not particularly limited, and for example, (a) a coiled stent made of a single linear metal or polymer material, (b) a stent obtained by cutting out a metal tube with a laser, and (c). Examples thereof include a stent assembled by welding a linear member with a laser, and (d) a stent made by weaving a plurality of linear metals. The design of the stent is proposed, for example, in WO 2010/029928.

The stent preferably has a mesh structural portion.

Based on the expansion mechanism, the above-mentioned stent generally includes (a) a balloon-expandable stent in which a stent is attached to the outer surface of a balloon and transported to a lesion, and the stent is expanded by a balloon at the lesion, and (b) a stent. It can be classified as a self-expanding stent that expands by itself by loading it into a catheter having a sheath member that suppresses the expansion of the stent, transporting the stent to the lesion, and removing the sheath member at the lesion.

In the present invention, a self-expandable stent can be preferably used as the tubular medical device.

Hereinafter, the self-expandable stent will be described in more detail, but the medical device used in the present invention is not limited thereto.

FIG. 4 is a perspective view of the self-expandable stent 11a, and the self-expandable stent 11a shown in FIG. 4 is a tubular body having a mesh structure portion and shows an expanded state.

The reference outer diameter of the self-expandable stent 11a is D ₁ and the reference length is L ₁ . The reference outer diameter D ₁ is the outer diameter before being contracted and loaded into the lumen of the transport device.

The reference outer diameter D ₁ and the reference length L ₁ of the self-expandable stent 11a are appropriately selected according to the inner diameter and length of the lumen in the lesion, and differ depending on the lumen to be treated. For example, as the self-expandable stent 11a for the treatment of lower limb arteries, a stent having a reference outer diameter D ₁ of, for example, 4.0 to 12.0 mm and a reference length L ₁ of, for example, 20 to 200 mm is preferably used. Be done.

A shape memory alloy can be preferably used as the material of the self-expandable stent 11a.

The self-expandable stent 11a has a Radial Resistive Force (hereinafter sometimes referred to as RRF) that exhibits an expanding force that the stent tries to resist compression in order to effectively expand and maintain the body lumen in which it is placed. It is preferably 0.200 to 1.500 N / mm. If the RRF is too small, it will be difficult to dilate and maintain the body lumen to be placed due to insufficient dilation force, and there is a risk of falling out of the placement position. The RRF is more preferably 0.400 N / mm or more. The RRF is preferably as large as possible from the viewpoint of dilation and maintenance of the body lumen. However, if the RRF is too large, the force to dilate the body lumen becomes too large, and the body lumen may be rather damaged. The RRF is more preferably 1.200 N / mm or less. The radius of the RRF is 0.1 mm / sec and a radius of the self-expandable stent 11a to the diameter obtained by subtracting 2.5 mm from the reference outer diameter D _{1 in} an atmosphere of 37.0 ± 1.0 ° C. the contraction径荷heavy when the reduced diameter by compressing from the direction, the value obtained by dividing the reference length L ₁ of the self-expanding stent 11a.

The RRF can be measured, for example, by using a device commercially available from MSI, Blockwise and the like.

The self-expandable stent 11a is preferably highly flexible so that it can follow the dynamics of the body lumen in which it is placed. Specifically, using a tensile compression tester (for example, "EZ-Test" manufactured by Shimadzu Corporation), in an atmosphere of 37.0 ± 1.0 ° C., with a flat indenter, at a speed of 0.1 mm / sec. The major axis compression strength when compressed from the major axis direction of the self-expandable stent 11a to a displacement amount of 7% of the reference length L ₁ is preferably 0.030 to 1.500 N. The major axis compression strength is more preferably 0.050 N or more, and more preferably 1.300 N or less.

The self-expandable stent 11a can be produced, for example, by laser-cutting a pipe made of a nickel-titanium alloy, expanding the diameter, heat-treating it to form a desired shape, and finally electropolishing it. ..

Next, the tubular medical device transfer device manufactured by the present invention will be described. Hereinafter, the case where a self-expandable stent is used as a tubular medical device will be described in detail, but the present invention is not limited thereto.

As a self-expanding stent transfer device, an OTW (Over The Ware) type catheter structure and an RX (Rapid Exchange) type catheter structure are generally known. Of these, the OTW-type self-expanding stent-carrying catheter structure will be described with reference to FIG. 5, but the tubular medical device-carrying device according to the present invention is not limited thereto.

FIG. 5 is a schematic view showing the entire self-expandable stent transfer device.

