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Publication numberWO1997007740 A1
Publication typeApplication
Application numberPCT/US1996/013629
Publication date6 Mar 1997
Filing date22 Aug 1996
Priority date24 Aug 1995
Also published asEP0847249A1, EP0847249A4, US6799075, US20010009970
Publication numberPCT/1996/13629, PCT/US/1996/013629, PCT/US/1996/13629, PCT/US/96/013629, PCT/US/96/13629, PCT/US1996/013629, PCT/US1996/13629, PCT/US1996013629, PCT/US199613629, PCT/US96/013629, PCT/US96/13629, PCT/US96013629, PCT/US9613629, WO 1997/007740 A1, WO 1997007740 A1, WO 1997007740A1, WO 9707740 A1, WO 9707740A1, WO-A1-1997007740, WO-A1-9707740, WO1997/007740A1, WO1997007740 A1, WO1997007740A1, WO9707740 A1, WO9707740A1
InventorsVictor I. Chornenky, Michael R. Forman
ApplicantInterventional Innovations Corporation
Export CitationBiBTeX, EndNote, RefMan
External Links: Patentscope, Espacenet
X-ray catheter
WO 1997007740 A1
Abstract
A catheter for emitting radiation is disclosed, comprising a catheter shaft (104), and an x-ray unit (102) attached to the distal end of the catheter shaft. The x-ray unit comprises an anode (112), and a cathode (110) coupled to an insulator (108) to define a vacuum chamber (106). The cathode is preferably a field emission cathode of graphite or graphite coated with titanium carbide, for example. The anode is preferably tungsten, and the insulator is preferably pyrolytic boron nitride. The x-ray unit is preferably coupled to a voltage source through a coaxial cable. The anode is preferably a heavy metal such as tungsten. The cathode may also be a ferroelectric material. The x-ray unit can have a diameter less than about 4mm, and a length less than about 15 mm. Methods of use of the catheter are also disclosed. The catheter of the present invention can be used to irradiate the site of an angioplasty procedure to prevent restenosis. It can also be used to treat other conditions in any vessel, lumen or cavity of the body.
Claims  (OCR text may contain errors)
We claim :
1. A catheter for emitting x-ray radiation comprising: a flexible catheter shaft having a distal end; an x-ray unit coupled to the distal end, wherein the x-ray unit comprises an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber.
2. The catheter of claim 1, wherein the cathode is a field emission cathode.
3. The catheter of claim 1, wherein the catheter shaft compriseε a coaxial cable.
4. The catheter of claim 1, wherein the inεulator is chosen from the group consiεting of beryllium oxide, aluminum oxide, or pyrolytic boron nitride.
5. The catheter of claim l, wherein the cathode and the anode are coupled to a voltage generator.
6. The catheter of claim l, further compriεing a guide wire lumen.
7. The catheter of claim 6, wherein the guide wire lumen extendε partially through the catheter εhaft.
8. The catheter of claim 6, wherein the guide wire lumen extendε partially through the x-ray unit.
9. The catheter of claim l, further compriεing a meanε for centering the x-ray unit within a lumen.
10. The catheter of claim 1, wherein the cathode iε a ferroelectric material.
11. An x-ray catheter compriεing: a flexible catheter shaft for being advanced through lumens of the vascular system, the catheter shaft having a distal end; an x-ray unit coupled to the distal end, the x- ray unit compriεing an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber.
12. The catheter of claim ll, wherein the insulator compriseε pyrolytic boron nitride.
13. The catheter of claim ll, wherein the anode compriεeε tungεten or platinum and the cathode comprises graphite.
14. The catheter of claim 11, wherein the cathode is a field emisεion cathode.
15. The catheter of claim 12, wherein the cathode and anode are coupled to a voltage generator.
16. The catheter of claim 15, wherein the catheter εhaft compriεes a coaxial cable coupling the anode and cathode to the voltage generator.
17. The catheter of claim 16, further comprising meanε for centering the x-ray unit within a lumen.
18. A catheter for the emission of x-ray radiation comprising: a flexible catheter shaft having a distal end; an x-ray generating unit coupled to the distal end, the x-ray generating unit compriεing an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber, and wherein the x-ray generating unit haε a diameter less than about 4 mm.
19. The catheter of claim 18, wherein the x-ray generating unit has a diameter of about 1 mm.
20. The catheter of claim 19, wherein the x-ray generating unit haε a length of about 7 mm.
21. The catheter of claim 18, wherein the x-ray generating unit haε a length less than about 15 mm.
22. The catheter of claim 18, wherein the insulator compriseε pyrolytic boron nitride.
23. An x-ray catheter for use in irradiating the wall of a lumen comprising: a flexible catheter shaft having a diεtal end; an x-ray generating unit; and meanε for centering the x-ray generating unit within the lumen.
24. A method for preventing reεtenoεiε of a lumen compriεing:
(a) advancing an x-ray catheter through a lumen to a first location adjacent an intended εite of the lumen, wherein the x-ray catheter comprises a flexible catheter shaft with a distal end and an x-ray generating unit coupled to the distal end, the x-ray generating unit comprising an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber;
(b) causing the emisεion of an effective doεe of x-ray radiation to prevent reεtenoεiε; and
(c) removing the catheter.
25. The method of claim 24, wherein εtep (b) compriεeε cauεing the emission of radiation within a particular energy range to achieve a particular depth of penetration.
26. The method of claim 24, wherein the causing step (b) further compriεeε applying a predetermined voltage between the anode and the cathode to achieve the particular depth penetration.
27. The method of claim 24, further compriεing irradiating tiεsue at a rate of 1-100 grayε per minute.
28. The method of claim 27, wherein the irradiating step is conducted for about l minute.
29. The method of claim 24, wherein step (b) compriseε cauεing the emiεεion of x-rays having an energy of about 8-10 KeV.
30. The method of claim 24, further comprising centering the x-ray unit within the lumen prior to the step
(b) .
31. The method of claim 24, wherein the advancing step comprises advancing the x-ray catheter through a lumen of the vascular system through an exchange tube.
32. The method of claim 24, wherein the advancing step comprises advancing the x-ray catheter through a lumen of the vascular system over a guide wire and through a guide catheter.
33. The method of claim 32, wherein a portion of the x-ray catheter is advanced over the guide wire.
34. The method of claim 24, further compriεing poεitioning the x-ray unit at a εecond location and cauεing the emiεsion of x-ray radiation at the second location. l
35. The method of claim 24, further compriεing poεitioning the x-ray unit at a plurality of locationε and causing the emisεion of x-ray radiation at each of the plurality of locations.
5 36. The method of claim 24, further comprising conducting an angioplasty procedure prior to step (a) , wherein the intended site of step (a) iε the site of the angioplasty procedure.
37. A method for providing x-ray radiation 10 treatment comprising: advancing an x-ray catheter through a lumen to an intended εite, wherein the x-ray unit comprises a flexible catheter shaft with a distal end and an x-ray generating unit coupled to the distal end, the x-ray
15 generating unit comprising an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber; cauεing the emiεεion of an effective dose of x- ray radiation; and
20 removing the catheter.
38. The catheter of claim 2, wherein the cathode is chosen from the group consisting of graphite, titanium carbide, carbides, metals, and graphite coated with titanium carbide.
25 39. The catheter of claim 1, further comprising a guide wire lumen extending through the catheter shaft.
40. The catheter of claim 2, wherein the cathode compriseε silicon and the x-ray unit further comprises a grid proximate the cathode. 30
41. The catheter of claim 2, wherein the cathode compriεeε εilicon needleε.
42. The catheter of claim 11, wherein the x-ray unit irradiateε tiεsue at a rate of at least about 1 gray per minute. 35
43. The catheter of claim 1, wherein the anode is coupled to a wall of the insulator, wherein the wall is tapered towards the anode.
44. The catheter of claim 3, wherein: the coaxial cable comprises an outer conductor and a central conductor; the insulator has a tubular portion with proximal and distal ends, the coaxial cable being coupled to the proximal end, the anode being coupled to the proximal end and to the central conductor of the coaxial cable, and the cathode being coupled to the distal end; the catheter further comprises a conductive surface surrounding the tubular insulator, coupling the cathode to the outer conductor of the coaxial cable; and the inεulator and cathode define an annular region proximate the coupling between the cathode and the insulator, the annular region being εcreened from an electrical field generated between the anode and the cathode by the conductive εurface and a portion of the cathode.
45. The catheter of claim 44, wherein the inεulator compriεes a wall depending from the proximal end of the tubular portion, the wall being angled toward the anode and the vacuum chamber.
Description  (OCR text may contain errors)

