HK1076410A

HK1076410A - Optically driven therapeutic radiation source

Info

Publication number: HK1076410A
Application number: HK05107872.3A
Authority: HK
Inventors: 马克．丁斯莫尔
Original assignee: Photo-Electron Corporation
Priority date: 2001-06-19
Filing date: 2002-06-19
Publication date: 2006-01-20

Description

Optically driven therapeutic radiation source

Technical Field

The present invention relates to therapeutic radiation sources, and more particularly to a radiation source using an optically driven thermionic cathode, which has reduced power, improved efficiency, and reduced size.

Background

In the medical field, therapeutic radiation, such as x-ray radiation and gamma-ray radiation, is used to diagnose, treat and administer relief therapy to a patient. Conventional medical radiation sources for such treatments include large, fixed-position machines, and small, portable, production radiation detectors. Prior art treatment systems utilize computers to generate complex treatment protocols.

Conventional radiation systems for medical treatment utilize a high-energy remote radiation source and direct a beam of radiation to a target area, such as a tumor, within the body of a patient. This type of treatment is called teletherapy because the radiation source is located at a predetermined distance from the target. This treatment has the drawback that the tissue arranged between the radiation source and the target is exposed to the radiation. Teletherapy radiation sources, which radiate from a source outside the target area to the target area within the patient, often cause significant damage not only to the target area or tissue, but also to all surrounding tissue between the entry site, the target area and the exit site. Brachytherapy, on the other hand, is a form of treatment in which the radiation source is located close to, or in some cases within, the area being treated. The term brachytherapy, derived from the ancient greek near ("brachy") vocabulary, offers significant advantages over teletherapy because radiation is applied primarily to only a predetermined tissue volume, without significantly affecting tissue adjacent to and outside the treated volume. The term brachytherapy is commonly used to describe the use of radioactive "seed-shaped tubules", i.e. encapsulated radioisotopes, which are placed directly in or close to the target area to be treated. However, handling and disposal of such radioisotopes can pose significant hazards to both the operator and the environment.

The term "x-ray brachytherapy" is defined as for x-ray radiation therapy, in which an x-ray source is located close to or within the region being treated. An x-ray brachytherapy system employing a small low power radiation source that is insertable into and generating radiation from a patient is disclosed in U.S. patent 5153900 to Nomikos et al, U.S. patent 5369679 to Sliski et al, and U.S. patent 5422926 to Smith et al, all of which are owned by the assignee of the present applicant and are hereby incorporated by reference. The x-ray brachytherapy system disclosed in the above-mentioned patent includes a miniaturized insertable probe capable of generating x-ray radiation local to the target tissue, so that the radiation does not need to pass through the patient's skin, bone or other tissue before reaching the target tissue. The insertable probe emits low power x-rays from a nominal "point" source located within or near the desired region to be affected. Thus, in x-ray brachytherapy, x-rays are applied to treat a predetermined tissue volume, which does not significantly affect the tissue adjacent to the treatment volume. Moreover, the x-rays of the predetermined dose geometry may be generated around the predetermined location while avoiding the use and manipulation of radioisotopes. Furthermore, x-ray brachytherapy allows the operator to always control the dose of emitted x-ray radiation.

x-ray brachytherapy typically involves positioning an insertable probe within or near a tumor, or within a site of tumor or partial tumor resection, to treat tissue near the site by local elevation of radiation. An x-ray detector of the type generally disclosed in U.S. patent 5153900 includes a housing, and a hollow tubular detector or catheter extending from the housing along an axis and having an x-ray emitting target at a distal end. The detector may encapsulate an electron source, such as a thermionic cathode. In another form of x-ray brachytherapy apparatus, as disclosed in U.S. patent 5428658, the x-ray detector may include a flexible detector, such as a flexible fiber optic cable encased in a metal sleeve. The x-ray detector can also include a substantially rigid capsule coupled to the distal end of the flexible detector. A capsule encloses the electron source and the x-ray emitting target element. The electron source may be a photocathode. In a photocathode configuration, a photoemissive substance is irradiated by an LED or laser source, resulting in the generation of free electrons. Typically, a flexible fiber optic cable transmits light from a laser source or LED to a photocathode.

In the devices disclosed in us patent 5133900 and 5428658, an accelerating electric field can be formed between the electron source and the target element. The generated electric field is used to accelerate electrons emitted by the electron source towards the target element. The target element emits radiation in response to incident electrons from the electron source.

In one form of conventional thermionic cathode, the filament is resistively heated with an electric current. The cathode is then heated to generate electrons by thermionic emission. In one form of conventional x-ray machine employing such resistively heated thermionic cathodes, the cathode assembly may consist of a thoriated tungsten coil approximately 2mm in diameter and 1 to 2 cm in length. Thoriated tungsten coils thermionically emit electrons when resistively heated with a current of 4A or higher. In one configuration, the coil is surrounded by a metallic focusing cup that focuses the electron beam to a small spot on an opposing anode that also serves as a target. The electron beam is focused to a spot diameter on the anode, typically in the range of from about 0.3 to.2.5 mm. In many applications, most of the energy from the electron beam is converted to heat at the anode. To accommodate such heating, high power medical x-ray sources often utilize liquid cooling and a rapidly rotating anode. Thus, a target region of increased efficiency is formed, allowing a small focal spot while minimizing the efficiency of local heating.