The self-expandable stent transfer device 21 transports the self-expandable stent 11a to a lesion portion (for example, a stenosis portion) of the body lumen, is insertable into the body lumen, is elongated, and has flexibility. are doing. Note that FIG. 5 shows a state in which the self-expandable stent 11a is loaded inside the outer shaft 22, and the self-expandable stent 11a is in a contracted state.

The self-expandable stent transfer device 21 shown in FIG. 5 has a self-expandable stent 11a, an outer shaft 22, and an inner shaft 23.

(Outer shaft)
The self-expandable stent 11a is housed in the outer shaft 22 in a contracted state, and the outer diameter of the self-expandable stent 11a is D ₄ , which is equal to the inner diameter of the outer shaft 22.

When the self-expandable stent 11a housed in the outer shaft 22 is transported to the lesion and released from the outer shaft 22, the outer diameter D ₄ of the contracted self-expandable stent 11a exceeds the inner diameter of the outer shaft 22. To expand the diameter.

The outer shaft 22 has flexibility to follow the lumen to be inserted (for example, a blood vessel, etc.), kink resistance, and tensile strength to the extent that the self-expandable stent transfer device 21 does not stretch when pulled during the procedure. It is preferably formed of a material.

The inside of the outer shaft 22 is preferably made of a material that reduces movement resistance (sliding resistance) in order to easily release the self-expandable stent 11a in contact with the inner peripheral surface to the lesion portion. The outside of the is preferably a material having low friction so that the movement operation in the lumen can be easily performed.

From this point of view, the outer shaft 22 is preferably formed of, for example, a tube having a three-layer structure in which the outer layer and the inner layer are formed of a resin material and a reinforcing layer is embedded between the outer layer and the inner layer. As the reinforcing layer, for example, it is preferable to use a layer of a metal wire.

For the outer layer of the outer shaft 22, it is preferable to use various elastic resin materials such as polyethylene, fluororesin (for example, PTFE, PFA, etc.), polyamide, polyamide-based elastomer, polyurethane, polyester, and silicone.

For the inner layer of the outer shaft 22, for example, it is preferable to use a low friction material such as fluororesin (for example, PTFE, PFA, etc.) or polyethylene.

For the reinforcing layer (layer of the metal wire) of the outer shaft 22, it is preferable to use various metal materials such as stainless steel, nickel-titanium alloy, tungsten, gold, and platinum.

The metal wire includes at least one structure of a braided structure or a coil structure, and is preferably formed from the proximal end to the distal end of the outer shaft 22.

The proximal end (proximal end side) means the hand side of the operator, and the distal end (distal end side) means the opposite side of the proximal end, that is, the side close to the lesion or the like.

The portion of the outer shaft 22 that holds the self-expandable stent 11a may be formed of a tube having a one-layer structure or a structure in which two or more layers are laminated, and is preferably a one-layer structure or a two-layer structure. For such a layer, it is preferable to use the material used for the outer layer and / or the material used for the inner layer.

(Inner shaft)
The inner shaft 23 is a tubular member having a guide wire lumen, and a pusher member 24 is attached to the proximal side of the self-expandable stent 11a, a tip tip 25 is attached to the distal side, and an operating member 26 is attached to the proximal side. Is preferable.

At least a part of the tubular member having the guide wire cavity of the inner shaft 23 is inserted into the lumen of the outer shaft 22. Then, a guide wire is inserted into the lumen formed in the inner shaft 23, and the self-expandable stent transfer device 21 can be guided to the lesion portion.

The inner shaft 23 preferably has flexibility, kink resistance, and tensile strength such that it does not stretch when the catheter is pulled during the procedure, so that it follows the lumen of the outer shaft 22.

For the inner shaft 23, for example, it is preferable to use various elastic resin materials such as polyethylene, fluororesin (for example, PTFE, PFA, etc.), polyamide, polyamide-based elastomer, polyurethane, polyester, polyimide, and silicone.

The inner shaft 23 may be reinforced with a metal material, and examples of the metal material include stainless steel, nickel-titanium alloy, tungsten, gold, platinum, and the like.

(Pusher member)
The pusher member 24 is preferably of a size that is mounted (adhered or welded) around the inner shaft 23 and fits within the lumen of the outer shaft 22.

The self-expandable stent 11a can be released from the outer shaft 22 according to the relative movement of the outer shaft 22 and the inner shaft 23.

The shape of the pusher member 24 is, for example, a ring shape, and the outer diameter thereof is preferably equal to or greater than the inner diameter of the contracted self-expandable stent 11a and less than or equal to the inner diameter of the outer shaft 22 (that is, the outer diameter D ₄ ). By setting the outer diameter of the pusher member 24 to this range, the pusher member 24 covers the peripheral edge of the self-expandable stent 11a, so that the self-expandable stent 11a can be efficiently released.