X-RAY CATHETER

This application claims the benefit of U.S. Provisional Application Nos. 60/006,708 filed November 14, 1995, and 60/002,722 filed August 24, 1995.

FIELD OF THE INVENTION The present invention relates generally to catheters and, more particularly, to catheters for irradiating vessels, lumens or cavities of a body, such as cardiovascular tissue to reduce the incidence of restenosis, and to treat other conditions.

BACKGROUND OF THE INVENTION Restenosis of an artery or vein after percutaneous transluminal coronary angioplasty (PTCA) or percutaneous transluminal angioplasty (PTA) occurs in about one-third of the procedures, requiring the procedure to be repeated. Various types of drugs or other agents are being investigated for use in preventing restenosis. Heparin, an anticoagulant and inhibitor of arterial smooth muscle proliferation, is one such drug. Dexamethasone may also prevent smooth muscle proliferation. Integralin, which prevents aggregation of platelets, may also be useful. Other anticoagulants and antiproliferative agents are being investigated for efficacy, as well. Such drugs can be delivered before or after the angioplasty procedure. The delivery of lytic agents such as urokinase, streptokinase and recombinant tissue type plasminogen activator (rTPA) to dissolve thrombi in arteries and veins is also being investigated.

Because of blood flow through the artery, drugs delivered to the site of an angioplasty procedure, for example, can be rapidly dissipated and removed from the site before they can be sufficiently absorbed to be effective. Catheters have therefore been developed to directly drive the drug into the desired site through a balloon or to maintain the delivered drug agent proximate the desired site by isolating the region with occlusion balloons. See, for example, U.S. Patent Nos. 5,087,244, 4,824,436, and 4,636,195, to olinsky.

The use of sufficient pressures to drive the drug into the tissue or plague, however, may damage the arterial wall. Passive delivery into a region isolated by occlusions balloons, on the other hand, is slow and may not enable sufficient absorption of the medication. Passive delivery can be particularly inappropriate for drug delivery in an artery because blood flow can only be occluded in an artery for a limited period of time.

Stents have also been used after angioplasty to prevent an opened blood vessel from closing. The use of stents, however, has only shown a small decrease in the incidence of restenosis. Stents are also difficult to properly position and are expensive.

The use of radiation has also been investigated to reduce restenosis after PTCA or PTA. One technique is Photodynamic Therapy (PDT) , wherein photosensitive drugs delivered to the angioplasty site are activated by irradiation with ultraviolet (UV) or visible light.

Another approach was to expose vascular tissue to UV light within a wavelength band of DNA absorption (240-280 nm) by a laser to disable or destroy the DNA of the tissue. This would impair or destroy the ability of the vascular tissue to proliferate. This approach had only limited success, however, because UV light does not penetrate vascular tissue sufficiently to prevent proliferation or migration of smooth muscle tissue.

Beta-irradiation of the vessel after angioplasty with radioactive guide wires or implanted stents is another technique. U.S. Patent No. 5,199,939 to Dake et al. , for example, discloses a catheter with radioactive pellets at its distal end to irradiate the site of an angioplasty procedure to prevent restenosis. The need for a radioactive source in the catheter lab, however, requires protection against radioactive hazards to personnel and costly compliance with regulations. It is also difficult to control the depth of penetration of the radiation by this method. U.S. Patent No. 4,143,275 to Mallozzi et al. , discloses an x-ray device for delivering radiation to remote locations of the human body such as the interior of the heart. The x-ray radiation is generated by irradiating a target material, such as iron, calcium, chromium, nickel, aluminum, lead, tungsten or gold, by a laser to vaporize the metal. X-ray radiation iε emitted from the ionized vapor plasma. The target is located outside the body and the x- rays are directed to a desired location within the body through a hollow guide. The patent discusses use of such a device to produce radiographs, to irradiate tumors or to alter tissue. It is believed, however, that x-ray radiation generated by this method would have photon energy of about 1-2 KeV at best, which is too low to penetrate biological tissue deeper than about 20-30 microns. In addition, the patent does not disclose how to produce a guide which is both flexible enough to be advanced through the cardiovascular system and able to transmit adequate x-ray radiation to an intended site without excessive losses. U.S. Patent No. 5,153,900 to Nomikos, et al. , discloses a miniaturized low power x-ray source for interstitial insertion for the treatment of tumors. The device comprises a housing with an elongated cylindrical, rigid probe. An anode and cathode are located in the housing and a target is located at the distal end of the probe. The cathode and target must lie along the same axis. Electrons emitted by the cathode, which can be a thermionic emitter or a photocathode, impinge on the target, causing the emission of x-ray radiation. A rigid probe is unsuitable for use in the cardiovascular system.

U.S. Patent No. 5,428,658 to Oettinger, et al. , a continuation of the patent to Nomikos, discussed above, discloses a flexible probe comprising a flexible optical fiber within a metallic tube. The optical fiber has a photoemissive coating at its terminal end. A target is located distal to the terminal end of the optical fiber, within an evacuated shell. The flexible probe is said to enable threading down a pathway, such as the trachea, or around structures, such as nerves or blood vessels. Such a device is not sufficiently flexible for advancement through the cardiovascular system, nor is it believed that such a device can be made small enough to access the site of a PTCA procedure. U.S. Patent No. Re 34,421 to Parker, et al. discloses an x-ray microtube comprising a glass tube having a diameter less than one inch, for insertion into the body for treating a tumor. While asserting that the diameter can be as small as 1/8 inch, Parker does not address any of the problems associated with such a small device, such as electrical flashover. It is questionable whether such a device could be made small enough to access the site of a PCT procedure, and still function. Glass also has too high a coefficient of absorption of x-ray radiation to enable delivery of sufficient x-ray radiation to prevent restenosis in a reasonable period of time. Parker also does not disclose any way to advance its x-ray source through the cardiovascular system, or any other channel of the body.