To achieve good heat conduction and efficient heat dissipation, the anode is typically made of copper. In addition, the anode region on which the electron beam is incident must be made of a material with a high atomic number in order for the x-rays to be efficiently generated. To meet the requirements of thermal conductivity, efficient heat dissipation, and efficient x-ray generation, tungsten alloys are typically embedded in copper.

It is desirable that the electron source is heated as efficiently as possible, i.e. the thermionic cathode reaches as high a temperature as possible with as low a power as possible. In conventional x-ray tubes, for example, thermal evaporation of the helical cathode filament of the tube is often the cause of tube damage. Moreover, the anode heated to a high temperature causes a decrease in radiation output. During longer exposures of the x-ray source, for example, from about 1 second to about 3 seconds of exposure, the anode temperature may rise sufficiently to cause it to glow brightly, with localized surface melting and localized corrosion, thereby reducing radiation output.

While photocathodes avoid these problems, one disadvantage of using photocathodes is that it is difficult to manufacture photocathodes. The photocathode must have sufficient quantum efficiency, which relates to the number of electrons generated per incident photon. The degree of efficiency must be balanced with the actual incident light intensity. For practical substances, reasonable quantum efficiencyGreater than 10^-3The fabrication of the photocathode should be performed in vacuum. In one form, vacuum fabrication is performed using fiber optic cables located in a bell jar, as disclosed in U.S. patent 5428658, which is owned by the assignee of the present applicant and incorporated herein by reference. For example, Ag-O-Cs photoactive surfaces can be fabricated in a conventional manner. As a result, the fiber optic cable is inserted into the tubular probe without exposure to air, and the end of the fiber optic cable is vacuum sealed to the probe.

It is an object of the present invention to provide an increased efficiency, miniaturized radiation source with significantly reduced power requirements. It is another object of the invention to provide a compact radiation source in which the electron source can generate electrons with minimal heat loss and without the need for a vacuum-fabricated photocathode. It is a further object of the invention to provide a miniaturized radiation source in which laser energy is used to heat a thermionic cathode instead of heating the thermionic cathode by conventional resistive heating. In this way, a sufficient number of electrons can be generated to create the electron flow required to generate therapeutic radiation at the target, while significantly reducing the power requirements necessary for the radiation source.

In order to reduce the power required by the above-described laser-heated therapeutic radiation sources, the thermal losses of the thermionic cathode must be minimized. Heat loss in laser heated thermionic cathodes includes 1) heat loss by thermal conduction; 2) heat loss due to the portion of incident laser radiation that remains unabsorbed; and 3) heat loss by thermal radiation.

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

In the devices disclosed in us patents 5133900 and 5428658, the electron source and the target element are enclosed within a substantially rigid membrane cartridge. The electron source generates an electron beam along a beam path, and the target element is positioned within the beam path. An accelerating electric field may be formed within the capsule. The accelerating electric field is used to accelerate the electrons emitted by 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 vacuum is widely used to insulate high voltages in devices such as the x-ray detectors 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 an electrical breakdown is said to occur. Moreover, efficient generation of x-rays requires an electron path directly from the cathode to the target. If electrons are deflected to the wall by the dielectric charging effect, the efficiency of x-ray generation is reduced and the stability of the x-ray output is compromised.

It is therefore important to create a substantially uniform voltage gradient in the region between the electron source and the target to avoid such electrical breakdown and to maximize and stabilize the x-ray output. It is therefore another object of the present invention to provide a highly efficient miniaturized therapeutic radiation source that forms a substantially uniform voltage gradient within the vacuum region between the electron source and the target.

Disclosure of Invention

The present invention relates to a compact therapeutic radiation source having a low power, electron beam activated radiation source. In particular, the apparatus of the invention comprises a thermionic cathode heated by a source of optical radiation, preferably a laser.

A therapeutic radiation source according to the present invention includes a radiation generator assembly, a source of optical radiation, and a detector assembly. The optical radiation source is preferably a laser which produces a substantially monochromatic coherent radiation beam. The radiation generator assembly includes an electron source for emitting electrons to generate electrons along a beam path having a nominally straight reference axis, and a target element positioned in the beam path. The electron source is preferably a light-driven thermionic cathode having an electron-emitting surface.

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

The target element includes means for emitting therapeutic radiation in response to accelerated electrons incident from the electron beam. In a preferred embodiment, the target element is spaced apart from and opposite the electron emission surface of the thermionic cathode. The target element comprises at least one radiation emitting element adapted to emit therapeutic radiation in response to accelerated electrons incident from said thermionic cathode. The therapeutic radiation source further comprises means for providing an accelerating voltage so as to generate an accelerating electric field for accelerating the accelerated electrons emitted by said electron source towards the target element.

