WO2002102459A1

WO2002102459A1 - Optically driven therapeutic radiation source

Info

Publication number: WO2002102459A1
Application number: PCT/US2002/019416
Authority: WO
Inventors: Mark Dinsmore
Original assignee: Photoelectron Corporation
Priority date: 2001-06-19
Filing date: 2002-06-19
Publication date: 2002-12-27
Also published as: JP2005502160A; EP1409078A4; CN1596140A; EP1409078A1; JP4268037B2

Abstract

A miniaturized, optically driven therapeutic radiation source (100) includes a laser-heated thermionic cathode (122). A fiber optic cable (113) directs a light beam from an optical source (104) to impinge upon a surface of the thermionic cathode (122), heating the surface to a temperature sufficient to cause thermionic emission of electrons. The target element (128) emits therapeutic radiation, such as x-rays, in response to incident accelerated electrons from the electron beam. In one embodiment, the electron source (122) and the target element (128) are disposed within a capsule (130) which defines a substantially evacuated region. The inner surface of the capsule is coated with a semiconductor coating, so that a uniform voltage gradient is maintained within the evacuated capsule. In another embodiment, the thermionic cathode (300) is formed of a spiral-shaped conductive element (310) having a plurality of spaced-apart turns. Interstitial spacing between adjacent turns essentially eliminates heat transfer across the spacin.

Description

OPTICALLY DRIVEN THERAPEUTIC RADIATION SOURCE

Field of the Invention

The present invention relates to therapeutic radiation sources, and more particularly to a reduced power, increased efficiency miniaturized radiation source that utilizes an optically driven thermionic cathode.

Background of the Invention

In the field of medicine, therapeutic radiation such as x-ray radiation and γ-ray radiation is used for diagnostic, therapeutic and palliative treatment of patients. The conventional medical radiation sources used for these treatments include large, fixed position machines as well as small, transportable radiation generating probes. The current state-of-the-art treatment systems utilize computers to generate complex treatment plans.

Conventional radiation systems used for medical treatment utilize a high power remote radiation source, and direct a beam of radiation at a target area, such as a tumor inside the body of a patient. This type of treatment is referred to as teletherapy because the radiation source is located a predefined distance from the target. This treatment suffers from the disadvantage that tissue disposed between the radiation source and the target is exposed to radiation. Teletherapy radiation sources, which apply radiation to target regions internal to a patient from a source external to the target regions, often cause significant damage not only to the target region or tissue, but also to all surrounding tissue between the entry site, the target region, and the exit site. Brachytherapy, on the other hand, is a form of treatment in which the source of radiation is located close to or in some cases within the area receiving treatment. Brachytherapy, a word derived from the ancient Greek word for close ("brachy"), offers a significant advantage over teletherapy, because the radiation is applied primarily to treat only a predefined tissue volume, without significantly affecting the tissue adjacent to the treated volume. The term brachytherapy is commonly used to describe the use of a radioactive "seed," i.e. encapsulated radioactive isotopes which can be placed directly within or adjacent the target tissue to be treated. Handling and disposal of such radioisotopes, however, may impose considerable hazards to both the handling personnel and the environment.

The term "x-ray brachytherapy" is defined for purposes of this application as x-ray radiation treatment in which the x-ray source is located close to or within the area receiving treatment. An x-ray brachytherapy system, which utilizes a miniaturized low power radiation source that can be inserted into, and activated from within, a patient's body, is disclosed in U.S. Patent No. 5,153,900 issued to Nomikos et al., U.S. Patent No. 5,369,679 to Sliski et al., and U.S. Patent No. 5,422,926 to Smith et al., all owned by the assignee of the present application, all of which are hereby incorporated by reference. The x-ray brachytherapy system disclosed in the above-referenced patents includes a miniaturized, insertable probe which is capable of generating x-ray radiation local to the target tissue, so that radiation need not pass through the patient's skin, bone, or other tissue prior to reaching the target tissue. The insertable probe emits low power x-rays from a nominal "point" source located within or adjacent to the desired region to be affected. In x-ray brachytherapy, therefore, x-rays can be applied to treat a predefined tissue volume without significantly affecting the tissue adjacent to the treated volume. Also, x- rays may be produced in predefined dose geometries disposed about a predetermined location. X-ray brachytherapy offers the advantages of brachytherapy, while avoiding the use and handling of radioisotopes. Also, x-ray brachytherapy allows the operator to control over time the dosage of the delivered x-ray radiation.

X-ray brachytherapy typically involves positioning the insertable probe into or adjacent to the tumor, or into the site where the tumor or a portion of the tumor was removed, to treat the tissue adjacent the site with a local boost of radiation. X-ray probes of the type generally disclosed in U.S. Patent No. 5,153,900 include a housing, and a hollow, tubular probe or catheter extending from the housing along an axis and having an x-ray emitting target at its distal end. The probe may enclose an electron source, such as a thermionic cathode. In another form of an x-ray brachytherapy device, as disclosed in U.S. Patent No. 5,428,658, an x-ray probe may include a flexible probe, such as a flexible fiber optic cable enclosed within a metallic sheath. The x-ray probe may also include a substantially rigid capsule that is coupled to a distal end of the flexible probe. The capsule encloses an electron source and an x-ray emissive target element. The electron source may be a photocathode. In a photocathode configuration, a photoemissive substance is irradiated by a LED or a laser source, causing the generation of free electrons. Typically, the flexible fiber optic cable couples light from a laser source or a LED to the photocathode.

In the devices disclosed in U.S. Patents Nos. 5,133,900 and 5,428,658, an accelerating electric field may be established between the electron source and the target element. The established electric field acts to accelerate the electrons emitted from the electron source toward the target element. The target element emits radiation in response to incident electrons from the electron source.

In one form of a conventional thermionic cathode, a filament is heated resistively with a current. This in turn heats the cathode so that electrons are generated by thermionic emission. In one form of a conventional x-ray machine that uses such resistively heated thermionic cathodes, the cathode assembly may consist of a thoriated tungsten coil approximately 2 mm in diameter and 1 to 2 cm in length. When resistively heated with a current of 4 A or higher, the thoriated tungsten coil thermionically emits electrons. In one configuration, this coil is surrounded by a metal focusing cup which concentrates the beam of electrons to a small spot on an opposing anode which also functions as the target. The beam is focused on the anode to a spot diameter, usually ranging anywhere from about 0.3 to 2.5 millimeters. In many applications, most of the energy from the electron beam is converted into heat at the anode. To accommodate such heating, high power medical x-ray sources often utilize liquid cooling and a rapidly rotating anode. An increased effective target area is thereby established, permitting a small focal spot while minimizing the effects of localized heating. To achieve good thermal conductivity and effective heat dissipation, the anode typically is fabricated from copper. In addition, the area of the anode onto which an electron beam is incident must be made from a material of high atomic number, in order for x-rays to be generated efficiently. To meet the requirements of thermal conductivity, effective heat dissipation, and efficient x-ray generation, a tungsten alloy is typically embedded in the copper. It is desirable that the electron source be heated as efficiently as possible, namely that the thermionic cathode reach as high a temperature as possible using as little power as possible. In conventional x-ray tubes, for example, thermal vaporization of the tube's coiled cathode filament is frequently responsible for tube failure. Also, the anode heated to a high temperature can cause degradation of the radiation output. During relatively long exposures from an x-ray source, e.g. during exposures lasting from about 1 to about 3 seconds, the anode temperature may rise sufficiently to cause it to glow brightly, accompanied by localized surface melting and pitting which degrades the radiation output.

While a photocathode avoids such problems, one disadvantage of using a photocathode is the difficulty in fabricating the photocathode. A photocathode must have a sufficient quantum efficiency, where quantum efficiency relates to the number of electrons generated per incident light quantum. The degree of efficiency must be balanced to the intensity of available incident light. For practical substances, with reasonable quantum efficiencies above 10^" , the fabπcation of the photocathode should be performed in a vacuum. As disclosed in U.S. Pat. No. 5,428,658, owned by the assignee of the present application and hereby incorporated by reference, in one form the vacuum fabrication can be carried out with the fiber optic cable positioned in a bell jar. By way of example, an Ag-O-Cs photosurface can be fabricated in the conventional manner. Subsequently, without exposure to air, the fiber optic cable can be inserted into the tubular probe, and the end of the fiber optic cable can be vacuum sealed to the probe.