SUMMARY OF THE INVENTION In accordance with a preferred embodiment of the present invention, an x-ray catheter is disclosed which is small and flexible enough to access an intended site within a vascular system of the body, such as the coronary arteries of the cardiovascular system. The x-ray catheter can operate at the high voltages required for generating x-ray radiation of an effective spectrum for preventing restenosis and treating other conditions. It also has walls highly transmissive to x-ray radiation so that an effective dosage can be delivered in a short period of time. in accordance with the present invention, a catheter for emitting x-ray radiation is disclosed comprising a flexible catheter shaft having a distal end and an x-ray unit coupled to the distal end. The x-ray unit comprises an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber. The insulator is preferably pyrolytic boron nitride, which is highly transmissive to x-ray radiation. The cathode is preferably a field emission cathode of graphite, graphite coated with titanium carbide, or other carbides. The cathode can also comprise silicon and the x- ray unit can include a grid. The cathode can be a ferroelectric material, as well. The anode is preferably tungsten. The catheter shaft is preferably a coaxial cable. A guide wire may be provided extending through the catheter shaft, partially through the catheter shaft or partially through the x-ray unit, in a rapid exchange configuration. The catheter further preferably comprises a means for centering the x-ray unit within a lumen. In accordance with another embodiment of the invention, an x-ray catheter is disclosed comprising a flexible catheter shaft for being advanced through lumens of a vascular system.

Another embodiment of the present invention comprises an x-ray generating unit having a diameter less than about 4 mm.

Yet another embodiment of the present invention comprises a catheter shaft, an x-ray generating unit and means for centering the x-ray generating unit within the lumen.

A method is also disclosed in accordance with the present invention for preventing restenosis of a lumen or treating other conditions, comprising advancing an x-ray catheter through a lumen to a first location adjacent an intended site of the lumen, wherein the x-ray catheter comprises a flexible catheter shaft with a distal end and an x-ray generating unit coupled to the distal end. The x-ray generating unit comprises an anode, a cathode and an insulator, wherein the anode and cathode are coupled to the insulator to define a vacuum chamber. The method further comprises causing the emission of an effective dose of x-ray radiation and removing the catheter. The catheter can be inserted after conducting an angioplasty procedure. The catheter can be advanced over a guide wire and through a guide catheter, or through an exchange tube.

DESCRIPTION OF THE FIGURES Fig. IA is a cross-sectional view of an x-ray catheter in accordance with a first embodiment of the present invention; Fig. IB is a cross-sectional view of a preferred catheter shaft for use in the present invention;

Fig. 2A is a graph of an exemplary voltage applied between the anode and grid electrode versus time; Fig. 2B is a graph of an exemplary voltage applied between the grid electrode and rear electrode of the cathode versus time;

Fig. 2C is a graph of the current flow from the cathode to the anode versus time, for the voltages of Figs. 2A and 2B;

Fig. 2D is a graph of the power of the emitted x-ray radiation for the voltages of Figs. 2A and 2B;

Fig. 3A is an alternative cathode in accordance with a second embodiment of the invention; Fig. 3B is an enlarged cross-section of one needle of Fig. 3A;

Fig. 4 is a graph of photon energy versus the Linear Attenuation Coefficient, μ;

Fig. 5 is a cross-sectional view of the distal portion of a third embodiment of the present invention;

Fig. 6 is a cross-sectional view of mandrel for use in chemical vapor deposition of the insulator of the embodiment of Fig. 5;

Fig. 7 is a cross-sectional view of the distal portion of a fourth embodiment of the present invention; Fig. 8 is a cross-sectional view of the distal portion of a fifth embodiment of the present invention;

Figs. 9-11 are side views of the distal portions of the catheter of the present invention, including several centering devices for centering the x-ray unit within a lumen;

Fig. 14 is a cross-sectional view of a distal portion of a catheter in accordance with the present invention, in a rapid exchange configuration wherein the guide wire passes through the distal tip of the x-ray unit; and

Fig. 15 is a partial cross-sectional view of another catheter in accordance with the present invention in a rapid exchange configuration wherein the guide wire enters and exits the catheter shaft proximal to the x-ray unit.

DESCRIPTION OF THE INVENTION

Fig. IA is a cross-sectional view of an x-ray catheter 10 in accordance with a first embodiment of the present invention. The x-ray catheter 10 comprises a flexible catheter shaft 12 adapted for insertion into blood vessels or other body vessels. The shaft 12 can be polyethylene, polyurethane, polyether block amide, nylon 12, polyamide, polyamide copolymer, polypropylene, polyester copolymer, polyvinyl difluoride or silicon rubber, for example.

A miniature x-ray unit 14 is secured at the distal end of the catheter shaft 12 by an adhesive, for example. The x-ray unit 14 comprises a vacuum chamber 16, a cathode 18, which emits electrons, and an anode 20, which receives the emitted electrons. The anode 20 abruptly decelerates the impinging electronε, causing the emission of x-ray radiation by the Bremsstrahlung effect, as is known in the art. About 0.1-0.2% of the kinetic energy of the impinging electrons is emitted in the x-ray range of about 0.5-5 Angstroms in the preferred embodiments of the present invention.

In this embodiment, the anode 20 preferably has the shape of an inverted cone. The walls of the anode 20 preferably have an angle of about 16° with respect to the surface of the cathode 18. The anode 20 is preferably a heavy metal, such as gold or tungsten, for example.

The cathode 18 comprises a base 19 which in this embodiment is preferably a ferroelectric material, as discussed below. The base 19 can also be doped or undoped silicon, or other such materials, which is also discussed below.

A grid electrode 24 is coupled to the surface of the base 19 facing the anode 20. A rear electrode 27 is coupled to the rear of the base 19. Wires 26, 28 and 30 extend from the rear electrode 27, anode 20 and the grid 24, respectively, through the catheter shaft 12, to a high voltage generator 32. The generator 32 preferably operates in the 0-30 kilovolt (Kv) range. The wires 26, 28 and 30 can be soldered in place. Separate lumens 34, 36, 38 can be provided through the catheter shaft 12 for each wire or a single lumen can be provided for a coaxial cable comprising the three wires. A coaxial cable can form the catheter shaft as well, as in the embodiments of Figs. 5 and 7.