In one embodiment, the radiation generator assembly further comprises a substantially rigid capsule enclosing the electron source and the target element. Preferably, the electron source is located at its proximal end and the target element is located at its distal end. The capsule defines a substantially evacuated interior region extending along a nominally straight beam axis between a 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 provided at the distal end of the capsule. The total resistance of the inner surface of the capsule is preferably high enough to limit the dissipated power to less than 10% of the 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 a preselected maximum and ground potential. The weakly conductive or semiconductive coating applied to the inner surface of the capsule is also adapted to reduce secondary emission of electrons striking the coated inner surface of the capsule. The weakly conductive or semiconductive coating is also adapted to reduce the electric field near the triple junction, thus reducing the potential for electrical flashover at the triple junction of the thermionic cathode. Sufficient current is carried in the coating to prevent charge build-up by field emission and subsequent avalanche and breakdown.

The probe assembly includes an optical transmission structure, preferably a fiber optic cable, having a proximal end and a distal end. The distal end of the fiber optic cable is connected to the radiation generator assembly. The fiber optic cable transmits optical radiation generated by the source and incident at the proximal end to the distal end. The fiber optic cable directs a beam of transmitted optical radiation to impinge upon the surface of the thermionic cathode, wherein the beam of optical radiation has a power level sufficient to heat at least a portion of the surface of the thermionic cathode to an electron emission temperature, thereby causing thermionic emission of electrons from the surface. In one embodiment, the probe assembly includes a flexible metal sleeve enclosing a fiber optic cable.

In one embodiment, the means for providing an acceleration voltage is a power supply having a first terminal and a second terminal, and has driving means for generating an output voltage between the first terminal and the second terminal. In one form, the power supply may be electrically connected to the target element via the first and second terminals. A first terminal of the power supply may be electrically connected to the electron emission surface of the thermionic cathode and a second terminal electrically connected to the target element to generate an electric field that accelerates electrons emitted from the thermionic cathode toward the target element.

In one embodiment of the invention, the apparatus of the invention comprises one or more reflector elements arranged at predetermined positions along the inner surface of the housing. The reflector element is operable to reflect incident laser radiation that is not absorbed by the thermionic cathode back to the thermionic cathode, thereby increasing the efficiency of the therapeutic radiation source.

Drawings

Figure 1(a) is a schematic 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 with a resistively heated thermionic cathode.

Figure 2 is a schematic block diagram of an overview of a therapeutic radiation source constructed in accordance with the present invention having a laser heated thermionic cathode.

Figure 3 is a schematic view of one embodiment of a therapeutic radiation source constructed in accordance with the present invention showing a laser source, a detector assembly, and a radiation generator assembly embodying the present invention.

FIG. 4 is an enlarged schematic view of one embodiment of a detector assembly and radiation generator assembly constructed in accordance with the invention.

Figure 5 is an enlarged view of one end of a radiation generator assembly according to the present invention illustrating an electron source with a laser heated thermionic cathode.

Figure 6 is an enlarged view of one embodiment of an electron source embodying the invention and illustrating reflector elements which reflect laser radiation not absorbed by the thermionic cathode back to the laser heated thermionic cathode.

Figure 7 shows an embodiment of the invention in which the therapeutic radiation source comprises a spiral laser heated thermionic cathode.

Figure 8(a) shows a plan view of a spiral thermionic cathode constructed in accordance with the present invention.

Figure 8(b) shows a side view of a spiral thermionic cathode constructed in accordance with the present invention.

FIG. 9 provides an enlarged view of the distal end of a radiation generator assembly and probe assembly constructed according to one embodiment of the present invention, wherein a weakly conductive or semiconductive coating is applied to the inner surface of a rigid capsule enclosing the electron source and target element.

Fig. 10(a) is an enlarged view of field lines of a voltage gradient in a vacuum cassette.

Fig. 10(b) shows a triple junction of a thermionic cathode used in the present invention.

Detailed Description

The present invention relates to a small low power therapeutic radiation source for use in the diagnosis, treatment and administration of comfort therapy to a patient. In the present invention, a laser is used to heat the thermionic cathode to an electron emission temperature. The power requirements of the therapeutic radiation source are significantly reduced compared to systems provided with resistively heated thermionic cathodes. The therapeutic radiation produced by the apparatus of the present invention may include, but is not limited to, x-rays. In medical applications, the device may be fully or partially implanted or surface mounted on a desired region of the host to irradiate a preselected area with therapeutic radiation. The device of the present invention can be operated at lower voltages, for example, voltages in the range of about 10KeV to 90KeV, and electron currents, for example, in the range of from about 1nA to about 1 μ A.

Figure 1(a) shows a therapeutic radiation source 10 that generates and transmits therapeutic radiation in the form of x-rays. The compact, low power x-ray source 10 shown in fig. 1(a) is a prior art x-ray brachytherapy system having a thermionic cathode heated using conventional resistive heating. One suitable system is described in detail, for example, in the above-mentioned us patent 5153900 entitled "small low power x-ray source". The system includes a housing 12, and an elongated cylindrical probe 14 extending from the housing 12 along a reference axis 16 and having a target assembly 26 at a distal end. The probe 14 may be flexible or rigid and integral with the housing 12. The housing 12 encloses a high voltage power supply 12A. The detector 14 is a hollow tube and encloses an electron source 20. The electron source 20 includes a thermionic cathode 22, and the thermionic cathode 22 may be driven by a floating low voltage power supply. In one embodiment, the electron source 20 may also include a ring-shaped focusing electrode 23, in which case the thermionic cathode 22 is positioned proximate to the ring-shaped focusing electrode 23, the ring-shaped focusing electrode 23 being typically at the same potential as the cathode. The detector 14 extends along the same axis as the cathode 22 and the focusing electrode 23.