It is an object of this invention to provide an increased efficiency, miniaturized radiation source having significantly reduced power requirements. It is another object of this invention to provide a miniaturized radiation source in which the electron source can generate electrons with minimal heat loss, without requiring a vacuum-fabricated photocathode. It is yet another object of this invention to provide a miniaturized radiation source in which laser energy is used to heat a thermionic cathode, instead of heating a thermionic cathode via conventional ohmic heating. In this way, electrons can be produced in a quantity sufficient to form an electron current necessary for generating therapeutic radiation at the target, while significantly reducing the requisite power requirements for the radiation source.

In order to reduce the power requirements for the laser-heated therapeutic radiation source discussed above, it is necessary to minimize heat loss by the thermionic cathode. Heat loss in a laser-heated thermionic cathode includes 1) heat lost by thermal conduction; 2) heat loss caused by the portion of the incident laser radiation that remains unabsorbed; and 3) heat loss by thermal radiation.

It is another object of this invention to increase the efficiency of a laser-heated thermionic cathode in a radiation source, by reducing the amount of heat that is lost due to incident laser radiation that remains unabsorbed by the thermionic cathode. It is yet another object of this invention to reduce heat loss that is caused by thermal conduction in a laser heated thermionic cathode, thereby further increasing the efficiency of a laser-driven therapeutic radiation source and reducing the power requirements therefor.

In the devices disclosed in U.S. Patents Nos. 5,133,900 and 5,428,658, the electron source and the target element are enclosed within a substantially rigid capsule. The electron source generates an electron beam along a beam path, and the target element is positioned in the beam path. An accelerating electric field may be established within the capsule. The accelerating electric field acts to accelerate the electrons emitted from the electron source toward the target element. The target element emits therapeutic radiation in response to incident electrons from the electron source.

The capsule defines a substantially evacuated interior region extending along the electron beam axis. Typically, the inner surface of the capsule is lined with an electrical insulator. Although the vacuum is used extensively for the insulation of high voltages in devices such as the x-ray probes described above, the reliability of the vacuum is limited by the operational risk of an unpredictable "sparking" or "arcing" between the electrodes, when the insulating capability of the vacuum gap is suddenly lost and electrical breakdown is said to have occurred. Also, the efficient production of x-rays requires that the electron path be directly from the cathode to the target. If the electrons are deflected to the walls by effects of insulator charging, the efficiency of x-ray production is reduced, and stability of the x-ray output is compromised. It is therefore important to establish a substantially uniform voltage gradient in the region between the electron source and the target, in order to avoid such electrical breakdown and to maximize and stabilize the x-ray output. It is yet another object of this invention to provide a high efficiency, miniaturized therapeutic radiation source having a substantially uniform voltage gradient within the vacuum region between the electron source and the target.

Summary of the Invention

The present invention is directed to a miniaturized source of therapeutic radiation having a low power, electron-beam activated radiation source. In particular, the apparatus of the present invention includes a thermionic cathode heated by a source of optical radiation, preferably a laser.

A therapeutic radiation source in accord with the present invention includes a radiation generator assembly, a source of optical radiation, and a probe assembly. The source of optical radiation is preferably a laser that generates a substantially monochromatic, coherent beam of radiation. The radiation generator assembly includes an electron source for emitting electrons to generate an electron beam along a beam path along a nominally straight reference axis, and a target element positioned in the beam path. The electron source is preferably an optically driven thermionic cathode having an electron emissive surface.

In one embodiment of the invention, the optically-driven thermionic cathode is spiral- shaped. In this way, heat loss due to thermal conduction within the thermionic cathode is minimized. In this embodiment, the spiral-shaped thermionic cathode is preferably made of a spiral-shaped conductive element. The spiral-shaped conductive element has a plurality of spaced apart turns, and defines an interstitial spacing between each successive turn of said conductive element. Because the spiral-shaped conductive element is enclosed within the substantially evacuated interior region, heat transfer across the interstitial spacing between each adjacent turn of the conductive element is essentially eliminated. By minimizing heat lost by thermal conduction, the efficiency of the miniaturized thermionic cathode is increased.

The target element includes means for emitting therapeutic radiation in response to incident accelerated electrons from said electron beam. In a preferred embodiment, the target element is spaced apart from and opposite the electron emissive surface of the thermionic cathode. The target element includes at least one radiation emissive element adapted to emit therapeutic radiation in response to incident accelerated electrons from the thermionic cathode. The therapeutic radiation source also includes means for providing an accelerating voltage so as to establish an accelerating electric field which acts to accelerate electrons emitted from said electron source toward the target element. In one embodiment, the radiation generator assembly further includes a substantially rigid capsule which encloses the electron source and the target element. Preferably, the electron source is at its proximal end, and the target element is at its distal end. The capsule defines a substantially evacuated interior region extending along the nominally straight beam axis, between the thermionic cathode at the proximal end of the capsule and a target element at the distal end of the capsule. The capsule preferably includes a radiation transmissive region, which may be disposed at the distal end of the capsule. The total resistance of the inner surface of the capsule is preferably high enough to limit dissipated power to less than 10 % total target power. In one embodiment of the invention, the inner surface of the evacuated capsule is coated with a weakly conductive or semiconductive coating to provide a substantially smooth voltage gradient within the capsule, between the preselected maximum value and the ground potential. The weakly conductive or semiconductive coating, applied to the inner surface of the capsule, is also adapted to reduce secondary emissions of electrons striking the coated inner surface of the capsule. The weakly conductive or semiconductive coating is further adapted to reduce the electrical field in the vicinity of the triple junction point, thus reducing the possibility of electrical flashover the triple junction point of the thermionic cathode. Sufficient current is carried in the coating to prevent charge buildup from field emission, and subsequent avalanche and breakdown.

The probe assembly includes an optical delivery structure, preferably a fiber optic cable, having a proximal end and a distal end. The distal end of the fiber optic cable is coupled to the radiation generator assembly. The fiber optic cable transmits optical radiation, generated by the source and incident on the proximal end, to the distal end. The fiber optic cable directs a beam of the transmitted optical radiation to impinge upon a surface of the thermionic cathode, wherein the beam of optical radiation has a power level sufficient to heat at least a portion of a surface of the thermionic cathode to an electron emitting temperature so as to cause thermionic emission of electrons from the surface. In one embodiment, the probe assembly includes a flexible metallic sheath enclosing the fiber optic cable.

In one embodiment, the means for providing the accelerating voltage is a power supply having a first terminal and a second terminal, and having drive means for establishing an output voltage between the first terminal and the second terminal. In one form, the power supply may be electrically coupled to the target element by way of the first and second terminals. The first terminal of the power supply can be electrically coupled to the electron emissive surface of the thermionic cathode, and the second terminal electrically coupled to the target element, thereby establishing an electric field which accelerates electrons emitted from the thermionic cathode toward the target element. In one embodiment of the invention, the apparatus of the present invention includes one or more reflector elements disposed at predetermined locations along an inner surface of the housing. The reflector elements are operative to reflect incident laser radiation unabsorbed by the thermionic cathode back to the thermionic cathode, thereby increasing the efficiency of the therapeutic radiation source.

Brief Description of the Drawings

FIGURE 1 (a) is a diagrammatic perspective view of a therapeutic radiation source having a resistively heated thermionic cathode. FIGURE 1 (b) is a schematic representation of a therapeutic radiation source having a resistively heated thermionic cathode.

FIGURE 2 is a schematic block diagram of an overview of a therapeutic radiation source constructed according to the present invention, and having a laser-heated thermionic cathode. FIGURE 3 is a diagrammatic view of one embodiment of a source of therapeutic radiation constructed according to the present invention, illustrating a laser source, a probe assembly and a radiation generator assembly embodying the present invention.

FIGURE 4 is an enlarged diagrammatic view of one embodiment of a probe assembly and a radiation generator assembly, constructed according to the present invention.