The vacuum chamber 16 preferably comprises a wall 22 of beryllium, beryllium oxide, aluminum, aluminum oxide, pyrolytic boron nitride, graphite or other such metal or ceramic materials, which is transparent to x-rays. If a metal, such as beryllium or aluminum is used as the wall 22 of the vacuum chamber 16, an insulative layer (not shown) would be provided to electrically insulate the anode 20 and cathode 18, as iε known in the art. Aluminum oxide, pyrolytic boron nitride and other ceramics are insulators. A transparent biocompatible coating 25 of a polymeric material such as polyethylene, polyurethane or Teflon (R) , for example, is also provided over the wall 22. A vacuum tie off (not shown) depends from the vacuum chamber 16, which is sealed after the desired vacuum within the chamber is achieved. A soft, resilient material 48 may be provided at the distal tip of the x-ray unit 14, as is known in the art. The material can be ultra low density polyethylene or nylon, for example. A lumen 40 extending longitudinally through the catheter shaft 12 can also be provided to accommodate a guide wire 42. A port 44 can be provided through the shaft 12 for the guide wire 42 to exit the shaft 12. A tube 48 can be attached by adhesive or thermal bonding to the shaft 12 at the port 44 to provide a guide for the guide wire 42 around the x-ray unit 14. The tube 48 may be adhered to the wall of the x-ray unit 14, as well. The tube 48 may extend through the soft material 46 at the distal tip of the x-ray unit 14. The lumens in Fig. 1 are shown in the same plane for illustrative purposes. If multiple lumens are provided, they would preferably be arranged symmetrically within the catheter, as shown in Fig. IB. In this embodiment, the base 19 of the cathode 18 is preferably a ferroelectric material, as described in Riege, H., "Electron emission from ferroelectrics - a review," Nuclear Instrumentε and Methods in Physics Research A340 (1994), pp. 80-89; Gundel, H., et al. , "Faεt Polarization Changes in Ferroelectrics and Their Application, " Nuclear Inεtruments and Methods in Physics Research A280 (1989) , pp. 1-6; Gundel, H., et al. , "Time-dependent electron emission from ferroelectrics by external pulsed electric fields, " J. Appl. Phys. 69(2) 15 January 1991, pp. 975-982; and Asano, Jun-ichi, et al. , "Field-Excited Electron Emisεion from Ferroelectric Ceramic in Vacuum," Jpn. J. Appl. Phys. Vol. 31 (1992), pp. 3098-3101, Part 1, No. 9B, which are all incorporated by reference herein. As described in thoεe articleε, ferroelectric materialε, εuch aε lead-zirconium- titanate (PZT) and lead-lanthanum-zirconium-titanate (PLZT) and triglycineεulfate (TGS) , for example, emit electrons from their surfaceε when the εpontaneouε ferroelectric polarization of these materialε iε rapidly reversed. High voltage, submicroεecond pulεeε can cauεe εuch reverεals, aε can mechanical preεεure pulεeε, thermal heating or laεer illumination. The use of a laser to cause polarization reversal is diεcuεεed in Geissler, K., et al. , "Intense laser-induced self-emission of electrons from ferroelectrics," Physics Letters A 176 (1993), pp. 387-392, North Holland, which is also incorporated by reference herein. Ferroelectric cathodeε do not require aε high vacuum aε other types of cathodes. A vacuum of about 10" - 10~4 Torr iε εufficient. Ferroelectric cathodeε are alεo εimple to manufacture and are reliable.

Preferably, the polarization εwitching is caused by applying an electrical pulse acrosε the ferroelectric material. Preferably, voltage pulεeε are applied between the rear electrode 27 and the grid electrode 24. Poεitive or negative pulεeε, or a combination of poεitive and negative pulses, can be used, depending on the configuration and original orientation of the polarization of the ferroelectric material. The reversal of ferroelectric polarization can be achieved by applying a voltage pulεe of between about 1-3 Kv to the ferroelectric cathode 18 via the rear electrode 27 and the grid electrode 24. The pulses are preferably applied for 5-100 nanosecondε. The polarization of the ferroelectric material 19 can be switched at a rate of between about 1 kHz-5 MHz. Electrical current densitieε as high as 100 Amps per square centimeter can be generated. With a polarization switching rate of about 100 kHz, for example, and a diameter of ferroelectric material 19 of about 1 mm, an average anode current of about 10 milliamperes can be generated.

Preferably, a conεtant voltage or voltage pulεes are applied between the anode and the cathode, aε well, to control the energy of the emitted x-ray radiation, and hence the depth of penetration of the radiation into tissue. A voltage of about 10-30 Kv is preferred in coronary applications, aε diεcusεed further, below.

In thiε embodiment, the grid electrode 24 is preferably εilver, aluminum or gold. About one-half of itε area iε transparent or open to electrons. The grid 24 can be deposited on a layer of ferroelectric material, such as

PZT, PLZT or TGS, as is known in the art. The dimensions of the cathode 18 depend on the application. For use in coronary arteries, for example, the ferroelectric material 19 can have a diameter of about 1-2 mm. For uεe in larger blood veεεelε, εuch aε the femoral artery, the diameter of the ferroelectric material 19 could be up to 3 mm. The thickneεε of the ferroelectric material 19 can be between about 50-1,000 micronε. About 200-500 micronε is preferred. The grid 24 is preferably about 0.5-10 microns thick, with about the same diameter as the ferroelectric material 19. The electrode 27 is about 1 micron thick. The distance between the anode 20 and cathode can be about 0.2-5 mm.

Experimental data suggestε that reεtenoεis after PTCA can be limited by irradiation by about 2000 centigrays (cGy) . (See, for example, Tim A. Fiεchel et al. , "Low-Dose, beta-particle emission from "stent" wire resultε in complete, localized inhibition of εmooth muεcle cell proliferation," Circulation, Vol. 90, No. 6, December 1994, and Wiedermann, Joεeph G., et al. , "Intracoronary Irradiation Markedly Reduces Neointimal Proliferation After Balloon Angioplastε in Swine: Perεiεtent Benefit at 6-Month Follow-Up," JACC Vol. 25, No. 6, May 1995, 1451-6, which are incorporated by reference, herein) . It iε believed that the x-ray unit in accordance with this and the other embodiments of the preεent invention diεclosed herein can emit over 2000 centigrays of x-ray radiation in about one minute, to a cylindrical region of a lumen with a length of about 5 mm. Treatment of a typical leεion in a coronary artery, which can be 1-2 centimeterε long, can require repoεitioning of x-ray unit several times to irradiate the entire lesion. A lesion 1-2 centimeters long can therefore be irradiated in about 2-5 minutes. The x-ray catheter of the present invention can deliver sufficient x-ray radiation to a leεion in a εhort period of time which minimizeε the inconvenience and diεcomfort of the patient and cost of the procedure.