The probe 14 may Be a hollow vacuum cylinder made of a beryllium (Be) material cap and a molybdenum-rhenium (Mo-Re), molybdenum (Mo) or mu metal body. The length of the probe 14 may be determined according to the body region to be treated. For example, the cartridge may be 15 cm long with an inner diameter of 4mm and an outer diameter of 5 mm. Different geometries of the probe 14 may be used for different body regions. The body of the probe 14 may be made of a magnetically shielded material such as a mu-metal. Alternatively, the probe 14 may be made of a non-magnetic metal, preferably having a high value of Young's modulus and elastic limit. Examples of such materials include molybdenum, rhenium, or alloys of these materials. The inner or outer surface of the probe may be coated with a high permeability magnetic alloy, such as permalloy (approximately 80% nickel and 20% iron), to provide magnetic shielding. Alternatively, a thin cylinder of mu metal may be mounted on the probe 14 or substituted therefor. The prior art x-ray device 10 can be used in environments where there are dc and ac magnetic fields generated by electrical energy, the earth's field, or other magnetized bodies that nominally enable the deflection of an electron beam from the detector axis.

Fig. 1(b) is a schematic view of the prior art x-ray source apparatus 10 shown in fig. 1 (a). In this schematic view, the housing 12 is shown as being separable into a first portion 12' and a second portion 12 ". Enclosed within the first housing portion 12' is a rechargeable battery 12B; a charging network 12D for batteries, the charging network 12D being adapted for use with an external charger 50; and a telemetry network 12E, the telemetry network 12E adapted to function in response to an external telemetry device 52. The first housing portion 12' is connected to the second housing portion 12 "by suitable communication means. The second housing section 12 "includes a high voltage power supply 12A, a controller 12C, and a detector 14, as well as an electron beam generator. In the illustrated prior art apparatus, the electron beam generator comprises a photocathode 22 driven by an associated light source driver 55, and a diode laser 56 and associated lens assembly 58. In operation, the laser 56 illuminates the photocathode 22, which in turn generates electrons that are then accelerated toward the anode 24. The anode 24 attracts electrons, which are then directed 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 beam current, and the temporal parameters or to provide preselected voltages, beam currents, and temporal parameters.

As shown in fig. 1(b), external telemetry device 52 and telemetry network 12E may cooperate to allow either dynamic or predetermined external control of power source 12A and temporary parameters. Alternatively, the housing 12 "may not be implanted in the host, and only the detector 14 may extend into the patient. In this case, the controller 12C may be used to directly control the operation of the device, without the need for the network 12E.

In the prior art x-ray source 10 described above, the x-ray emitting elements of the target 26 are adapted to be near or within the irradiation region. To achieve satisfactory x-ray penetration through the body wall to the tumor site, the emitting element close to the target area, e.g. the tumor, does not require the high voltages of prior art machines. The low voltage also focuses the radiation in the target tumor and limits damage to surrounding tissue and surface skin at the point of penetration.

Figure 2 is a schematic block diagram illustrating an overview of one embodiment of a therapeutic radiation source 100 constructed in accordance with the present invention. The therapeutic radiation source 100 includes a laser-heated thermionic cathode, as opposed to prior art therapeutic radiation sources (shown in fig. 1(a) and 1 (b)) that include a resistance-heated thermionic cathode or photocathode. Heating the thermionic cathode 122 with a laser, instead of an electric current, can significantly reduce the power requirements of a therapeutic radiation source 100 constructed in accordance with the present invention.

In general, the therapeutic radiation source 100 includes a radiation generator assembly 101, an optical radiation source 104, and a detector assembly 106. Preferably, the optical radiation source 104 is a laser, and thus, the optical radiation generated by the source 104 is substantially monochromatic and coherent. The laser may be, for example, a diode laser, although other lasers known in the art may be used, such as a Nd: YAG laser, a Nd: YVOt laser, or a molecular laser. Alternatively, other high intensity light sources may be used, such as LEDs (light emitting diodes).

The radiation generator assembly 101 includes an electron source 122, and a target element 128, the target element 128 comprising means for emitting therapeutic radiation in response to incident accelerated electrons. Preferably, the electron source 122 is a thermionic cathode 122. The detector assembly 106 includes an optical transmission 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 that electrons are thermionically emitted.

Figures 3 and 4 show a schematic view of one embodiment of a therapeutic radiation source 100 constructed in accordance with the present invention. In the embodiment shown in FIG. 3, the therapeutic radiation source 100 includes a laser source 104, a detector 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 located in the beam path 109. The therapeutic radiation source 100 further comprises 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 detector assembly 106 connects the laser source 104 and the high voltage power supply 112 to the radiation generator assembly 101. Fig. 3 provides an overall view of the therapeutic radiation source 100, while fig. 4 provides an enlarged view of 1) the radiation generator assembly 101, and 2) the distal end of the detector assembly 106.