FIGURE 5 is an enlarged view of one end of a radiation generator assembly embodying the present invention, illustrating an electron source having a laser-heated thermionic cathode. FIGURE 6 is an enlarged view of one embodiment of an electron source embodying the present invention, illustrating reflector elements which reflect back to the laser-heated thermionic cathode laser radiation that was unabsorbed by the thermionic cathode.

FIGURE 7 shows an embodiment of the present invention in which the therapeutic radiation source includes a spiral-shaped, laser-heated thermionic cathode.

FIGURE 8 (a) shows a plane view spiral-shaped thermionic cathode, constructed in accordance with the present invention.

FIGURE 8 (b) shows a side view of a spiral-shaped thermionic cathode, constructed in accordance with the present invention. FIGURE 9 provides an enlarged view of a radiation generator assembly, and the distal end of the probe assembly, constructed in accordance with one embodiment of the present invention in which a weakly conductive or semiconductive coating is applied to the inner surface of the rigid capsule that encloses the electron source and the target element.

FIGURE 10 (a) illustrates an enlarged view of the field lines for the voltage gradient within an evacuated capsule. FIGURE 10 (b) illustrates the triple junction point of the thermionic cathode used in the present invention.

Detailed Description The present invention is directed to a miniaturized, low power therapeutic radiation source which can be used for diagnostic, therapeutic and palliative treatment of patients. In the present invention, a laser is used to heat a thermionic cathode to an electron emitting temperature. The power requirements for the therapeutic radiation source are significantly reduced, as compared to systems in which the thermionic cathode is resistively heated. Therapeutic radiation generated by the apparatus of the present invention may include, but is not limited to, x-rays. In medical applications, the apparatus may be fully or partially implanted into, or surface mounted onto a desired area of a host, so as to irradiate a pre-selected region with therapeutic radiation. The apparatus of the present invention can operate at a relatively low voltage, for example in the range of approximately 10 keV to 90 keV, with electron currents for example in the range of from approximately 1 nA to about 1 μA.

FIGURE 1 (a) shows a therapeutic radiation source 10 that generates and delivers therapeutic radiation in the form of x-rays. The miniaturized, low power x-ray source 10 shown in FIGURE 1 (a) is a prior art x-ray brachytherapy system having a thermionic cathode that is heated using conventional resistive heating. A suitable system is described in detail for example in the above-referenced U.S. Pat. No. 5,153,900, entitled "Miniaturized Low Power X-Ray Source." The system includes a housing 12 and an elongated, cylindrical probe 14 extending from the housing along a reference axis 16 and having a target assembly 26 at its distal end. The probe 14 may be flexible or rigid, and is integral with the housing 12. The housing 12 encloses a high voltage power supply 12 A. The probe 14 is a hollow tube, and encloses an electron source 20. The electron source 20 includes a thermionic cathode 22 that may be driven by a floating low voltage power supply. In one embodiment, the electron source 20 may also include an annular focusing electrode 23, in which case the thermionic cathode 22 is located in proximity to the annular focusing electrode 23, which is typically at nearly the same potential as the cathode. The probe 14 extends along the same axis as the cathode 22 and the focusing electrode 23. The probe 14 may be a hollow evacuated cylinder made of a beryllium (Be) cap and molybdenum-rhenium (Mo-Re), molybdenum (Mo) or mu-metal body. The length of the probe 14 may be determined in view of the body region to be treated. For example, the cylinder may be 15 cm in length, with an interior diameter of 4 mm, and an exterior diameter of 5 mm. Different geometries of the probe 14 may be used for different body regions. The main body of the probe 14 can be made of a magnetically shielding material such as a mu-metal. Alternatively, the probe 14 can be made of a non-magnetic metal, preferably having relatively high values for Young's modulus and elastic limit. Examples of such material include molybdenum, rhenium, or alloys of these materials. The inner or outer surface of probe can then be coated with a high permeability magnetic alloy such as permalloy (approximately 80% nickel and 20% iron), to provide magnetic shielding. Alternatively, a thin sleeve of mu-metal can be fitted over, or inside of, the probe 14. The prior art x-ray apparatus 10 can then be used in environments in which there are dc and ac magnetic fields due to electrical power, the field of the earth, or other magnetized bodies nominally capable of deflecting the electron beam from the probe axis. FIGURE 1 (b) is a schematic representation of the prior art x-ray source apparatus 10 shown in FIGURE 1 (a). In this schematic representation, the housing 12 is represented as being divided into a first portion 12' and a second portion 12". Enclosed within the first housing portion 12' is a rechargeable battery 12B, a recharge network 12D for the battery, which is adapted for use with an external charger 50, and a telemetry network 12E, adapted to be responsive to an external telemetry device 52 to function. The first housing portion 12' is coupled by suitable communication means to the second housing portion 12". The second housing portion 12" includes a high voltage power supply 12 A, the controller 12C, and the probe 14, as well as an electron beam generator. In the illustrated prior art apparatus, the electron beam generator includes a photocathode 22 driven by an associated light source driver 55 and diode laser 56 and associated lens assembly 58. In operation, the laser 56 illuminates the photocathode 22, which in turn generates electrons which are then accelerated toward the anode 24. The anode 24 attracts the electrons, then passes them through its central aperture toward the target assembly 26. The microprocessor 12C controls the power supply 12A and the light source driver 55 to dynamically adjust the cathode voltage, the electron beam current, and temporal parameters, or to provide pre-selected voltage, beam current, and temporal parameters.

As illustrated in FIGURE 1 (b), the external telemetry device 52 and telemetry network 12E may cooperate to permit external control, either dynamic or predetermined, over the power supply 12A and temporal parameters. Alternatively, the housing 12" may not be implanted into a host, and only the probe 14 may extend into a patient's body. In this case, the controller 12C may be used directly to control the operation of the apparatus, and there is no need for the network 12E.

In the above-described prior art x-ray source 10, the x-ray emissive element of the target

26 is adapted to be adjacent to or within the region to be irradiated. The proximity of the emissive element to the targeted region, e.g. the tumor, eliminates the need for the high voltages of prior art machines, in order to achieve satisfactory x-ray penetration through the body wall to the tumor site. The low voltage also concentrates the radiation in the targeted tumor, and limits the damage to surrounding tissue and surface skin at the point of penetration.

FIGURE 2 is a schematic block diagram of an overview of one embodiment of a therapeutic radiation source 100, constructed according to the present invention. The therapeutic radiation source 100 includes a laser-heated thermionic cathode, in contrast to the prior art therapeutic radiation sources (shown in FIGURES 1 (a) and 1 (b)), which included a resistively heated thermionic cathode, or a photocathode. Heating the thermionic cathode 122 with a laser, instead of a current, significantly reduces the power requirements for a therapeutic radiation source 100 constructed in accordance with the present invention. In overview, the therapeutic radiation source 100 includes a radiation generator assembly

101, a source 104 of optical radiation, and a probe assembly 106. Preferably, the source 104 of optical radiation is a laser, so that the optical radiation generated by the source 104 is substantially monochromatic, and coherent. The laser may be a diode laser, by way of example; however other lasers known in the art may be used, such as a Nd:YAG laser, a Nd: YVO₄ laser, or a molecular laser. Alternatively, other sources of high intensity light may be used, such as LEDs (light emitting diodes).

The radiation generator assembly 101 includes an electron source 122, and a target element 128 that includes means for emitting therapeutic radiation in response to incident accelerated electrons. Preferably, the electron source 122 is a thermionic cathode 122. The probe assembly 106 includes an optical delivery structure 113, such as a fiber optic cable. The optical delivery structure 113 directs a beam of laser radiation generated by the laser source 104 onto the thermionic cathode 122. The laser beam heats the thermionic cathode 122 so as to cause thermionic emission of electrons.