For the pusher member 24, various elastic resin materials such as polyethylene, fluororesin (for example, PTFE, PFA, etc.), polyamide, polyamide-based elastomer, polyurethane, polyester, and silicone can be used.

The pusher member 24 preferably contains a material that is at least partially opaque to X-rays. By using at least a part of an X-ray opaque material, the self-expandable stent carrier 21 can be operated under fluoroscopy, so that the self-expandable stent 11a can be safely and efficiently placed in the lumen of the body. Can be transported to the lesion site and released.

Examples of the X-ray opaque material include contrast-enhancing materials such as X-ray contrast-enhancing substances and contrast-enhancing substances such as ultrasonic contrast-enhancing substances. Specifically, for example, gold, platinum, tungsten, tantalum, iridium, palladium or an alloy thereof, or gold-palladium alloy, platinum-iridium, NiTiPd, NiTiAu, barium sulfate, bismuth compound, tungsten compound, etc. Can be mentioned.

(Tip tip)
The tip tip 25 is attached (for example, bonded or welded) to the tip of the inner shaft 23. The tip 25 facilitates the passage of the self-expanding stent carrier 21 through a lesion (eg, a stenosis).

It is preferable that the tip tip 25 has flexibility enough to follow the lumen to be inserted (for example, a blood vessel or the like) and rigidity in the major axis direction enough to pass through the narrowed portion.

For the tip 25, various elastic resin materials such as polyethylene, fluororesin (for example, PTFE, PFA, etc.), polyamide, polyamide-based elastomer, polyurethane, polyester, and silicone can be used.

The tip 25 preferably contains a material that is at least partially opaque to X-rays. By using at least a part of an X-ray opaque material, the self-expandable stent carrier 21 can be operated under fluoroscopy, so that the self-expandable stent 11a can be safely and efficiently placed in the lumen of the body. Can be transported to the lesion site and released.

Examples of the X-ray opaque material include contrast-enhancing materials such as X-ray contrast-enhancing substances and contrast-enhancing substances such as ultrasonic contrast-enhancing substances. Specifically, for example, gold, platinum, tungsten, tantalum, iridium, palladium or an alloy thereof, or gold-palladium alloy, platinum-iridium, NiTiPd, NiTiAu, barium sulfate, bismuth compound, tungsten compound, etc. Can be mentioned.

(Operating member)
The operating member 26 is attached (eg, bonded or welded) to the proximal end of the inner shaft 23.

The shape of the operating member 26 is not particularly limited as long as it is easy to grip for the operator to operate, and for example, as shown in FIG. 5, the shape may include a grip portion that can be gripped by hand. preferable.

As the material of the operating member 26, for example, a metal material such as stainless steel, aluminum, iron, nickel, or titanium, or a resin material such as polyetheretherketone, polyamide, polyimide, or polycarbonate can be used.

Hereinafter, the present invention will be described in more detail with reference to Examples, but the present invention is not limited by the following Examples, and the present invention shall be carried out with modifications to the extent that it can be adapted to the gist of the above and the following. Of course, they are all within the technical scope of the present invention.

In manufacturing a device for transporting a tubular medical device, the loading load when the tubular medical device was loaded into the lumen provided in the transport device was measured. Specifically, a self-expandable stent was used as a tubular medical device, and the self-expandable stent was inserted into the lumen of a catheter which is a transport device to manufacture a tubular medical device transport device.

(Example 1)
As a tubular medical device, a self-expandable stent made of nickel-titanium alloy having a reference outer diameter D ₁ of 6.0 mm and a reference length L ₁ of 100 mm was prepared. The design of the prepared self-expanding stent is detailed below with reference to FIGS. 7-9. 7 is an expanded view showing the expanded state of the self-expandable stent, FIG. 8 is a partially enlarged view showing a substantially corrugated component in the expanded state of the self-expandable stent, and FIG. 9 is an expanded view of the self-expandable stent. A partially enlarged view showing the connecting portion in the state is shown respectively.

The self-expandable stent design has an annular section 71 consisting of substantially corrugated components 72 aligned with respect to the longitudinal axis and a connecting portion 73 connecting the approximately corrugated components 72 in adjacent annular sections 71. The connecting portion 73 connects the vertices 87 of the substantially corrugated components to each other. In the expanded state, a line connecting adjacent connecting portions 73 connected to each other without interposing another connecting portion 73 forms two spirals opposite to each other with respect to the longitudinal axis, and these spirals The wavelengths are different from each other. More specifically, 55 annular sections 71 are aligned with respect to the longitudinal axis, the number of substantially corrugated components 72 in the annular section 71 is 16, and the number of connecting portions 73 included in the annular section 71. Is four.