Referring to fig. 3 and 4, the probe assembly 106 includes an optical transmission structure 113 having a proximal end 113A and a distal end 113B. The light delivery structure 113 is enclosed within a flexible conductive conduit 105. The distal end 113B of the light delivery structure 113 is affixed to the radiation generator assembly 101. In a preferred embodiment, the light delivery structure 113 is a flexible fiber optic cable that extends from a proximal end 113A to a distal end 113B. In this embodiment, the flexible conduit 105 encasing the fiber optic cable 113 is a small diameter flexible metal sleeve.

In a preferred embodiment, the fiber optic cable 113 includes a conductive outer surface 200. For example, the outer surface of the fiber optic cable 113 can be made conductive by applying a conductive coating. The conductive outer surface 200 of the fiber optic cable 113 provides the connection of the high voltage power supply 112 to the thermionic cathode 122. In this embodiment, the radiation generator assembly 101 also has an electrically conductive outer surface. Preferably, the outer conductive surfaces of the flexible metal sleeve 105 and the radiation generator assembly 101 are set to ground potential to reduce the risk of electrical shock to the device. The flexible sleeve 105 connects the target element 128 and the high voltage power supply 112 to a ground return path, thereby creating 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 metal sleeve 105 may have a diameter of about 1.4 mm. A layer of dielectric material 210 provides insulation between the outer surface of the fiber optic cable 113 and the inner surface of the metal sleeve 105.

As shown in fig. 3 and 4, the radiation generator assembly 101 includes an electron source 122 and a target element 128. A radiation generator assembly 101, for example, about 0.5 to about 2 centimeters in length, extends from the distal end of the probe assembly 106 and includes a sleeve or capsule 130 enclosing the electron source 122 and the target element 128. According to one embodiment, the bellows 130 is rigid in nature and is generally cylindrical in shape. In this embodiment, the cylindrical capsule 130 enclosing the other elements of the radiation generator assembly 101 is considered to provide a substantially rigid enclosure 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 a housing 130, the electron source 122 being disposed at a proximal end of the capsule 130, and the target element 128 being 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. The inner surface of the radiation generator assembly 101 is lined with an electrical insulator or semiconductor, while the outer surface of the assembly 101 is electrically conductive as described above. The low secondary emission, controlled sheet resistance semiconductor film maximizes the high voltage breakdown voltage of the system. According to a preferred embodiment, the radiation generator assembly 101 is sealed to the end of the detector assembly and evacuated. According to another embodiment, the integral probe assembly 106 is evacuated.

In the illustrated embodiment of the invention, the electron source 122 is a thermionic cathode 122 having an electron emitting surface. In an alternative form of the invention (not shown) an annular focus electrode may also be provided. In an alternative embodiment, the thermionic cathode 122 may be positioned near an annular focusing electrode that is at substantially the same potential as the cathode.

In the embodiment shown in fig. 3 and 4, the means for creating an accelerating electric field is a high voltage power supply 112. The power supply 112 has a first terminal 112A and a second terminal 112B, and has a driving device for generating an output voltage between the first terminal 112A and the second terminal 112B. In one form, the power supply 112 can be electrically connected to the target member via first and second terminals. A first terminal 112A of the power supply 112 can be electrically connected to the electron emission surface of the thermionic cathode 122, and a second terminal 112B can be electrically connected 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 metal sleeve 105 to create an acceleration potential difference between the thermionic cathode 122 and the grounded target element 128. In this manner, electrons emitted from the thermionic cathode 122 are accelerated toward the target element 128 and generate an electron beam. The electron beam is preferably thin (e.g., 1mm or less in diameter) and is formed along a beam path 109 that extends to a nominally straight reference axis of 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 meets three criteria: 1) the size is small; 2) high efficiency to enable the use of battery power; and 3) independently variable x-ray tube voltage and current to program the cell for a particular application. 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. High frequency switching mode power converters are preferably used to meet these requirements. The most suitable layout technique for generating low power and high voltage is a resonant voltage converter working in conjunction with a high voltage Cockroft-Walton type amplifier. Low power dissipation, switch mode power supply controller Integrated Circuits (ICs) are currently available to control such layout techniques with few auxiliary components. A more detailed description of typical power supplies suitable for use as power supply 112 is found in U.S. patents 5153900 and 5428658.

The target element 128 is preferably spaced apart from and opposite the electron emission surface of the thermionic cathode 122 and has at least one radiation emitting material adapted to emit therapeutic radiation in response to accelerated electrons incident from the electron emission surface of the thermionic cathode 122. In a preferred embodiment, the emitted therapeutic radiation consists of x-rays, however, it should be noted that the scope of the present 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 with a high Z, x radiation-emitting elemental film or layer, such as tungsten (W), uranium (U), or gold (Au), on the side exposed to the incident electron beam. For example, when electrons are accelerated to 30KeV, a 2 micron thick gold layer absorbs substantially all incident electrons while transporting about 95% of any 30KeV x-rays, about 88% of any 20KeV x-rays, and about 83% of any 10KeV x-rays generated in the layer. In this example, the beryllium (Be) substrate is 0.5mm thick. With this configuration, 95% of the x-rays generated in a direction perpendicular to and toward the beryllium (Be) substrate and passing through the gold layer then pass through the beryllium (Be) substrate and are transmitted outward at the distal end of the detector assembly 106.