FIGURES 3 and 4 illustrate a diagrammatic view of one embodiment of the therapeutic radiation source 100 constructed according to the present invention. In the embodiment illustrated in FIGURE 3, the therapeutic radiation source 100 includes a laser source 104, a probe assembly 106, and a radiation generator assembly 101. The radiation generator assembly 101 includes an electron source 122 that generates an electron beam along a beam path 109, and a target element 128 positioned in the beam path 109. The therapeutic radiation source 100 also includes means for providing an accelerating voltage between the electron source 122 and the target element 128. In the illustrated embodiment, the means for providing the accelerating voltage is a high voltage power supply 112. The probe assembly 106 couples the laser source 104 and the high voltage power supply 112 to the radiation generator assembly 101. FIGURE 3 provides an overall view of the therapeutic radiation source 100, whereas FIGURE 4 provides an enlarged view of 1) the radiation generator assembly 101, and 2) the distal end of the probe assembly 106.

Referring to both FIGURES 3 and 4, the probe assembly 106 includes an optical delivery structure 113 having a proximal end 113A and a distal end 113B. The optical delivery structure 113 is enclosed within a flexible, electrically conductive catheter 105. The distal end 113B of the optical delivery structure 113 is affixed to the radiation generator assembly 101. In a preferred embodiment, the optical delivery structure 113 is a flexible fiber optic cable, extending from the proximal end 113A to the distal end 113B. In this embodiment, the flexible catheter 105 that encloses the fiber optic cable 113 is a small-diameter, flexible, metallic sheath. In a preferred embodiment, the fiber optic cable 113 includes an electrically conductive outer surface 200. For example, the outer surface of the fiber optic cable 113 may be made conductive by applying an electrically conductive coating. The electrically conductive outer surface 200 of the fiber optic cable 113 provides a connection to the thermionic cathode 122 from the high voltage power supply 112. In this embodiment, the radiation generator assembly 101 also has an electrically conductive outer surface. Preferably, both the flexible metallic sheath 105 and the outer conductive surface of the radiation generator assembly 101 are set at ground potential, in order to reduce the shock hazard of the device. The flexible sheath 105 couples a ground return from the target element 128 to the high voltage power supply 112, thereby establishing a high voltage field between the thermionic cathode 122 and the target element 128. In an exemplary embodiment, the fiber optic cable 113 may have a diameter of about 200 microns, and the flexible metallic sheath 105 may have a diameter of about 1.4 mm. A layer 210 of dielectric material provides insulation between the outer surface of the fiber optic cable 113 and the inner surface of the metallic sheath 105.

As shown in FIGURES 3 and 4, the radiation generator assembly 101 includes the electron source 122, and the target element 128. The radiation generator assembly 101, which can be for example about 0.5 to about 2 cm in length, extends from the distal end of the probe assembly 106 and includes a shell or capsule 130 which encloses the electron source 122 and the target element 128. According to one embodiment, the capsule 130 is rigid in nature and generally cylindrical in shape. In this embodiment, the cylindrical capsule 130 enclosing the other elements of the radiation generator assembly 101 can be considered to provide a substantially rigid housing for the electron source 122 and the target element 128. In this embodiment, the electron source 122 and the target element 128 are disposed within the housing 130, with the electron source 122 disposed at a proximal end of the capsule 130, and the target element 128 disposed at a distal end of the capsule 130. The capsule 130 defines a substantially evacuated interior region extending along the beam axis 109, between the electron source 122 at the proximal end of the capsule 130 and the target element 128 at the distal end of the capsule 130. The inner surface of the radiation generator assembly 101 is lined with an electrical insulator or semiconductor, while the external surface of the assembly 101 is electrically conductive, as mentioned earlier. A low secondary emission, controlled sheet resistance semiconducting film maximizes the high voltage breakdown voltage of the system. According to a preferred embodiment, the radiation generator assembly 101 is hermetically sealed to the end of the probe assembly, and evacuated. According to another embodiment, the entire probe assembly 106 is evacuated. In the illustrated preferred embodiment of the invention, the electron source 122 is a thermionic cathode 122 having an electron emissive surface. In an alternative form of the invention (not shown), an annular focusing electrode may also be provided. In the alternative embodiment, the thermionic cathode 122 may be located in close proximity to the annular focusing electrode, which may be at nearly the same potential as the cathode. In the embodiments illustrated in FIGURES 3 and 4, the means for establishing an accelerating electric field is the high voltage power supply 112. The power supply 112 has a first terminal 112A and a second terminal 112B, and has drive means for establishing an output voltage between the first terminal 112A and the second terminal 112B. In one form, the power supply 112 may be electrically coupled to the target element by way of the first and second terminals. The first terminal 112A of the power supply 112 is electrically coupled to the electron emissive surface of the thermionic cathode 122, and the second terminal 112B is electrically coupled to the target element 128.

In the illustrated embodiment, the high voltage power supply 112 provides a high potential difference across the conductive outer surface 200 of the fiber optic cable, and the metallic sheath 105, to establish an acceleration potential difference between the thermionic cathode 122 and the grounded target element 128. In this way, electrons emitted from the thermionic cathode 122 are accelerated toward the target element 128, and an electron beam is generated. The electron beam is preferably thin (e.g. 1 mm or less in diameter), and is established along a beam path 109 along a nominally straight reference axis that extends to the target element 128. The target element 128 is positioned in the beam path 109. The distance from the electron source 122 to the target element 128 is preferably less than 2 mm.

The high voltage power supply 112 preferably satisfies three criteria: 1) small in size; 2) high efficiency, so as to enable the use of battery power; and 3) independently variable x-ray tube voltage and current, so as to enable the unit to be programmed for specific applications. Preferably, the power supply 112 includes selectively operable control means, including means for selectively controlling the amplitude of the output voltage and the amplitude of the beam generator current. A high-frequency, switch-mode power converter is preferably used to meet these requirements. The most appropriate topology for generating low power and high voltage is a resonant voltage converter working in conjunction with a high voltage, Cockroft- Walton-type multiplier. Low-power dissipation, switch-mode power-supply controller-integrated circuits (IC) are currently available for controlling such topologies with few ancillary components. A more detailed description of an exemplary power supply suitable for use as the power supply 112 is provided in U.S. Patent Nos. 5,153,900 and 5,428,658.

The target element 128 is preferably spaced apart from and opposite the electron emissive surface of the thermionic cathode 122, and has at least one radiation emissive material adapted to emit therapeutic radiation in response to incident accelerated electrons from the electron emissive surface of the thermionic cathode 122. In a preferred embodiment, the emitted therapeutic radiation consist of x-rays, however it should be noted that the scope of this invention is not limited to x-rays, and other forms of therapeutic radiation may also be generated. In one embodiment, the target element 128 is a small beryllium (Be) substrate, coated on the side exposed to the incident electron beam with a thin film or layer of a high-Z, x-ray emissive element, such as tungsten (W), uranium (U) or gold (Au). By way of example, when the electrons are accelerated to 30 keV-, a 2 micron thick gold layer absorbs substantially all of the incident electrons, while transmitting approximately 95% of any 30 keV-, 88% of any 20 keV-, and 83% of any 10 keV-x-rays generated in that layer. In this embodiment, the beryllium substrate is 0.5 mm thick. With this configuration, 95% of the x-rays generated in directions normal to and toward the beryllium substrate, and having passed through the gold layer, are then transmitted through the beryllium substrate and outward at the distal end of the probe assembly 106. In some forms of the invention, the target element 128 may include a multiple layer film, where the differing layers may have different emission characteristics. By way of example, the first layer may have an emission versus energy peak at a relatively low energy, and the second underlying layer may have an emission versus energy peak at a relatively high energy. With this form of the invention, a low energy electron beam may be used to generate x-rays in the first layer, to achieve a first radiation characteristic, and high energy electrons may be used to penetrate through to the underlying layer, to achieve a second radiation characteristic. As an example, a 0.5 mm wide electron beam may be emitted at the cathode and accelerated to 30 keV, with 0.1 eV transverse electron energies, and may arrive at the target element 128, with a beam diameter of less than 1 mm at the target element 128. X-rays are generated in the target element 128 in accordance with pre-selected beam voltage, current, and target element composition. The x-rays thus generated pass through the beryllium substrate with minimized loss in energy. As an alternative to beryllium, the target substrate may be made of carbon, ceramic such as boron nitride, or other suitable material which permits x-rays to pass with a minimum loss of energy. An optimal material for target substrate is carbon in its diamond form, since that material is an excellent heat conductor. Using these parameters, the resultant x-rays have sufficient energy to penetrate into soft tissues to a depth of a centimeter or more, the exact depth dependent upon the x-ray energy distribution. In another embodiment of the invention, the target may be a solid, high-Z material, with x-rays being emitted in an annular beam perpendicular to the tube axis. In the above embodiments, the probe assembly 106, along with its associated radiation generator assembly 101, can be coated with a biocompatible outer layer, such as titanium nitride on a sublayer of nickel. For additional biocompatibility, a sheath of, for example, polyurethane can be fitted over the probe.