In operation, the high voltage generator 32 preferably applieε voltage pulεeε between the anode 20 and grid 24, and between the rear electrode 27 and grid 24. In Fig. 2A, exemplary voltage pulεes applied between the anode 20 and grid 24, VAG, are plotted versus time. The voltage pulses in this example are about 10-12 Kv. The voltage pulses between the anode 20 and grid 24 can be applied for about 0.1-1.0 microseconds, every 10 microseconds. Fig. 2B plots exemplary voltage pulseε VGR, applied between the grid electrode 24 and the rear electrode 27 versus time. The voltage difference here is about 2.0 Kv. Fig. 2B also εhowε a negative pulse 49 which is preferably applied to reεtore the negative charge on the εurface of the ferroelectric material 19 adjacent the grid 24. Fig. 2C illustrates qualitatively the current IA flowing from the ferroelectric material 19 to the anode 20 for the voltage pulseε εhown in Figs. 2A and 2B. The length of each current pulse generated for the range of voltage pulseε of 0.1-1 microεecond, iε about 10-100 nanoεecondε. The current pulses cause the emission of pulseε of x-ray radiation with peak power in this example of up to about 30 watts, as shown in Fig. 2D. In a second embodiment of the invention, shown in

Fig. 3A, the cathode 18 may also be a field emisεion cathode 50 comprising multiple needles 52 and optionally a grid electrode 54. Fig. 3B is an enlarged cross-sectional view of a single needle 52, of Fig. 3A. The base 55 and needles 52 can be doped or undoped silicon. The grid 54 can be niobium. If a grid 54 iε provided, a layer 57 of an insulator, εuch aε silicon dioxide (Si02) / is preferably deposited over the base 55 of εilicon. The grid 54 of niobium iε depoεited over the εilicon dioxide layer 57. A rear electrode 59 iε coupled to the rear of the base 55. A wire 58 iε coupled to the rear electrode 59. A wire 56 is coupled to the grid 54. Returning to Fig. 3A, a vacuum tie- off 60 is shown, as well. The anode 20 can be the same as described above.

The radius of the tips of the needles 52 is between about 5-100 Angstroms. The height of the needles is about 0.5-1.0 microns. The grid 54, which iε about 0.5 micronε thick, iε preferably poεitioned εlightly above the top of the needle 52, aε εhown in Fig. 3B. The openingε in the grid 54 have a diameter of about 2 micronε. The layer of εilicon dioxide is about 1-2 microns thick. A vacuum of between about 10~7-10~8 Torr is preferred for a field emitting cathode including silicon. The needles 52 emit electronε when negative potential is applied between the rear electrode 59 and the grid electrode 54. A triggering voltage of about 100-500 volts may be used, for example. The voltage can be constant or pulsed. If no grid electrode iε provided, the high voltage can be provided directly between the anode and the needles 52.

The radiation emitted by the anode 18 passes through the vacuum chamber wall 22 and coating 25, into εurrounding tiεsue. Irradiation reduces the ability of smooth muscle cell to proliferate, inhibiting restenosis, as discussed above. Fig. 4 is a graph of Photon Energy (kev) versus the Linear Attenuation Coefficient μ (cm-1) for bone 62, muscle 64 and lung tissue 66. (See, Anthony Brinton Wolbarst, Physics of Radiology, Appleton and Lange, 1993, p. 108; Johnε, H.E., Cunningham, JR.: The Phyεicε of Radiology, 4th ed., Springfield, IL; Charleε C. Thomaε, 1983, Appendix A.) The greater the coefficient μ, the more effectively the medium absorbs and scatters photonε. The depth of penetration of radiation iε the depth at which the intenεity of the impinging radiation drops to 1/e of its original value. The depth of penetration of x-ray radiation of a particular energy iε equal to l/μ. Generally, the coefficient μ increaεeε with increaεing effective atomic number of the material. While muεcle and lung tissue have nearly identical chemical composition, the attenuation in muscle tisεue iε about 3 timeε greater than the attenuation in lung tiεεue, becauεe muεcle tiεεue is about 3 times denεer than lung tisεue. The energy of x-ray radiation is preferably adjusted so that it penetrates slightly into the adventitia tisεue of the blood vessel about 2 mm deep. Penetration into the cardiac muscle tissue beyond the coronary artery, for example, should be minimized. The energy can be adjusted by varying the voltage applied between the anode and cathode. The preferred voltage range of 10-30 Kv yields x-ray radiation with a peak energy of about 8-10 KeV, which is appropriate in coronary applicationε.

Fig. 5 iε a croεε-sectional view of the diεtal portion of an x-ray catheter 100 in accordance with a third embodiment of the present invention. The x-ray catheter 100 compriεeε an x-ray unit 102 coupled to a high voltage coaxial cable 104. The x-ray unit 102 has a vacuum chamber 106, defined by an insulator 108, a cathode 110 and an anode 112. The insulator 108 comprises a base portion 114 coupled to a tubular, preferably cylindrical wall portion 116 with an open end 118. The cathode 110, which is a cold, field emiεsion cathode, is coupled to the open end 108. The insulator 108 is preferably alumina, beryllium oxide or more preferably, pyrolytic boron nitride. The boron nitride must be pyrolytic, as opposed to sintered, because only the pyrolytic boron nitride is vacuum tight at the wall thicknesεeε required. The cathode 110 is preferably graphite. The anode 112 is preferably tungεten or tungεten coated with a layer of platinum. A one micron layer of platinum is sufficient. The vacuum is preferably 10~5 Torr or better.

The cathode 110 is preferably graphite, carbides, such as titanium carbide, silicone, metals, or graphite coated with titanium carbide. The cathode 110 preferably includes one or a plurality of protrusionε 110a with a sharp tip extending towardε the anode 112 along a central axiε of the x-ray unit 102. The protruεion 110a locally enhanceε the electrical field and improves the emission of electrons, as iε known in the art. The protruεion 110a can compriεe the same material as the cathode 110, or can be another of the cathode materials mentioned above.

The anode 112, which iε preferably in the εhape of a rod, extendε along the central axiε of the x-ray unit 102. The rod 112 haε a depending portion 112a received within a cylindrical groove 114a extending through the base portion 114. Preferably, the base 114 has a portion 114b, which tapers toward the anode 112. An angle of about 45° can be used, for example. The anode 112 also can have a portion 112b tapered toward the cylindrical portion 114b of the base. Such a configuration displaceε the electrical field from the anode-vacuum-inεulator triple junction, decreaεing the riεk of electrical flashover during operation. The anode 112 is preferably a heavy metal. Tungsten is preferred.

The cathode 110 and anode 114 are coupled to the high voltage generator 32 of Fig. 1, described above, through the high voltage coaxial cable 104. The coaxial cable 104 comprises a central conductor 120, which is coupled to a proximal end of the anode 114, and an external conductor 122, which is coupled to the cathode 110. A conductive coating 124 is provided over the external surface of a portion of the cathode 110 and the external εurface of the inεulator 108 to couple the cathode 110 to the external conductor 122. A εilver coating with a thickneεε of about 0.1-1.0 micronε may be uεed. Gold may be uεed as well. Insulation 126, such as Teflon (R) , silicone, rubber, fluorinated ethylene propylene (FEP) or polyethylene, for example, is typically provided between the external conductor 122 and the central conductor 120. The x-ray unit 102 can be attached to the coaxial cable 114 with an adhesive, for example. The cathode's "triple junction point" (the junction between the cathode, the insulator and the vacuum) , which in this embodiment iε an annular region surrounding the cathode 110 proximate the open end 118 of the insulator 108, is screened from the high electrical field between the anode 112 and the cathode 110 by the conductive coating 124 and the side of the cathode 110. This decreaseε the incidence of electrical flaεhover, enabling the uεe of higher voltageε.