In some versions of the invention, the target element 128 may comprise a multi-layer film, and the different layers may have different emission characteristics. For example, a first layer may have a lower energy emission pair energy peak and a second lower layer may have a higher energy emission energy peak. For this form of the invention, a low energy electron beam may be used to generate x-rays in the first layer to achieve the first radiation characteristic and high energy electrons may be used to penetrate the underlying layer to achieve the second radiation characteristic. For example, a 0.5mm wide electron beam can be emitted at the cathode and accelerated to 30KeV, with a lateral electron energy of 0.1eV, and can reach the target element 128 where the beam diameter is less than 1 mm. X-rays are generated within the target element 128 based on a preselected beam voltage, current, and target element composition. The x-rays thus generated pass through a beryllium (Be) substrate with minimal energy loss. As an alternative to beryllium, the target substrate may be made of carbon, a ceramic such as boron nitride, or other suitable material that allows x-rays to pass through with minimal energy loss. The preferred material for the target substrate is carbon in the form of diamond because this material is an excellent thermal conductor. With these parameters, the generated x-rays have sufficient energy to penetrate to a depth of one centimeter or more into soft tissue, with the exact depth depending on the x-ray energy distribution. In another embodiment of the invention, the target may be a solid high-Z material, with the x-rays emitted in a circular beam perpendicular to the tube axis.

In the above-described embodiments, the detector assembly 106, along with its associated radiation generator assembly 101, may be coated with a biocompatible outer layer, such as titanium nitride on a nickel underlayer. For additional biocompatibility, a polyurethane sleeve, for example, may be fitted over the probe.

Figure 5 shows an electron source constructed in accordance with the present invention and including a laser heated thermionic cathode 122. The cathode disk is held in place by end molding or laser welding. The thermionic cathode 122 has an electron emission 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 substrate, such that an electron emissive surface is formed thereon. For example, the substrate 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, for example, a coating of mixed barium and strontium oxides may be applied to a metal substrate such as nickel or a nickel alloy. The metal matrix may be made of other materials including group VI metals, such as tungsten.

Getter 155 may be located within housing 130. The getter 155 helps to form and maintain a high quality vacuum state. The getter has an activation temperature, which will then react with stray gas molecules in the vacuum. It is preferred that the getter used has an activation temperature that is not too high so that the x-ray device is not damaged when heated to the activation temperature.

The fiber optic cable 113 is adapted to transmit laser radiation onto the distal end 113B of the fiber optic cable 113, the laser radiation being generated by the laser source 104 (shown in FIG. 3) and incident on the proximal end 113A of the fiber optic cable 113. The fiber optic cable 113 is also adapted to deliver a beam of transmitted laser radiation to impinge upon the electron emitting surface of the thermionic cathode 122. The laser radiation beam has a power level sufficient to heat at least a portion of the electron emission surface to an electron emission temperature so as to cause thermionic emission of electrons from the surface.

In operation, a laser beam striking the fiber optic cable 113 impinges on the surface of the thermionic cathode 122 and rapidly heats the surface to an electron emission temperature that is below the melting point of the metallic cathode 122. When the surface reaches the electron emission temperature, electrons are thermionically emitted from the surface. The high voltage field between the cathode 122 and the target element 128 (shown in fig. 3 and 4) accelerates these electrons, forcing them to strike the surface of the target element 128 and generate x-rays. In one embodiment of the invention, a Nd: YAG laser is coupled to SiO with a diameter of 400 microns₂Within the optical fiber. A 20KV power supply was used and a thermionic cathode made of tungsten was used. Only a few watts of power are required to generate an electron current greater than 100 mua.

In addition to driving the thermionic cathode with laser energy, another way to increase the efficiency of laser heating the thermionic cathode is to minimize heat loss due to incident laser radiation that remains unabsorbed by the thermionic cathode. Figure 6 shows an embodiment of an electron source embodying the invention which includes a reflector element 160 which reflects incident laser radiation not absorbed by the thermionic cathode 122 back to the thermionic cathode 122. Fig. 6 shows an illustrative incident ray 152 of laser radiation that is not absorbed and scattered by the thermionic cathode 122. Scattered radiation 153 of the laser radiation impinges on the inner surface of the capsule 130 enclosing the radiation generator assembly 101. By positioning the reflector element 160 at a predetermined location along the inner surface of the capsule 130, incident laser radiation that is not absorbed by the electron emitting surface of the thermionic cathode 122 is reflected back to the thermionic cathode 122 by the reflector element 160, thereby effectively forming an optical cavity within the capsule. Thus significantly increasing the coupling efficiency of the laser radiation incident on the thermionic cathode 122.

Figure 7 illustrates an embodiment of the invention in which the therapeutic radiation source includes a spiral-shaped laser-heated thermionic cathode 222. As with the previous embodiments, the capsule 230 defines a substantially evacuated region extending along the beam axis 209 between the cathode 222 at the proximal end of the capsule 230 and the target element 228 at the distal end of the capsule, with the incident laser light being transmitted through the fiber optic cable 213.