FIGURE 5 illustrates an electron source constructed according to the present invention, and including a laser-heated thermionic cathode 122. The cathode disc can be held in place by means of swage of the end or by laser welding. The thermionic cathode 122 has an electron emissive surface, and is typically formed of a metallic material. Suitable metallic materials forming the cathode 122 may include tungsten, thoriated tungsten, other tungsten alloys, thoriated rhenium, and tantalum. In one embodiment, the cathode 122 may be formed by depositing a layer of electron emissive material on a base material, so that an electron emissive surface is formed thereon. By way of example, the base material may be formed from one or more metallic materials, including but not limited to Group VI metals such as tungsten, and Group II metals such as barium. In one form, the layer of electron emissive material may be formed from materials including, but not limited to, aluminum tungstate and scandium tungstate. The thermionic cathode 122 may also be an oxide coated cathode, where a coating of the mixed oxides of barium and strontrium, by way of example, may be applied to a metallic base, such as nickel or a nickel alloy. The metallic base may be made of other materials, including Group VI metals such as tungsten.

Getters 155 may be positioned within the housing 130. The getters 155 aid in creating and maintaining a vacuum condition of high quality. The getter has an activation temperature, after which it will react with stray gas molecules in the vacuum. It is desirable that the getter used have an activation temperature that is not so high that the x-ray device will be damaged when heated to the activation temperature.

The fiber optic cable 113 is adapted to transmit laser radiation, generated by the laser source 104 (shown in FIGURE 3) and incident on the proximal end 113A of the fiber optic cable 113, to the distal end 113B of the fiber optic cable 113. The fiber optic cable 113 is also adapted to deliver a beam of the transmitted laser radiation to impinge upon the electron-emissive surface of the thermionic cathode 122. The beam of laser radiation must have a power level sufficient to heat at least a portion of the electron-emissive surface to an electron emitting temperature so as to cause thermionic emission of electrons from the surface. In operation, the laser beam shining down the fiber optic cable 113 impinges upon the surface of the thermionic cathode 122, and rapidly heats the surface to an electron emitting temperature, below the melting point of the metallic cathode 122. Upon reaching of the surface of a electron emitting temperature, electrons are thermionically emitted from the surface. The high voltage field between the cathode 122 and the target element 128 (shown in FIGURES 3 and 4) accelerates these electrons, thereby forcing them to strike the surface of the target element 128 and produce x-rays. In one embodiment of the invention, a Nd:YAG laser was coupled into a SiO2 optical fiber having a diameter of 400 microns. A 20 kV power supply was used, and a thermionic cathode made of tungsten was used. Only a few watts of power was needed to generate over 100 μA of electron current. Another way to increase the efficiency of the laser heated thermionic cathode, besides using laser energy to drive the thermionic cathode, is to minimize heat loss due to incident laser radiation that remains unabsorbed by the thermionic cathode. FIGURE 6 illustrates one embodiment of an electron source embodying the present invention, in which reflector elements 160 are included which reflect back to the thermionic cathode 122 incident laser radiation that was unabsorbed by the thermionic cathode 122. FIGURE 6 shows an illustrative incident ray 152 of laser radiation which is unabsorbed and scattered by the thermionic cathode 122. The scattered ray 153 of laser radiation impinges upon the inner surface of the capsule 130 enclosing the radiation generator assembly 101. By placing reflector elements 160 at predetermined locations along the inner surface of the capsule 130, incident laser radiation that remained unabsorbed by the electron emissive surface of the thermionic cathode 122 is reflected back by the reflector elements 160 to the thermionic cathode 122, so that an optical cavity is effectively created within the capsule. The coupling efficiency of the incident laser radiation to the thermionic cathode 122 is thereby significantly increased.

FIGURE 7 shows an embodiment of the present invention in which the therapeutic radiation source includes a spiral- shaped, laser-heated thermionic cathode 222. As in previous embodiments, a capsule 230 defines a substantially evacuated region extending along a beam axis 209, between the cathode 222 at the proximal end of the capsule 230 and a target element 228 at the distal end of the capsule 230, and incident laser light is transmitted through a fiber optical cable 213. The spiral-shaped thermionic cathode 222 preferably has a plurality of spaced apart turns, an interstitial spacing being defined between each successive turn. Heat loss in the cathode due to thermal conduction is minimized, due to the spiral-shaped configuration of the cathode. The thermionic cathode 222 includes a spiral-shaped conductive element having a plurality of spaced apart turns that define an interstitial spacing between adjacent turns. The conductive element may be a wire, by way of example. The conductive element may also be a photochemically machined flat spiral of cathode material. The spiral arrangement of the wire results in a reduction of conductive heat loss in the cathode.

For a disc-shaped or planar tungsten thermionic cathode, the percentage of incident radiation that is absorbed at an incident spot on the cathode is typically about 40 %. Of the 40 % absorbed, however, further losses are caused because of thermal conduction within the cathode. In the illustrated embodiment, heat loss through thermal conduction is minimized, because the cathode 222 is in the shape of a spiral coil having a plurality of spaced apart turns. Heat loss through thermal conduction is substantially reduced, as compared to heat conduction within a disk-shaped thermionic cathode, because no heat transfer occurs across the vacuum between adjacent, spaced apart turns of the conductive element forming the cathode.

As mentioned earlier, with an optically drive thermionic cathode as featured in the present invention, only a few watts of power are needed to generate over 100 μA of electron current, even with a disc-shaped, planar cathode. Using an infrared diode laser in conjunction with a spiral-shaped, half millimeter etched cathode, about 100 μA of electron current can be achieved with only 180 mW of power, thereby substantially reducing the power requirements for the apparatus.

FIGURES 8(a) and 8(b) illustrate in more detail a spiral-shaped cathode 300 constructed in accordance with the present invention. FIGURE 8(a) illustrates a planar view of the spiral- shaped cathode 300, whereas FIGURE 8(b) illustrates a side view. In a preferred embodiment, the spiral-shaped cathode 300 may be fabricated by using photoetching techniques known in the art. The spiral-shaped cathode 300 includes a conductive element 310 arranged in a spiral shape. The material forming the spiral-shaped conductive element is preferably a high melting point metal adapted to withstand high temperature uses. Suitable materials forming the cathode may include tungsten, thoriated tungsten, other tungsten alloys, tantalum, rhenium, thoriated rhenium, and molybdenum. Preferably, the spiral-shaped conductive element 310 forms a planar coil, although other forms of conductive coils may be used, such as helical coils. Spiral coils of various shapes can be used. For example, each of the plurality of spaced apart turns may have a substantially circular shape, when viewed from the longitudinal direction. Alternatively, the spiral coil may have other transverse sectional shapes, such as oval, square, or rectangular.

The spiral-shaped conductive element 310 has a plurality of spaced apart turns, which define an interstitial spacing 330 between each successive turn. The conductive element 310 may have a length of about 2 mm to about 7 mm, although other dimensions are also within the scope of this invention. The distance between adjacent turns of the conductive element 310 may be about 25 microns to about 50 microns, although other dimensions are also within the scope of this invention. Since the spiral-shaped cathode 300 is disposed within the vacuum within the capsule 230, heat transfer across the interstitial spacing 330 between adjacent turns of the conductive element 310 is essentially eliminated. In this way, heat loss in the thermionic cathode 300 that is caused by thermal conduction is substantially reduced.

In an exemplary embodiment, the spiral-shaped thermionic cathode 300 was fabricated using a conductive wire 0.002 mm thick, and 7.4 mm long. In this embodiment, the conductive wire defined two spaced-apart turns. The power loss caused by thermal conduction was only 0.126 Watts, as compared to planar, disk-shaped cathodes, in which the power loss due to thermal conduction was about 1.1 Watts. The power loss caused by thermal radiation was about 140 mW.