Further, the central axes of the inner radius of curvature between the struts forming the adjacent substantially corrugated component 72 are on the same straight line in the circumferential direction, and the strut inner radius of curvature center 81 and the outer radius of curvature center of the substantially corrugated component 72 are formed. The distance to 82 is 30 μm. Further, the outer radius of curvature 84 between the struts of the substantially corrugated component 72 is 110 μm, and the inner radius of curvature 83 is 30 μm. The width 86 of the apex of the substantially corrugated component 71 is 110 μm, the width 85 of the strut is 80 μm, and the wall thickness of the strut is 200 μm.

The nickel-titanium alloy complies with ASTM F2063-05.

The RRF of the self-expanding stent was calculated by the following procedure. Using an RRF measuring machine (“TTR2” manufactured by Blockwise), a self-expandable stent minus 2.5 mm from the standard outer diameter D ₁ (6.0 mm) in an atmosphere of 37.0 ± 1.0 ° C. Measure the diameter reduction load when the diameter is reduced by compressing from the radial direction at a speed of 0.1 mm / sec up to the diameter (that is, 3.5 mm), and the diameter reduction load is the reference length of the self-expandable stent. The RRF was calculated by dividing by L ₁ (100 mm). As a result of the calculation, the RRF was 1.001 N / mm.

The long-axis compressive strength of the self-expanding stent was measured by the following procedure. Using a tensile compression tester (“EZ-Test” manufactured by Shimadzu Corporation), in an atmosphere of 37.0 ± 1.0 ° C., with a flat indenter, at a speed of 0.1 mm / sec, the long axis of the self-expandable stent. The load when compressed (that is, compressed to 7 mm) from the direction to a displacement of 7% of the reference length L ₁ (100 mm) was measured as the major axis compression strength. As a result of the measurement, the long axis compression strength was 0.081N.

The nickel-titanium alloy is an alloy that exhibits a two-step phase transformation, and becomes a martensite phase, an R phase, and an austenite phase depending on temperature and / or stress.

Table 1 below shows the phase transformation temperatures when the nickel-titanium alloy is cooled or heated.

The phase transformation temperature of the self-expanding stent was measured using a differential scanning calorimeter (“DSC7020” manufactured by Hitachi High-Tech Science) according to ASTM F2004-05 and ASTM F2005-05. The sample mass of the self-expandable stent was 20 to 30 mg, the scanning temperature rate was 10 ° C./min, and the scanning temperature range was −120 ° C. to 100 ° C. The measurement results are shown in Table 1 below.

In Table 1 below, R _s is the transformation start temperature from the austenite phase to the R phase during cooling, R _p is the transformation peak temperature from the austenite phase to the R phase (exothermic peak temperature), and R _f is the austenite phase to the R phase during cooling. The transformation end temperature to, M _s is the transformation start temperature from the R phase to the martensite phase during cooling, M _p is the transformation peak temperature from the R phase to the martensite phase (exothermic peak temperature), and M _f is the R phase during cooling. from transformation finish temperature to the martensite phase, R _s 'is transformation start temperature to the R-phase from martensite phase during heating, R _p' is the transformation peak temperature to the R-phase from martensite phase (endothermic peak temperature), R _f 'is transformation finish temperature to the R-phase from martensite phase during heating, a _s is transformation start temperature to the austenite phase from R-phase during heating, a _p is the transformation peak temperature to the austenite phase from R-phase (endothermic peak temperature ) And A _f indicate the temperature at which the transformation from the R phase to the austenite phase ends during heating, respectively.

As the outer shaft 22, a polyamide-based elastomer was used for the outer layer, polytetrafluoroethylene (PTFE) was used for the inner layer, and a tube having a braided structure using a flat wire of stainless steel having a width of 100 μm and a thickness of 25 μm was used as the reinforcing layer. ..

The inner diameter D ₄ of the lumen formed on the outer shaft 22 was 1.80 mm.

In the procedure shown in FIG. 1, a self-expandable stent 11 was placed in the outer shaft 22 using a tubular medical device contraction device 31 (“RFL225” manufactured by Blockwise).

Hereinafter, the charging procedure will be described with reference to FIG.

FIG. 1 (i) shows a state in which the self-expandable stent 11 having an outer diameter D ₁ is arranged in the tubular medical device contraction device 31. In (i), the temperature T ₁ was set to 37.0 ± 1.0 ° C. The reference outer diameter D ₁ of the self-expandable stent 11 is 6.0 mm.