The spiral thermionic cathode 222 preferably has a plurality of spaced apart turns defining interstitial spaces between each adjacent turn. Due to the helical configuration of the cathode, heat losses within the cathode due to thermal conduction are minimized. The thermionic cathode 222 includes a spiral-shaped conductive element having a plurality of spaced apart turns defining interstitial spaces between adjacent turns. The conductive element may be, for example, a wire. The conductive element may also be a flat spiral cathode material that is photochemically machined. The spiral configuration of the wires results in a reduction of the conductive heat loss in the cathode.

For a disk-shaped or planar tungsten thermionic cathode, the percentage of incident radiation absorbed at the point of incidence on the cathode is typically about 40%. However, for 40% of the absorption, further heat loss occurs due to heat 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 helical coil having several spaced apart turns. The heat loss through thermal conduction is greatly reduced compared to thermal conduction in 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 previously mentioned, only a few watts of power are required to generate electron currents in excess of 100 μ A for the optically driven thermionic cathode of the present invention, even for a disk-shaped, planar cathode. With an infrared diode laser in combination with a spiral half-millimeter etched cathode, an electron current of about 100 μ A can be obtained with only 180mW of power, thereby greatly reducing the power requirements of the device.

Fig. 8(a) and 8(b) show details of a spiral cathode 300 constructed in accordance with the present invention. Fig. 8(a) shows a plan view of the spiral cathode 300, and fig. 8(b) shows a side view thereof. In a preferred embodiment, the spiral cathode 300 may be fabricated by using photolithography techniques known in the art. The spiral cathode 300 includes conductive elements 310 arranged in a spiral. The material forming the spiral-shaped conductive element is preferably a high melting point metal suitable for withstanding high temperature use. Suitable materials for 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, such as a helical coil, may be used. Spiral coils of different shapes may be used. For example, each of the plurality of spaced apart turns may have a substantially circular shape when viewed longitudinally. Alternatively, the helical coil may have other transverse cross-sectional shapes, such as elliptical, square, or rectangular.

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

In one exemplary embodiment, a 0.002mm thick and 7.4mm long wire is used to fabricate the spiral thermionic cathode 300. In this embodiment, the wire defines two spaced apart loops. Compared to a planar, disk-shaped cathode, in which the power loss due to thermal conduction is about 1.1 watts, the power loss due to thermal conduction of the present invention is only 0.126 watts. The power loss due to thermal radiation is about 140 mW.

FIG. 9 provides an enlarged view of the distal end of a radiation generator assembly and probe assembly constructed according to an embodiment of the present invention, wherein a weakly conductive or semiconductive coating is applied to the inner surface of a rigid capsule enclosing the electron source and target element. As shown in fig. 9, the electron source 208 and the target element 228 are enclosed within a vacuum capsule 230. The inner surface of the capsule is coated with a layer 207 of weakly conductive or semiconductive material. Layer 207 or weakly conductive or semiconductive coating is adapted to provide a substantially smooth voltage gradient within the capsule.

In light-driven small therapeutic radiation sources, such as the device disclosed in U.S. patent 5428658, the inner surface of the capsule 230 is typically lined with an electrical insulator. In contrast, in the embodiment shown in fig. 9, the inner surface is coated with a weakly conductive or semiconductive coating 207. A weakly conductive or semiconductive coating 207 is used to prevent "spiking" in localized high electric field regions or acceleration regions inside the capsule 230, thereby substantially reducing the likelihood of electrical breakdown in the vacuum of the capsule 230. The weakly conductive or semiconductive coating 207 also substantially reduces the likelihood of electrons striking the inner wall of the capsule 230 causing avalanche and ultimately electrical breakdown to produce secondary emissions. The coating also ensures that the electron beam does not deviate from the target by charging effects on the insulating wall.

The illustrated embodiment of the invention features a light-driven high-efficiency therapeutic radiation source that maintains a substantially uniform voltage gradient within the vacuum region between the electron source and the target. A 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. This layer 207 of weakly conductive or semiconductive coating also serves to shield the "triple junction" of the thermionic cathode, i.e., the interface of the cathode, housing wall and vacuum, from high electric fields, thereby preventing electron field emission and subsequent high voltage breakdown.

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

As shown in fig. 10(a), the capsule 230 has an interior surface 310 that defines a hollow evacuated region 312. As described above, the high voltage power supply supplies an acceleration voltage for accelerating electrons emitted from the electron source toward the target element. Since it is safer to keep the target element at ground potential, the thermionic cathode can be negatively biased so that the target element is kept at a net positive voltage relative 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 weakly conductive or semiconductive coating 207. This is because the layer 207 of weakly conductive or semiconductive coating can be voltage gradient controlled in the vacuum region. In other words, 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, as shown by the voltage gradient field lines 330 in FIG. 10 (a). Coating 207 may be made of a weakly conductive or semiconductive material including, but not limited to, chromium oxide, vanadium pentoxide, or an ion-implanted metal such as platinum.