The cathode 110 can be coupled to the open end 118 of the inεulator 108 through a metal ring 130. The metal ring can comprise tungsten, platinum, or graphite covered by platinum. Coupling of the cathode 110 to the metal ring and coupling of the anode 112 to the insulator 108 is described further, below. A biocompatible layer 128 is provided over the external conductor 116, conductive layer 124, and the cathode 110. A thickness of lesε than about 0.002 incheε iε preferred. Preferably, the biocompatible coating 128 alεo actε aε an inεulating layer. The biocompatible coating may be εilicone or FEP, for example. A lubriciouε layer (not shown) of a hyaluronic coating, for example, may be provided aε well. The biocompatible coating mayhave εufficient lubricity without a further coating. Silicone, for example, iε a highly lubriciouε biocompatible coating. The coaxial cable 104 is chosen to have εufficient flexibility to be advanced through the cardiovaεcular or other εuch system, to an intended site. It has been found that εtandard high voltage coaxial cables are generally not flexible enough to be advanced through the cardiovascular system to the coronary arteries. It has further been found, however, that miniature high frequency coaxial cables are available with sufficiently small diameter (about 1.0-3.0 mm outer diameter) and sufficient flexibility to be advanced to the coronary arteries. Usually, such cables are used in high frequency applications at voltages lesε than εeveral kilovolts. Surprisingly, it has been found in connection with the present invention, that these cableε can hold direct current voltages as high as 75-100 Kv without breakdown, and consequently can be used with the x-ray unit of the present invention for operational voltages of up to 30-40 Kv. Such voltages are sufficient to generate x-ray radiation in appropriate energy ranges for the treatment of restenosis and other conditions. Suitable coaxial cables include CW2040-3050FR; CW2040-30; CW2040-3675-SR; and

CW2040-3275SR, diεtributed by Cooner Wire, Inc. Chatεworth, CA, for example. Cooner diεtributeε coaxial cableε for New England Electric Wire Corporation, Lisborn, New Hampshire.

An x-ray unit 102 in accordance with this embodiment of the invention can have a length leεε than about 15 mm and a diameter less than about 4.0 mm, depending on the application. The distance between the cathode 108 and the anode 110 can be between about 2.0-0.2 mm, depending on the size of the x-ray unit 102. The thickness of the cylindrical insulator wall 116 can be between about 0.2-0.5 mm. The diameter of the coaxial cable 104 can be about the same as the diameter of the x-ray unit 102. For use in preventing restenoεiε after dilatation of a coronary artery, which typically haε a diameter of about 3 mm, the x-ray unit 102 preferably haε a length of about 7 mm and a diameter of about 1.5 mm. In peripheral blood veεεelε, which are larger, the x-ray unit 102 preferably has a diameter of about 3.5 mm and a length of between about 7-15 mm. Larger x-ray unitε with greater diameters and lengths than those discuεεed above could also be made and used in accordance with the present invention.

To operate the x-ray unit 101 to prevent restenoεiε in a vessel of the cardiovascular syεtem, for example, direct current having a voltage of between about 10-30 Kv, can be applied to the central conductor 120. The external conductor iε connected to ground. Electrons emitted from the cathode 110 due to a field emisεion effect impact the anode 112, cauεing the emission of x-ray radiation of about 8-10 KeV, as discussed above. The radiation is primarily emitted radially, to the vessel wall. About 10-30 Kv is preferred for uεe in the prevention of reεtenoεiε. Higher voltageε will cauεe the emission of x-ray radiation of higher energy which can penetrate too deeply into the vessel wall, damaging cardiac tissue. Higher voltages may be used for other applications.

Voltageε at the higher end of the 10-30 Kv range are preferred because the uεe of higher voltageε enableε the generation of the same amount of radiation with less current than the use of a lower voltageε, and iε therefore more efficient. Higher voltageε alεo enable the generation of x- ray radiation of higher power. Higher power, however, can cause the generation of more heat, which can damage the tissue of a vesεel wall. In this embodiment, moεt of the heat iε generated at the anode 110 positioned at the center of the x-ray unit, as far from the vesεel wall aε poεεible.

Higher voltage alεo increaεeε the riεk of electrical flaεhover at the anode and cathode triple junctions. As discuεsed above, the anode 112 and cathode 110 are preferably configured to minimize the risk of flashover. Bulk electrical breakdown is also a risk with increaεed voltageε. Pyrolytic boron nitride haε a high dielectric εtrength, enabling the x-ray unit of the catheter to tolerate the voltageε uεed in this application without bulk electrical breakdown.The dielectric strength of pyrolytic boron nitride iε 200-600 KV/mm.

Pyrolytic boron nitride iε also particularly preferred as the inεulator 108 because it is highly transparent to soft x-rays and can therefore be efficiently uεed aε an x-ray window. The coefficient of linear abεorption of boron nitride at about 8 Kev, the average energy of the emitted radiation, iε 1.0 mm"1. About 8-10 KeV iε the preferred energy level of x-ray radiation in the treatment of reεtenoεis, as discuεsed above. Transmisεion of radiation through pyrolytic boron nitride with a thickneεε of about 0.3 mm is about 70%. This enables irradiation of tissue at a rate of at least about 1 gray per minute. Preferably, about 10-30 grays per minute of radiation at about 8-10 KeV are provided, enabling delivery of an effective amount of radiation to prevent reεtenoεiε to a leεion about 5 mm long in about 1 minute. It iε believed that x-ray radiation can be delivered at a rate of up to about 100 grays per minute with the x-ray unit of this embodiment. A lesion 1-2 cm long can be treated in about 2- 5 minutes by progressively repositioning the x-ray unit to irradiate additional portions of the leεion.

Poεitive electrical pulεes with a peak voltage of between about 15-30 Kv and 2-100 nanoseconds long can also be applied to the central conductor 120 of the coaxial cable 104 at a rate of between about 1-50 KHz. The high voltage pulses cause field emission. The pulses can further cause a vacuum electrical breakdown, causing electrons to flow from the cathode 110 to the anode 112 through a plaεma of vaporized cathode and anode material between the cathode 110 and the anode 112.