Next, FIG. 1 (ii) shows a state in which the diameter of the self-expandable stent 11 is reduced by applying an external force using the tubular medical device contraction device 31. In (ii), the temperature T ₂ is 25.0 ± 1.0 ° C., the speed is 0.1 mm / sec, and the diameter is reduced from the reference outer diameter D ₁ (6.0 mm) to 1.45 mm (outer diameter D ₂ ). I let you. The radial force F ₂ when the diameter was reduced to the outer diameter D ₂ (1.45 mm) was measured with a tubular medical device contracting device 31 and found to be 151.8 N.

The radial force F ₃ was 133.6 N when the outer diameter became 1.75 mm (outer diameter D ₃ ) while the diameter was reduced from the reference outer diameter D ₁ to the outer diameter D ₂ .

Next, FIG. 1 (iii) shows a state in which the diameter of the self-expandable stent 11 is expanded by weakening the external force applied from the tubular medical device contraction device 31. In (iii), the temperature T _{3 is set} to 25.0 ± 1.0 ° C., the speed is set to 0.1 mm / sec, and the outer diameter is increased from D ₂ (1.45 mm) to 1.75 mm (outer diameter D ₃ ). It was. Outer diameter D ₃ radial force F ₃ when expanded in diameter (1.75 mm) 'was 82.3N.

When the outer diameter becomes D ₃ while reducing the diameter from the standard outer diameter D ₁ to the outer diameter D ₂ , the radial force F ₃ is 133.6 N, and the diameter is expanded from the outer diameter D ₂ to the outer diameter D ₃ . 'because was 82.3N, radial force F ₃ and radial force F _3' is radial force F ₃ when the difference (F ₃ -F ₃ ') was 51.3N.

Next, in FIG. 1 (iv), an outer shaft 22 having an inner diameter D ₄ of 1.80 mm is prepared on one side of the tubular medical device shrinking device 31, and an outer diameter D is prepared from the other side of the tubular medical device shrinking device 31. _It shows a state in which the self-expandable stent 11 held at ₃ (1.75 mm) is pushed by a solid extrusion rod 32 and loaded into the lumen of the outer shaft 22. In (iv), the temperature T ₄ was set to 25.0 ± 1.0 ° C. The solid extrusion rod 32 was made of SUS304 having an outer diameter of 1.72 mm, and the loading speed was 10 mm / sec.

A force gauge is attached to the end opposite to the end where the extrusion rod 32 and the self-expandable stent 11 abut, and the maximum load (static friction load) when the self-expansion stent 11 starts to move by the extrusion rod 32 is loaded. Was measured as. As the force gauge, "FGC-10" manufactured by Nidec Symposium was used. As a result of measurement, the loading load was 11.9 N.

Next, FIG. 1 (v) shows a state in which the self-expandable stent 11 is loaded into the lumen of the outer shaft 22. The outer diameter of the self-expanding stent 11 fully loaded into the lumen of the outer shaft 22 was comparable to the inner diameter D ₄ (1.80 mm) of the outer shaft 22.

Then, as shown in FIG. 5, the inner shaft 23 to which the pusher member 24 is mounted is inserted into the lumen of the outer shaft 22 in which the self-expandable stent 11 is housed, and is inserted into the distal end of the inner shaft 23. The tip 25 was attached, the operating member 26 was attached to the proximal end of the inner shaft 23, and other members were appropriately attached to manufacture the tubular medical device transport device 21.

(Example 2)
In the first embodiment, the temperature T ₂ in FIG. 1 (ii) is 13.0 ± 1.0 ° C., the temperature T ₃ in FIG. 1 (iii) is 13.0 ± 1.0 ° C., and (1) in FIG. The self-expandable stent 11 was loaded into the lumen of the outer shaft 22 under the same conditions as in Example 1 above, except that the temperature T ₄ in iv) was changed to 13.0 ± 1.0 ° C., as shown in FIG. The tubular medical device transfer device 21 shown was manufactured.

When the self-expandable stent 11 (reference outer diameter D ₁ was 6.0 mm) was reduced to an outer diameter D ₂ (1.45 mm), the radial force F ₂ was 139.6 N.

The radial force F ₃ was 120.4 N when the outer diameter became 1.75 mm (outer diameter D ₃ ) in the middle of reducing the diameter from the reference outer diameter D ₁ to the outer diameter D ₂ .

The self-expanding stent 11 outer diameter D ₂ radial force F ₃ when the enlarged loosen the target outer diameter D ₃ from (1.45mm) (1.75mm) 'was 67.7N.