FIGURE 9 provides an enlarged view of a radiation generator assembly, and the distal end of the probe assembly, constructed in accordance with one embodiment of the present invention in which a weakly conductive or semiconductive coating is applied to the inner surface of the rigid capsule that encloses the electron source and the target element. As seen from FIGURE 9, an electron source 208 and a target element 228 are enclosed within an evacuated capsule 230. The interior surface of the capsule is coated with a layer 207 of weakly conductive or semiconductive material. The layer 207 or weakly conductive or semiconductive coating is adapted to provide a substantially smooth voltage gradient within the capsule.

In optically-driven, miniaturized therapeutic radiation sources, such as the devices disclosed in U.S. Patent No. 5,428,658, the interior surface of the capsule 230 is typically lined with an electrical insulator. In contrast, in the embodiment illustrated in FIGURE 9, the interior surface is coated with a layer 207 of semiconductive or weakly conductive coating. The layer 207 of weakly conductive or semiconductive coating serves to prevent localized high electric field regions or "spikes" within the accelerating region inside the capsule 230, thereby substantially reducing the chances of electrical breakdown within the vacuum inside the capsule

230. The layer 207 of weakly conductive or semiconductive coating also substantially reduces the chances of the secondary emissions from electrons striking the inner wall of the capsule 230 from causing a avalanche that eventually leads to an electrical breakdown. The coating also ensures that the electron beam is not deflected off the target by chargeup effects on the insulator wall.

The illustrated embodiment of the present invention features an optically-driven, high efficiency therapeutic radiation source that maintains a substantially uniform voltage gradient within the evacuated region between the electron source and the target. The layer 207 of weakly conductive or semiconductive coating allows a substantially smooth voltage gradient to be maintained between a predetermined maximum value of the accelerating voltage and ground potential. The layer 207 of weakly conductive or semiconductive coating also serves to screen the thermionic cathode's "triple junction point," i.e. the junction between the cathode, housing wall, and vacuum, from high electric fields, thus preventing electron field emission and subsequent high voltage breakdown.

FIGURE 10 (a) illustrates an enlarged view of the interior of the substantially rigid capsule 230 having a layer 207 of weakly conductive or semiconductive coating. In particular, FIGURE 10 (a) illustrates the electromagnetic field lines 330 within the interior of the capsule 230, showing a substantially smooth voltage gradient across the evacuated region within the capsule.

As seen from FIGURE 10 (a), the capsule 230 has an interior surface 310 that defines a hollow, evacuated region 312. As mentioned earlier, a high voltage power supply provides an accelerating voltage for accelerating the electrons emitted from the electron source toward the target element. Because it is safer to maintain the target element at ground potential, the thermionic cathode may be negatively biased, so that the target element is maintained at a net positive voltage with respect to the thermionic cathode. The accelerating voltage has a predetermined maximum value, typically about 90 keV.

The hollow interior surface 310 is coated with a layer 207 of weakly conductive or semiconductive coating. This is because the weakly conductive or semiconductive coating allows for voltage gradient control within the evacuated region. In other words, the weakly conductive or semiconductive coating 207 allows a substantially smooth voltage gradient to be maintained between the predetermined maximum value of the accelerating voltage, and the ground potential, as illustrated by the voltage gradient field lines 330 in FIGURE 10(a). The coating 207 may be made of weakly conductive or semiconductive materials, including but not limited to chromium sesquioxide, vanadium pentoxide, or ion implanted metals such as platinum.

The high-resistance coating 207, which is weakly conductive or semiconductive, and which is applied to the interior surface 310 of the capsule 230 in the optically-driven, miniaturized therapeutic radiation source of the present invention, improves the ability of the therapeutic source 200 to withstand high acceleration voltages, without breakdown. It also greatly enhances the x-ray output and stability, as the efficiency in propagating the electrons to the target is greatly increased. This is in contrast to prior art devices, such as disclosed in U.S. Patent No. 5,428,658, and in contrast to the miniaturized, optically-driven therapeutic radiation sources that do not include such a coating. These devices contained a film of insulating material on the inner surface of the evacuated capsule, rather than a high resistance, semiconductive (or weakly conductive) coating, so that no control was possible over the voltage gradient within the evacuated region.

As mentioned earlier, the reliability of the vacuum within the capsule 230 is limited by the operational risk of an unpredictable "sparking" or "arcing" between the electrodes, when the insulating capability of the evacuated region 312 is suddenly lost, and electrical breakdown is said to have occurred. Because of such practical limitation of the insulating capability of the vacuum, localized high voltage gradient regions or "spikes" may be caused. Such spikes may occur in the accelerating region within the evacuated capsule 230, as the emitted electrons are accelerated toward the target. In order to avoid such spikes in the electric field within the evacuated region 312, it is preferable that the inner surface 310 of the capsule 230 be lined with a weakly conductive or semiconductive material that can directly control the electric field. The present invention provides such as low dielectric constant material in the form of the resistive layer 207 of weakly conductive or semiconductive coating. There is a wide range of physical phenomena, such as electrode heating and thermal diffusion processes, and electron emissions, which contribute to the performance of the vacuum gap and its ability to withstand electrical breakdown. The breakdown voltage depends on a number of parameters, including but not limited to electrode material and geometry, surface preparation and geometry, vacuum quality, and vacuum gap spacing. In particular, the breakdown voltage is typically a function of the dielectric constant of the material forming the capsule 230. It is preferable that the material forming the capsule 230 have a high dielectric strength, in order to withstand a large electrical field without breakdown. Preferably, the dielectric strength of the capsule material, which may be a ceramic by way of example, is at least 100 kV / mm. The ceramic material that forms the capsule 230 may include, but is not limited to, glass, boron nitride, sapphire, fused silica, and diamond.

The layer 207 of weakly conductive or semiconductive coating also serves to reduce secondary emissions from the interior surface 310. Secondary emissions of electrons hitting the walls of the capsule enclosing the accelerating region may cause an avalanche, eventually leading to a breakdown. Such an avalanche is likely to happen when the inner surface 310 of the capsule 230 is lined with an insulator material, for example a material having a secondary emission coefficient > 3. The semiconductive coating 207, which in one exemplary embodiment may have a secondary emission coefficient < 1 , serves to prevent such an avalanche. Also, charge-up, and subsequent electric field intensification and breakdown, are eliminated by draining the charge through the weakly conductive or semiconductive layer. FIGURE 10 (b) illustrates the thermionic cathode's triple junction point 350. The weaker the electric field at the cathode, the more imperfections or irregularities can be tolerated on the surface of the thermionic cathode, without risking flashover. The triple junction point 350 may be screened from the high electrical field between the target element and the cathode 222 by the weakly conductive or semiconductive coating, thereby substantially reducing the chances of electrical flashover.

Referring back to FIGURE 10(a), the electric field within the evacuated accelerating region is controlled by a resistive, weakly conductive or semiconductive coating 207 on the inner surface of the capsule enclosing the x-ray generator assembly. The weakly conductive or semiconductive coating 207 generates a controlled voltage gradient across the vacuum within the capsule 230. Also, the weakly conductive or semiconductive coating 207 is adapted to reduce the strength of the electric field near the triple junction point of the thermionic cathode, thereby reducing the chances of electrical flashover. Finally, the weakly conductive or semiconductive coating 207 prevents the amplification of the secondary emissions of electrons that hit the inner wall of the capsule 230 by absorbing the emitted electrons, thereby preventing an avalanche of emissions that could lead to electrical breakdown. For these reasons, the chances for electrical flashover or electric breakdown within the evacuated capsule 230 is substantially reduced, thereby providing a significant advantage over optically driven therapeutic radiation sources in which the inner surface of the capsule 230 is lined with an insulator material. Also, the weakly conductive or semiconductive coating increases the x-ray production efficiency and stability by ensuring that the electrons emitted from the cathode proceed directly to the target, thereby maximizing the probability that they will hit the x-ray target, and maximizing the electrons' kinetic energy as they impact the target.