The anode 114 iε preferably attached to the inεulator 108 of pyrolytic boron nitride during formation of the inεulator 108 by chemical vapor depoεition (CVD) . During CVD, the deposited boron nitride chemically bonds to the anode material, forming a εtrong, vacuum tight seal. The seal formed by CVD has higher voltage hold-off because it does not have voidε which can locally enhance the electrical field and cauεe electrical flaεhover. A mandrel 250 for uεe in manufacturing the x-ray unit 102 by CVD iε εhown in Fig. 6. The mandrel 250 iε preferably graphite. A cavity 252 iε provided in the mandrel 250 for receiving the anode 114. The anode 114 iε εecured in an anode holder 254 of boron nitride, for example. The mandrel 250 includes a shoulder 254 for supporting the metal ring 130. The metal ring 210 is held in place by a cylindrical ring holder 256, alεo of boron nitride, for example, which is εupported by a mandrel holder 258 of graphite, for example. The assembly of Fig. 6 is placed in a CVD reactor for the deposition of boron nitride by CVD, as is known in the art. Chemical vapor deposition of boron nitride is deεcribed, for example, in Matεuda, et al. , "Syntheεiε and Structure of Chemically Vapour-Deposited Boron Nitride, " Journal of Materials Science 21 (1986) pp. 649-658; and Pouch, John J., et al. "Synthesiε Propertieε of Boron Nitride," Materialε Science Forum, Volumeε 54 and 55 (1990) pp. 141-152, for example, which are incorporated by reference, herein. The boron nitride iε deposited on the hot surface of the asεembly, cryεtallizing into a hexagonal εtructure. CVD of pyrolytic boron nitride can be performed by CVD Productε Incorporated, of Hudson, New Hampshire, for example.

It may be advantageous to deposit and impregnate boron onto the surface of the graphite mandrel 250 and tungsten anode 114 prior to depositing the boron nitride. To increaεe the chemical stability of the anode 114 during the deposition procedure, the tungsten could be coated with a layer of platinum about 1 micron thick.

After completion of the CVD process, the mandrel 250 is removed from the assembly by oxidation of the graphite, also as known in the art.

The cathode 110 is then vacuum brazed to the metal ring 130 with brazing materialε, which are discusεed below, sealing the chamber. Vacuum brazing iε also known in the art and can be provided by Koral Labs., Minneapolis, St. Paul, for example. The sealed chamber is then covered with the conductive coating 124 by metal vapor deposition, for example.

Such a procesε can be uεed for aεs production of large numberε of aεsemblies. A fourth embodiment of an x-ray unit 300 in accordance with the preεent invention is shown in Fig. 7. The x-ray unit 300 compriseε a vacuum chamber 302 defined by an insulator 304, preferably of pyrolytic boron nitride, a cathode 306, and an anode 308. The anode 308 is preferably tungsten.

The cathode 306 may be graphite, titanium carbide, graphite coated with titanium carbide or εtainleεs steel, for example. Graphite coated with titanium carbide is preferred. A coating of several microns may be used. Titanium coating can be provided by Lanxide Coated Productε, Inc., Newark, Delaware, for example. The cathode 306 preferably includeε an annular protruεion 306c for creating a cavity for containing the brazing material 316. The cathode 306 may alεo include a protruεion 306a directed towardε the anode 308, aε in the embodiment of Fig. 5.

The inεulator 304 compriεes a cylindrical wall 304a with an inclined depending wall 310 and a cylindrical wall 314 preferably parallel to the cylindrical wall 304a. The depending wall 310 is preferably angled towardε the interior of the vacuum chamber 302. The cylindrical wall 314 defineε a sleeve for receiving a depending portion 318 of the anode 308. The anode 308 is coupled to the cylindrical wall 314 through a brazing alloy 312. The cathode 306 is coupled to the open end 314 of the inεulator 304 through a brazing alloy 316, as well.

The depending portion 318 of the anode 308 preferably includes a slot 320 for receiving the central conductor 322 of a coaxial cable 324. The cathode 306 is coupled to the external conductor 326 of the coaxial cable 324 through a conductive layer 325, as in the embodiment of Fig. 5. A biocompatible coating is also provided over the coaxial cable 324, conductive layer 325 and cathode 306. A lubricious coating (not εhown) may be provided, aε well. Preformed pyrolytic boron nitride of the deεired εizeε and εhapes is available from CVD Products, Incorporated, for example.

Appropriate brazing alloys for coupling pyrolytic boron nitride to the tungsten anode 308 include Incusil-15 ABA and Incusil-ABA, for example, available from GTE Products Corporation, WESTGO Division, Belmont, CA. ("WESTGO") . Incusil-15 ABA compriεes 14.5% indium, 1.25% titanium, 23.5% copper and 60.75% silver. Incusil-ABA comprises 12.5% indium, 1.25% titanium, 27.5% copper and 59% silver. The brazing temperatures for both alloys iε about 750°C. The brazing material can be in the form of a cylindrical ring placed within the εleeve formed by the cylindrical wall 314 in Fig. 7. The brazing material spreads into the vertical region between the anode 308 and wall 314 during the brazing proceεε. Theεe alloyε can also be used to braze the cathode 110 to the metal ring 130 in the embodiment of Fig. 5.

Appropriate brazing alloys for coupling a cathode 308 of graphite or graphite coated with titanium carbide to pyrolytic boron nitride include Cuεin-1 ABA and Cuεil-ABA, alεo available from WESTGO. Cusin-1 ABA comprises 34.25% copper, 1.75% titanium, 1.0% tin and 63% silver. Cusil-ABA comprises 63% silver, 35.25% copper and 1.75% titanium. The brazing temperatures for both alloyε iε about 850°C. The brazing iε alεo conducted in a vacuum of about 10"5 Torr or better. Becauεe it requireε a higher brazing temperature, the graphite cathode 306 iε coupled to the pyrolytic boron nitride prior to the tungεten anode 308. The brazing material can be in the form of a ring or it can be εputtered onto the end of the pyrolytic boron nitride prior to vacuum brazing.

Instead of a cathode of graphite, the cathode can be PLZT or other such ferroelectric material, as discussed above. As above, the uεe of ferroelectric material requireε the uεe of voltage pulεeε. In Fig. 8, a fifth embodiment of the preεent invention iε εhown, compriεing a ferroelectric cathode 130 εupported by a conductive cap 132. The conductive cap 132 is coupled to the outer conductor 116 of the coaxial cable 114 by a conductive layer 118, aε above. The remainder of the x-ray catheter 150 iε the εame aε the embodiment of Fig. 5. Graphite iε preferred aε the conducting material becauεe it haε a low abεorption coefficient for x-ray, enabling transmission through the distal end of the x-ray unit.

It iε preferable to center the x-ray unit within the veεεel or lumen, to provide a uniform diεtribution of x-ray radiation around the circumference of the veεεel wall. Fig. 9 iε a side view of an x-ray catheter 400 in accordance with the preεent invention, with a centering device compriεing a plaεtic εleeve 402 with a plurality of reεilient polymeric εolid armε 404 depending from it at an angle. The sleeve 402 can be coupled to the outer, biocompatible layer of the coaxial cable 406 proximal to the x-ray unit 408 by adhesive or thermal bonding, for example. The diεtal endε of the arms 404 can optionally extend beyond the distal end of the x-ray unit 408. The arms 404 bear against the veεεel wall 410, centering the x-ray unit 408 within a vessel or lumen of the body.