When the outer diameter becomes D ₃ while reducing the diameter from the reference outer diameter D ₁ to the outer diameter D ₂ , the radial force F ₃ is 120.4 N, and the diameter is expanded from the outer diameter D ₂ to the outer diameter D ₃ . 'because was 67.7N, radial force F ₃ and radial force F _3' is radial force F ₃ when the difference (F ₃ -F ₃ ') was 52.7N.

As a result of measuring the maximum load (static friction load) when the self-expandable stent 11 starts to move by the extrusion rod 32, the loading load was 9.4N.

(Example 3)
In Example 1, except that the temperature T ₃ in FIG. 1 (iii) is changed to 13.0 ± 1.0 ° C. and the temperature T ₄ in FIG. 1 (iv) is changed to 13.0 ± 1.0 ° C. The self-expandable stent 11 was loaded into the lumen of the outer shaft 22 under the same conditions as in Example 1 to manufacture the tubular medical device transport device 21 shown in FIG.

The radial force F ₂ when the self-expandable stent 11 (reference outer diameter D ₁ was 6.0 mm) was reduced to the outer diameter D ₂ (1.45 mm) was 152.3 N.

The radial force F ₃ was 134.2 N when the outer diameter became 1.75 mm (outer diameter D ₃ ) in the middle of reducing the diameter from the reference outer diameter D ₁ to the outer diameter D ₂ .

Radial force F ₃ when the outer diameter was enlarged from D ₂ (1.45 mm) to the target outer diameter D ₃ (1.75 mm) of self-expanding stents 11 'was 70.3N.

When the outer diameter becomes D ₃ while reducing the diameter from the standard outer diameter D ₁ to the outer diameter D ₂ , the radial force F ₃ is 134.2 N, and the diameter is expanded from the outer diameter D ₂ to the outer diameter D ₃ . 'because was 70.3N, radial force F ₃ and radial force F _3' is radial force F ₃ when the difference (F ₃ -F ₃ ') was 63.9N.

As a result of measuring the maximum load (static friction load) when the self-expandable stent 11 starts to move by the extrusion rod 32, the loading load was 9.8 N.

(Comparative Example 1)
In the first embodiment, instead of the procedure shown in FIG. 1, a tubular medical device contracting device 31 (“RFL225” manufactured by Blockwise) is used in the procedure shown in FIG. 6, and a self-expanding type is used in the outer shaft 22. Stent 11 was charged.

FIG. 6 is a schematic view for explaining a method of manufacturing a conventional tubular medical device transfer device.

Hereinafter, the charging procedure will be described with reference to FIG.

FIG. 6 (i) shows a state in which the self-expandable stent 11 having an outer diameter D ₁ is arranged in the tubular medical device contracting device 31. In (i), the temperature T ₁ was set to 37.0 ± 1.0 ° C. The reference outer diameter D ₁ of the self-expandable stent 11 is 6.0 mm.

Next, FIG. 6 (ii) shows a state in which the diameter of the self-expandable stent 11 is reduced by applying an external force using the tubular medical device contraction device 31. In (ii), the temperature T _{2 is set} to 25.0 ± 1.0 ° C., the speed is set to 0.1 mm / sec, and the reference outer diameter D ₁ (6.0 mm) is reduced to 1.75 mm (target outer diameter D ₃ ). The diameter was increased. The radial force F ₃ when the outer diameter was reduced to D ₃ (1.75 mm) was 132.8 N.

Next, in FIG. 6 (iii), an outer shaft 22 having an inner diameter D ₄ of 1.80 mm is prepared on one side of the tubular medical device shrinking device 31, and an outer diameter D ₃ is prepared from the other side of the tubular medical device shrinking device 31. It shows a state in which the self-expandable stent 11 held at (1.75 mm) is pushed by a solid extrusion rod 32 and loaded into the lumen of the outer shaft 22. In (iii), the temperature T ₄ was set to 25.0 ± 1.0 ° C. The solid extrusion rod 32 was made of SUS304, had an outer diameter of 1.72 mm, and had a loading speed of 10 mm / sec.

A force gauge is attached to the end opposite to the end where the extrusion rod 32 and the self-expandable stent 11 abut, and the maximum load (static friction load) when the self-expansion stent 11 starts to move by the extrusion rod 32 is loaded. Was measured as. As the force gauge, "FGC-10" manufactured by Nidec Symposium was used. As a result of measurement, the loading load was 23.4N.

Next, FIG. 6 (iv) shows a state in which the self-expandable stent 11 is loaded into the lumen of the outer shaft 22. The outer diameter of the self-expanding stent 11 fully loaded into the lumen of the outer shaft 22 was comparable to the inner diameter D ₄ (1.80 mm) of the outer shaft 22.