In sum, the present invention significantly reduces the power requirements for such miniaturized therapeutic radiation sources, by heating a thermionic cathode with laser energy, instead of resistively heating the thermionic cathode with a current. The present invention also features the use of a spiral-shaped thermionic cathode, which is configured so as to minimize energy lost from the incident laser radiation due to thermal conduction within the thermionic cathode. In this way, the power requirements for generating therapeutic radiation in such miniaturized radiation sources are further reduced. Finally, the present invention also features the use of a weakly conductive or semiconductive coating, applied on the inner surface of the evacuated capsule. In this way, a substantially uniform voltage gradient is established in the region between the electron source and the target, so that high electric field regions or spikes are avoided within the evacuated capsule, the chances of electrical breakdown are reduced, and the electrons propagate directly to the target. The weakly conductive or semiconductive coating also reduces the chances for secondary emissions of electrons that strike the walls of the capsule to cause an avalanche that eventually causes an electrical breakdown. Field emitted currents are swept away, preventing charge-up and breakdown.

While the invention has been particularly shown and described with reference to specific preferred embodiments, it should be understood by those skilled in the art that various changes in form and detail may be made therein without departing from the spirit and scope of the invention as defined by the appended claims.

Claims

1. A therapeutic radiation source, comprising:

A. a probe assembly including an optical delivery structure having a proximal end and a distal end, said optical delivery structure being adapted for transmitting optical radiation incident on said proximal end to said distal end;

B. an optical source, including means for generating a beam of optical radiation directed to said proximal end of said optical delivery structure;

C. a radiation generator assembly coupled to said probe assembly, including: a. an electron source, responsive to optical radiation transmitted to said distal end of said optical delivery structure, for emitting electrons, the electron source including a thermionic cathode having an electron emissive surface; and b. a target element including at least one radiation emissive material adapted to emit therapeutic radiation in response to incident accelerated electrons from said electron source; and D. means for providing an accelerating voltage between said electron source and said target element so as to establish an accelerating electric field which acts to accelerate electrons emitted from said electron source toward said target element; wherein said optical delivery structure is adapted for directing a beam of optical radiation transmitted therethrough to impinge upon a surface of said thermionic cathode, and wherein said beam of transmitted optical radiation has a power level sufficient to heat at least a portion of said surface to an electron emitting temperature so as to cause thermionic emission of electrons from said surface.

2. A therapeutic radiation source according to claim 1, wherein said optical source is a laser, and wherein said beam of optical radiation is substantially monochromatic and coherent.

3. A therapeutic radiation source according to claim 1, wherein said electron emissive surface of said thermionic cathode is formed of a metallic material.

4. A therapeutic radiation source according to claim 4, wherein said metallic material includes tungsten, thoriated tungsten, tungsten alloys, thoriated rhenium, and tantalum.

5. A therapeutic radiation source according to claim 1, wherein the emitted electrons form an electron beam along a beam path, and wherein said target assembly is positioned in said beam path.

6. A therapeutic radiation source according to claim 1, wherein said electron beam is characterized by a current in the approximate range of about 1 nA to about 1mA.

7. A therapeutic radiation source according to claim 1, wherein said electrons incident on said target element from said electron emissive surface are accelerated by said accelerating electric field to energies in the approximate range of 10 keV to 90 keV.

8. A therapeutic radiation source according to claim 1, wherein the means for providing an accelerating voltage includes a power supply, having a first terminal and a second terminal, and a drive means for establishing an output voltage between said first terminal and said second terminal, said power supply being electrically coupled to said radiation generator assembly by way of said first terminal and said second terminal.

9. A therapeutic radiation source according to claim 8, wherein said first terminal of said power supply is electrically coupled to said electron emissive surface of said thermionic cathode and said second terminal of said power supply is electrically coupled to said target element, thereby establishing an electric field which acts to accelerate electrons emitted from said electron emissive surface of said thermionic cathode toward said target element.

10. A therapeutic radiation source according to claim 8, wherein said second terminal is at ground potential.

11. A therapeutic radiation source according to claim 8, wherein said power supply further includes selectively operable control means for selectively controlling the amplitude of said output voltage.

12. A therapeutic radiation source according to claim 6, further including selectively operable control means for selectively controlling the amplitude of said beam current.

13. A therapeutic radiation source according to claim 1, wherein said thermionic cathode includes a metallic base coated with an oxide.

14. A therapeutic radiation source according to claim 13, wherein said oxide includes barium oxide, strontium oxide, and calcium oxide, and said metallic base includes nickel.

15. A therapeutic radiation source according to claim 1, wherein said optical delivery structure includes a fiber optic cable.

16. A therapeutic radiation source according to claim 15, wherein said probe assembly includes a flexible, electrically conductive catheter enclosing said fiber optic cable.

17. A therapeutic radiation source according to claim 16, wherein the means for establishing an accelerating voltage includes a power supply, said power supply having a first terminal and a second terminal, said power supply being electrically coupled to said radiation generator assembly by way of said first terminal and said second terminal.

18. A therapeutic radiation source according to claim 16, wherein said electrically conductive catheter is adapted for coupling said second terminal of said power supply to said radiation generator assembly.

19. A therapeutic radiation source according to claim 17, wherein said fiber optic cable includes an electrically conductive outer surface, said electrically conductive outer surface being adapted for electrically coupling said first terminal of said power supply to said thermionic cathode.

20. A therapeutic radiation source according to claim 1 , further including a substantially rigid capsule, wherein said electron source and said target element are disposed within said capsule, and further wherein said capsule defines a substantially evacuated interior region extending along a beam axis between said thermionic cathode at a proximal end of said capsule and said target element at a distal end of said capsule.

21. A therapeutic radiation source according to claim 1 wherein said therapeutic radiation includes x-rays.

22. A source of therapeutic radiation, comprising:

A. a radiation generator assembly, including: a. an electron source for emitting electrons to generate an electron beam along a beam path, wherein said electron source includes a thermionic cathode having an electron emissive surface; and b. a target element positioned in said beam path, said target element being spaced apart from and opposite said electron emissive surface, said target element comprising at least one radiation emissive element adapted to emit therapeutic radiation in response to incident accelerated electrons from said electron beam; and c. a substantially rigid capsule, wherein said electron source and said target element are disposed within said capsule, and further wherein said capsule defines a substantially evacuated interior region extending along a beam axis between said thermionic cathode at a proximal end of said housing and a radiation transmissive window at a distal end of said housing; B. a source of laser radiation; C. a probe assembly coupled to said radiation generator assembly and including an optical delivery structure, said optical delivery structure having a proximal end and a distal end; wherein said optical delivery structure is adapted for transmitting laser radiation, generated by said source and incident on said proximal end, to said distal end, and for directing a beam of said transmitted laser radiation to impinge upon said electron emissive surface of said thermionic cathode, and wherein said beam of laser radiation has a power level sufficient to heat at least a portion of said surface to an electron emitting temperature so as to cause thermionic emission of electrons from said surface; D. one or more reflector elements disposed at predetermined locations along an inner surface of said capsule, said one or more reflector elements being operative to reflect incident laser radiation unabsorbed by said thermionic cathode back to said thermionic cathode.

23. A therapeutic radiation source according to claim 1, wherein power required to heat said electron emissive surface of said cathode so as to generate an electron beam forming a current of about 100 micro amps is between about 0.1 Watts to about 3.0 Watts.

24. A therapeutic radiation source according to claim 1, wherein said target element is spaced apart and opposite said electron emissive surface of said thermionic cathode.

25. A therapeutic radiation source according to claim 19, further comprising a layer of dielectric material between said electrically conductive outer surface of said fiber optical cable and an inner surface of said flexible catheter.

26. A therapeutic radiation source, comprising:

A. a radiation generator assembly, comprising: a. an electron source for emitting electrons to generate an electron beam along a beam path, said electron source including a thermionic cathode having an electron emissive surface, and b. a target positioned in said beam path, said target including means for emitting therapeutic radiation in response to incident accelerated electrons from said electron beam; wherein said thermionic cathode comprises a spiral-shaped conductive element;

B. a source of optical radiation; and C. optical delivery structure having an originating end and a terminating end and adapted for transmitting to said terminating end optical radiation generated by said source and incident on said originating end; and wherein said optical delivery structure are adapted for directing a beam of said transmitted optical radiation upon a surface of said thermionic cathode; and wherein said beam of optical radiation has a power level sufficient to heat at least a portion of said surface to an electron emitting temperature so as to cause thermionic emission of electrons from said surface.