A sheath 412 is preferably provided over the coaxial cable 406 for compressing the arms 404 during advancement of the x-ray unit 408 to the intended site. When the x-ray unit 408 is properly positioned, the εheath 410 iε retracted, releaεing the armε 404. Radiopaque bands 414 of gold or tantalum, for example, are preferably provided on the coaxial cable 406 and the sheath 412 to aεεist in tracking of the x-ray catheter 400 on a fluoroscope during a procedure. The bands 414 are preferably positioned on the coaxial cable 406 and the sheath 412 εuch that when the εheath 412 haε been sufficiently retracted to release the arms 404, the bands on the coaxial cable 406 and the sheath 412 are essentially aligned.

Fig. 10 is a partial, crosε-sectional view of the x- ray catheter 400 of Fig. 9, wherein the x-ray unit 408 is within the sheath 412 and the armε 404 are compreεεed. Saline or some other cooling agent can be delivered through the εpace 416 between the εheath 412 and the coaxial cable 406, as well. Alternatively, a compreεεible cage 418 can be provided over the x-ray unit 408 aε a centering device, aε εhown in Fig. 11. The cage 418 comprises a plurality of arms 420 with a first end 420a coupled to a first sleeve portion 422 and a second end 420b coupled to a εecond εleeve portion 324. The x-ray catheter unit 408 extendε into and lieε within the region defined by the armε 418. The arms 408 can be compreεsed by the sheath 412, as in Fig. 14. The εecond portion 424 can be coupled to the distal end of the x-ray unit 308. The material of the outer layer of the coaxial cable

406 and the material of the sheath 412 preferably comprise materials which slide easily with respect to each other. The outer layer of the coaxial cable 406 iε preferably coated with a lubriciouε material, εuch aε εilicone or a hyaluronic coating, as well.

Releaεable armε and cageε, methodε of their manufacture and suitable materials are discloεed in U.S.S.N. 08/488,216, filed on June 7, 1995 and aεsigned to the .pa assignee of the present inventor. U.S.S.N. 08/488,216 is incorporated by reference, herein.

Another method of centering the x-ray unit is a malecot device, as shown in Figs. 12-13. A sheath 450 of plastic material is attached to the distal portion 454a of an x-ray unit 454, which is εhown in Fig. 12. The coaxial cable 456 attached to the proximal end of the x-ray unit, iε alεo εhown in phantom. A plurality of lateral εlotε 457 are provided through portions of the sheath εurrounding the x- ray unit 454, Four equidiεtantly positioned slots 457 may be provided around the circumference of the sheath 450, two of which are shown in Fig. 12. The length of the slotε 457 dependε on the diameter of the veεsel at the intended site and the diameter of the sheath 450, and should be εufficient to enable the buckled portion of the εheath 450 to bear againεt the circumference of the veεsel wall. When the x- ray unit 454 is adjacent the intended site, the sheath 450 is advanced, causing a portion 458 of the εheath 450 between the slotε 457 to buckle outward, aε εhown in Fig. 13. The sheath 450 is advanced a sufficient diεtance for the portion 458 to buckle sufficiently to bear against the vessel wall, centering the x-ray unit 454. The distal tip 460 of the catheter may be of a εoft, reεilient material such as ultra low denεity polyethylene or nylon, for example, as iε known in the art. Any of the embodimentε of the x-ray catheter can be provided with a εoft tip.

The x-ray unit could alεo be placed within an expandable balloon.

The x-ray catheters of the embodiments of Figs. 5, 7 and 8 can be conveyed to the site of the dilatation procedure through an exchange tube after the dilatation catheter is removed. The exchange tube can be advanced to the intended site over the same guide wire used in the dilatation procedure. After the exchange tube is properly poεitioned, the x-ray catheterε of Figs. 5, 7 and 8 can be advanced through the exchange tube, to the intended site.

The x-ray catheter of the present invention can alεo be advanced over the same guide wire used by the dilatation catheter after the dilatation catheter is removed, through a guide catheter. Fig. 1 shows one such x-ray catheter 10. Fig. 14 is a cross-sectional view of another x-ray catheter 500 for use with a guide wire 502 in a rapid exchange configuration. The guide wire 502 enterε the x-ray unit 504 through an opening 506 in the cylindrical wall of the unit 404, extendε through the center of the unit 504 and a central paεεage 508 in a cathode 510, exiting through an opening at the diεtal end of the unit 504.

The cathode 510 of the x-ray unit 504 may be graphite, for example. The anode can compriεe a baεe 514 of tungεten, for example, with a plurality of rod-like protruεionε 516 arranged concentrically about the base, within a vacuum cavity 518 defined by an inεulator 520 and a cathode 510. The protrusions 516 extend toward the cathode 510. The insulator 520 is preferably of pyrolytic boron nitride. A tube 522 of insulative, vacuum tight material, may be provided through the vacuum chamber 518, providing a pasεage for the guide wire 502.

The baεe 514 of the anode haε a depending portion 514a, preferably coupled to the central electrode 417 of a coaxial cable 518. A conductive layer iε provided over the outer wallε of the insulator 520, to couple the cathode 510 to the outer electrode of the coaxial cable 518, as described in the embodiments, above. Fig. 15 is a side view of another embodiment of a rapid exchange x-ray catheter 600 in accordance with the present invention, wherein a portion of the catheter shaft 602 iε shown in crosε-section. Here, a lumen 601 is provided in the catheter shaft 602 with an entrance port 603 and an exit port 604 proximal to the x-ray unit 605. A guide wire 606 enters the lumen 601 through a port 603 and exits through a port 604. The x-ray catheter 600 can be tracked along the guide wire 606 to the intended site in a lumen or vesεel, through the lumen 601. The distance between the entrance port 602 and the exit port 604 can be about 10-20 cm, for example. Other lumens (not shown) can be provided for a coaxial cable or wires to couple the x-ray unit 605 to the high voltage generator 32 shown in Fig. 1, for example.

Such a catheter shaft 602 can be formed in a multi- lumen extrusion procesε, aε iε known in the art, wherein the lumenε extend longitudinally through the catheter εhaft 602. The portionε of the lumen diεtal and proximal to the intended locations of the exit port 604 and entrance port 602 can be closed, as is known in the art. The ports 603, 604 can then be made through the catheter shaft by a laser, for example.

While the above embodiments are described with respect to applying x-ray radiation to the εite of an angioplaεty procedure, the preεent invention can be uεed to apply radiation within the cardiovaεcular εyεtem for other purpoεeε, or to other vessels, lumens, or cavitieε in the body, wherever the application of radiation would be uεeful. The variouε embodiments set forth above are for the purpoεe of illuεtration. It will be appreciated by thoεe skilled in the art that various changes and modifications may be made to these embodiments without departing from the εpirit and εcope of the invention as defined by the claims, below.

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Classifications
International ClassificationA61N5/10, H01J35/32
Cooperative ClassificationH01J35/32, A61N5/1002, H01J2201/306, H01J2201/304, A61N5/1001, A61N2005/1003
European ClassificationA61N5/10B, H01J35/32
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