Then, as shown in FIG. 5, the inner shaft 23 to which the pusher member 24 is mounted is inserted into the lumen of the outer shaft 22 in which the self-expandable stent 11 is housed, and is inserted into the distal end of the inner shaft 23. The tip 25 was attached, the operating member 26 was attached to the proximal end of the inner shaft 23, and other members were appropriately attached to manufacture the tubular medical device transport device 21.

(Comparative Example 2)
In Comparative Example 1 above, except that the temperature T ₂ in FIG. 6 (ii) is changed to 13.0 ± 1.0 ° C. and the temperature T ₄ in FIG. 6 (iii) is changed to 13.0 ± 1.0 ° C. The self-expandable stent 11 was loaded into the lumen of the outer shaft 22 under the same conditions as in Comparative Example 1, and the tubular medical device transport device 21 shown in FIG. 5 was manufactured.

When the self-expandable stent 11 (reference outer diameter D ₁ was 6.0 mm) was reduced to the target outer diameter D ₃ (1.75 mm), the radial force F ₃ was 119.2 N.

As a result of measuring the maximum load (static friction load) when the self-expandable stent 11 starts to move by the extrusion rod 32, the loading load was 17.6 N.

(Comparative Example 3)
In the first embodiment, the temperature T ₂ in (ii) of FIG. 1 is 37.0 ± 1.0 ° C., the temperature T ₃ in (iii) of FIG. 1 is 37.0 ± 1.0 ° C., and (1) of FIG. A self-expandable stent 11 was loaded into the lumen of the outer shaft 22 under the same conditions as in Example 1 above, except that the temperature T ₄ in iv) was changed to 37.0 ± 1.0 ° C., as shown in FIG. The tubular medical device transfer device 21 shown was manufactured.

When the self-expandable stent 11 (reference outer diameter D ₁ was 6.0 mm) was reduced to an outer diameter D ₂ (1.45 mm), the radial force F ₂ was 168.7 N.

The radial force F ₃ was 149.1 N when the outer diameter became 1.75 mm (outer diameter D ₃ ) in the middle of reducing the diameter from the reference outer diameter D ₁ to the outer diameter D ₂ .

Radial force F ₃ when the outer diameter was enlarged from D ₂ (1.45 mm) to the target outer diameter D ₃ (1.75 mm) of self-expanding stents 11 'was 99.0N.

When the outer diameter becomes D ₃ while reducing the diameter from the reference outer diameter D ₁ to the outer diameter D ₂ , the radial force F ₃ is 149.1 N, and the diameter is expanded from the outer diameter D ₂ to the outer diameter D ₃ . 'because was 99.0N, radial force F ₃ and radial force F _3' is radial force F ₃ when the difference (F ₃ -F ₃ ') was 50.1N.

As a result of measuring the maximum load (static friction load) when the self-expandable stent 11 starts to move by the extrusion rod 32, the load load was 15.2 N.

The above results are summarized in Table 2 below.

It can be considered from Table 2 below as follows. Examples 1 to 3 are examples of manufacturing a tubular medical device transporting device under conditions that satisfy the requirements specified in the present invention, and Comparative Examples 1 to 3 are conditions that do not satisfy any of the requirements specified in the present invention. This is an example manufactured in.

As shown in Examples 1 to ₃ , the comparison is made by manufacturing a tubular medical device through diameter reduction and diameter expansion under the conditions of D ₁ > D ₄ > D ₃ > D ₂ after appropriately controlling the temperature. As shown in Examples 1 and 2, it can be seen that the loading load can be reduced as compared with the case where the tubular medical device is manufactured without expanding the diameter as D ₁ > D ₄ > D ₃ . That is, in Examples 1 to 3, since the can greatly reduce the radial force F ₃ 'with respect to radial force F _3, the loading load is reduced.

Further, Comparative Example 3 is an example in which the tubular medical device transport device is manufactured by the same procedure as in Examples 1 to 3, but since the temperature of each step is not appropriately controlled, it is radial with respect to the radial force F ₃ . not been able to sufficiently reduce the force F ₃ '. As a result, the loading load became larger than the loading load of Examples 1 to 3. That is, the temperature T _{2 at} which the step (2) is performed is set lower than the R-phase transformation start temperature (R _s ) as shown in Examples 1 and 3, or the R-phase transformation ends as shown in Example 2. Either make it lower than the temperature (R _f ), or make the temperature T _{3 at} which the step (3) is performed lower than the austenite phase transformation end temperature (A _f ) as shown in Example 1, or make it lower than the temperature (A _f ), or in Examples 2 and 3. As shown, it can be seen that the loading load can be reduced by lowering the temperature below the R'phase transformation end temperature (R _f '). In particular, Example 3, since the temperature is in the _{T 1> T 2> T 3} = T 4, in terms of a reduced cooling energy, sufficiently reduce the radial force F ₃ 'with respect to radial force F ₃ is made of.