27. A therapeutic radiation source according to claim 26, wherein said thermionic cathode is disposed at said input end of said housing.

28. A therapeutic radiation source according to claim 26, further comprising a radiation transmissive window at an output end of said housing, wherein therapeutic radiation emitted from said target is directed through said radiation transmissive window.

29. A therapeutic radiation source according to claim 26, wherein said spiral-shaped conductive element defines a plurality of spaced apart turns.

30. A therapeutic radiation source according to claim 29, wherein said conductive element defines an interstitial space between each successive turn.

31. A therapeutic radiation source according to claim 29, wherein said spiral-shaped conductive element forms a planar coil.

5 32. A therapeutic radiation source according to claim 29, wherein said spiral-shaped conductive element forms a helical coil.

33. A therapeutic radiation source according to claim 29, wherein the distance between adjacent turns of said conductive coil is from about 25 microns to about 50 microns.

10

34. A therapeutic radiation source according to claim 29, wherein each of said plurality of spaced apart turns has a transverse sectional shape that is substantially circular.

35. A therapeutic radiation source according to claim 26, wherein said optical delivery 5 structure comprises a fiber optical cable.

36. A therapeutic radiation source according to claim 26, wherein said fiber optical cable has a diameter between about 100 microns to about 200 microns.

0 37. A therapeutic radiation source according to claim 29, wherein said spiral-shaped conductive coil has a length between about 2 mm to about 7 mm.

38. A therapeutic radiation source according to claim 26, wherein the power required for heating said electron emissive surface of said cathode so as to generate an electron beam forming 5 a current of about 2 micro amps is between about 0.1 Watt to about 1.0 Watt.

39. A therapeutic radiation source according to claim 26, wherein said optical source is a laser, and wherein said beam of optical radiation is substantially monochromatic and coherent.

0 40. A therapeutic radiation source according to claim 26, wherein said therapeutic radiation comprises x-rays.

? 41. A therapeutic radiation source according to claim 26, wherein power loss caused by thermal conduction is less than 0.2 Watts.

42. A therapeutic radiation source, comprising:

A. a radiation generator assembly, comprising: a. an electron source for emitting electrons to generate an electron beam along a beam path, said electron source including a thermionic cathode having an electron emissive surface, and b. a target positioned in said beam path, said target including means for emitting therapeutic radiation in response to incident accelerated electrons from said electron beam; and c. a substantially rigid housing enclosing said thermionic cathode and said target, wherein said housing defines a substantially evacuated interior region extending along said beam path between an input end and an output end of said housing.

B. a source of optical radiation; and C. optical delivery structure having an originating end and a terminating end and adapted for transmitting to said terminating end optical radiation generated by said source and incident on said originating end, said optical delivery structure being adapted for directing a beam of said transmitted optical radiation upon a surface of said thermionic cathode, wherein said beam of optical radiation has a power level sufficient to heat at least a portion of said surface to an electron emitting temperature so as to cause thermionic emission of electrons from said surface; and wherein said thermionic cathode comprises a spiral-shaped conductive element having a plurality of spaced apart turns.

43. A therapeutic radiation source according to claim 42, wherein heat transfer across the spacing between each adjacent turn of said conductive element is essentially eliminated, thereby substantially reducing in said thermionic cathode heat loss caused by thermal conduction.

44. A therapeutic radiation source according to claim 26, further including means for establishing an accelerating electric field which acts to accelerate electrons emitted from said electron source toward said target.

45. A therapeutic radiation source, comprising:

A. an optical source;

B. a probe assembly, including an optical delivery structure having a proximal end and a distal end, said optical delivery structure being adapted for transmitting optical radiation incident on said proximal end to said distal end, said optical delivery structure being adapted for directing a beam of optical radiation transmitted therethrough to impinge upon a surface of said thermionic cathode;

C. a radiation generator assembly coupled to said probe assembly, including: a. an electron source, responsive to optical radiation transmitted to said distal end of said optical delivery structure, for generating an electron beam along a beam path, said electron source including a thermionic cathode having an electron emissive surface; and b. a target element positioned in said beam path, said target element including at least one x-ray emissive material adapted to emit x-rays in response to incident accelerated electrons from said electron source; and c. a substantially rigid capsule enclosing said electron source and said target element, wherein said capsule defines a substantially evacuated interior region extending along a beam axis between said thermionic cathode at a proximal end of said capsule and said target element at a distal end of said capsule; D. means for providing an accelerating voltage between said electron source and said target element so as to establish an accelerating electric field which acts to accelerate electrons emitted from said electron source toward said target element, said accelerating voltage at said source having a preselected maximum value relative to a reference potential at said target element; wherein said beam of transmitted optical radiation has a power level sufficient to heat at least a portion of said surface to an electron emitting temperature so as to cause thermionic emission of electrons from said surface; and wherein the inner surface of said capsule is coated with a semiconductive coating to provide a substantially smooth voltage gradient within said capsule between said preselected maximum value and the reference potential.

46. A therapeutic radiation source according to claim 45, wherein said semiconductive coating comprises a weakly conductive coating.

47. A therapeutic radiation source according to claim 45, wherein said semiconductive coating is formed of an ion-implanted metal.

48. A therapeutic radiation source according to claim 45, whereby a surface resistivity of said inner surface of said capsule establishes a relatively low power dissipation.

49. A therapeutic radiation source according to claim 45, wherein the dielectric constant of said inner surface of said capsule is sufficiently small so as to prevent spikes in said accelerating electric field.

50. A therapeutic radiation source according to claim 45, wherein a secondary electron emission coefficient of said inner surface of said capsule is less than 1.

51. A therapeutic radiation source according to claim 45, wherein said semiconductive coating applied to said inner surface of said capsule is adapted to screen from electrical flashover a triple junction point of said thermionic cathode.

52. A therapeutic radiation source according to claim 45, wherein said optical source is a laser, and wherein said beam of optical radiation is substantially monochromatic and coherent.

53. A therapeutic radiation source according to claim 45, wherein said optical delivery structure includes a fiber optic cable.

54. A therapeutic radiation source according to claim 53, wherein said probe assembly includes a flexible, electrically conductive catheter enclosing said fiber optic cable.

55. A therapeutic radiation source according to claim 45, wherein said semiconductive coating is adapted to reduce secondary emissions of electrons striking said inner surface of said capsule.

56. A therapeutic radiation source, comprising:

A. an optical source;

C. a radiation generator assembly coupled to said probe assembly, including: a. an electron source, responsive to optical radiation transmitted to said distal end of said optical delivery structure, for generating an electron beam along a beam path, said electron source including a thermionic cathode having an electron emissive surface; and b. a target element positioned in said beam path, said target element including at least one x-ray emissive material adapted to emit x-rays in response to incident accelerated electrons from said electron source; and c. a substantially rigid capsule enclosing said electron source and said target element, wherein said capsule defines a substantially evacuated interior region extending along a beam axis between said thermionic cathode at a proximal end of said capsule and said target element at a distal end of said capsule;

D. means for providing an accelerating voltage between said electron source and said target element so as to establish an accelerating electric field which acts to accelerate electrons emitted from said electron source toward said target element; wherein said beam of transmitted optical radiation has a power level sufficient to heat at least a portion of said surface to an electron emitting temperature so as to cause thermionic emission of electrons from said surface; and wherein the inner surface of said capsule is coated with a semiconductive coating to provide a substantially smooth voltage gradient within said capsule between said electron source and said target element.

57. A therapeutic radiation source according to claim 56, wherein said semiconductive coating is adapted to reduce secondary emissions of electrons striking said inner surface of said capsule.

58. A therapeutic radiation source according to claim 45, wherein said semiconductive coating is formed of a material selected from the group consisting of chromium sesquioxide, vanadium pentoxide, and platinum.