As described above, according to the manufacturing method of the present invention, even a tubular medical device having a large radial force can reduce the loading load when it is loaded into the tubular medical device transport device. As a result, a medical device delivery system having excellent operability and passability in the lumen of the body can be easily manufactured.

11 Tubular medical device 11a Self-expandable stent 21 Tubular medical device transport device 22 Outer shaft 23 Inner shaft 24 Pusher member 25 Tip tip 26 Operating member 31 Tubular medical device contraction device 32 Extruded rod 71 Annular section 72 Approximately corrugated component 73 Connecting part 81 Center of inner radius of curvature 82 Center of outer radius of curvature 83 Inner radius of curvature 84 Outer radius of curvature 85 Strut width 86 Apex width 87 Approximate waveform component apex L ₁ Reference length D ₁ Reference outer diameter D ₂ Outer diameter D ₃ outer diameter D ₄ inside diameter F ₂ radial force F ₃ radial force F ₃ 'radial force R _s R-phase transformation start temperature R _p R phase transformation peak temperature R _f R-phase transformation finish temperature M _s martensite transformation start temperature M _p Martens sites phase transformation peak temperature M _f martensite transformation finish temperature R _s 'R' phase transformation start temperature R _p 'R' phase transformation peak temperature R _f 'R' phase transformation finish temperature A _s austenite transformation start temperature A _p austenite Phase transformation peak temperature A _f Austenite phase transformation end temperature T ₁ temperature T ₂ temperature T ₃ temperature T ₄ temperature

Claims (10)

A method of manufacturing a device that transports a tubular medical device into the body.
The device includes a tube having an inner diameter D _4,
A method for manufacturing a tubular medical device transfer device, which comprises performing the following steps (1) to (4) in this order.
(1) A step of bringing a tubular medical device having an outer diameter of D ₁ to a temperature T ₁ in which an austenite phase is present.
(2) A preloading morphogenesis step in which an external force is applied to the tubular medical device to reduce the outer diameter from D ₁ to D ₂ at a temperature T ₂ lower than the temperature T ₁ .
(3) Loading that weakens the external force and increases the outer diameter of the tubular medical device from D ₂ to D ₃ at a temperature T ₃ lower than the austenite phase transformation end temperature A _{f of} the alloy constituting the tubular medical device. Morphogenesis step.
(4) at the temperature T _4, the tubular medical device having an outer diameter D _3, the step of loading into the lumen of the inner diameter D _4.
However, in the above (1) to (4), D ₁ > D ₄ > D ₃ > D ₂ . The manufacturing method according to claim 1, wherein the difference between the outer diameter D ₂ and the outer diameter D ₃ of the tubular medical device is 0.01 to 0.50 mm. The temperature T ₁ is The process according to claim 1 or 2 at a temperature higher than the austenitic phase transformation start temperature A _s of the alloy. The production method according to any one of claims 1 to 3, wherein the temperature T ₂ is lower than the R phase transformation start temperature R _{s of the} alloy. The production method according to any one of claims 1 to 4, wherein the temperature T ₂ is lower than the R phase transformation end temperature R _{f of the} alloy. Wherein the temperature T ₃ The production method according to claim 1 wherein a temperature lower than the R 'phase transformation finish temperature R _f' of the alloy. The production method according to any one of claims 1 to 6, wherein the temperatures T ₁ , T ₂ , T ₃ , and T ₄ satisfy the relationship of T ₁ > T ₂ > T ₃ = T ₄ . In the step (2), the radial force F _{3 at} the outer diameter D ₃ in the middle of reducing the outer diameter of the tubular medical device from D ₁ to D ₂
Wherein the (3) in step radial force F ₃ when the outer diameter of the tubular medical device was D ₃ by diameter '
The manufacturing method according to any one of claims 1 to 7, which satisfies the relationship of F ₃ > F ₃ '. The manufacturing method according to any one of claims 1 to 8, wherein the tubular medical device is a self-expandable stent. In the step (2), the outer diameter of the tubular medical device is set to D in the step (3). ₂ From D ₃ Difference when expanding the diameter to (D) ₃ -D ₂ ) Is 0.01 to 0.50 mm, so that the outer diameter of the tubular medical device is D. ₁ From D ₂ The manufacturing method according to any one of claims 1 to 9, wherein the diameter is reduced to.

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