A weakly conductive or semi-conductive high resistance coating 207 is applied to the inner surface of the capsule 230 within the light-driven compact therapeutic radiation source of the present invention, the weakly conductive or semi-conductive high resistance coating 207 improving the ability of the therapeutic source 200 to withstand high acceleration voltages without breakdown. It also greatly enhances x-ray output and stability because the efficiency of electron propagation to the target is greatly increased. This is in contrast to prior art devices such as that disclosed in us patent 5428658 and to compact light driven therapeutic radiation sources that do not include such a coating. These devices include a film of insulating material on the inner surface of the evacuated capsule, rather than a high resistance semiconductor (or weakly conductive) coating, and thus may not be able to control the voltage gradient in the evacuated region.

As previously mentioned, the reliability of the vacuum within the capsule 230 is limited by the operational risk of unpredictable "spikes" or "arcing" between the electrons, which is said to occur when the insulating capability of the vacuum region 312 is suddenly lost. Due to this practical limitation of the insulating capability of the vacuum, local high voltage gradient regions or "spikes" may result. Such spikes can occur in the acceleration region within the evacuated capsule 230 as the emitted electrons accelerate toward the target. To avoid such spikes in the electric field in the evacuated region 312, it is preferable that the inner surface 310 of the capsule 230 be lined with a weakly conductive or semiconducting material that can directly control the electric field. The present invention provides, for example, a low dielectric constant material in the form of a 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 emission, which contribute to the vacuum gap performance and ability to withstand electrical breakdown. The breakdown voltage depends on a number of parameters including, but not limited to, electrode material and geometry, surface treatment and geometry, vacuum quality, and vacuum gap spacing. In particular, the breakdown voltage is generally a function of the dielectric constant of the material forming the capsule 230. It is desirable that the material forming the capsule 230 have a high dielectric strength to withstand large electric fields without breakdown. Preferably, the dielectric strength of the bellows material, e.g., ceramic, is at least 100 KV/mm. Ceramic materials forming the capsule 230 include, but are not limited to, glass, boron nitride, sapphire, quartz glass, and diamond.

The weakly conductive or semiconductive high resistance coating 207 also serves to reduce secondary emissions from the inner surface 310. Secondary emission of electrons striking the walls of the capsule enclosing the acceleration region can lead to avalanches, ultimately leading to breakdown. Such an avalanche is likely to occur when the inner surface 310 of the capsule 230 is lined with an insulating material, such as a material having a secondary emission coefficient greater than 3. A semiconductor coating 207 having a secondary emission coefficient less than 1 is used to prevent such avalanches in an exemplary embodiment. Also, charge up, and subsequent electric field enhancement and breakdown, is eliminated by draining charge through the weakly conductive or semiconductive layer.

FIG. 10(b shows a triple junction 350 of a thermionic cathode the weaker the electric field at the cathode, the stronger the imperfections or irregularities tolerated on the surface of the thermionic cathode without the risk of electrical arcing the triple junction 350 can be shielded from the electric field between the target element and the cathode 222 by a weakly conductive or semiconductive coating, thereby substantially reducing the likelihood of electrical arcing.

Referring to fig. 10(a), the electric field in the vacuum acceleration region is controlled by a resistive weakly conductive or semiconductive coating 207 on the inner surface of the capsule that encapsulates the x-ray generator assembly. The weakly conductive or semiconductive coating 207 creates a controlled voltage gradient across the vacuum within the capsule 230. Also, the weakly conductive or semiconductive coating 207 is adapted to reduce the electric field strength near the triple junction of the thermionic cathode, thereby reducing the likelihood of electrical flashover. Finally, the weakly conductive or semiconductive coating 207 prevents the expansion of secondary emissions of electrons that strike the inner wall of the capsule 230 by absorbing the emitted electrons, thereby preventing an avalanche of emissions that may lead to electrical breakdown. For this reason, the likelihood of electrical flashover or electrical breakdown within the evacuated capsule 230 is substantially reduced, providing significant advantages over light-driven therapeutic radiation sources in which the inner surface of the capsule 230 is lined with an insulating material. Moreover, by ensuring that electrons emitted by the cathode enter the target directly, the weakly conductive or semiconductive coating increases the x-ray generation efficiency and stability, thereby maximizing the likelihood that electrons will strike the x-ray target and maximizing the kinetic energy at which electrons will strike the target.

In summary, the present invention significantly reduces the power requirements of such small therapeutic radiation sources by heating the thermionic cathode with laser energy instead of resistively heating the thermionic cathode with a current. The invention is also characterized by the use of a spiral thermionic cathode configured to minimize energy losses due to incident laser radiation due to thermal conduction within the thermionic cathode. In this way, the power requirements for generating therapeutic radiation in such a small radiation source are further reduced. Finally, the invention also features the use of a weakly conductive or semiconductive coating applied to the inner surface of the evacuated capsule. In this manner, a substantially uniform voltage gradient is formed in the region between the electron source and the target, thus avoiding the formation of high electric field regions and spikes within the evacuated capsule, reducing the likelihood of electrical breakdown, and the electrons being sent directly to the target. The weakly conductive or semiconductive coating also reduces the likelihood of secondary emission of electrons striking the wall of the capsule, thereby avoiding the generation of avalanches that ultimately lead to electrical breakdown. The current from the electric field sweeps, preventing charging and breakdown.

While the present invention has been particularly shown and described with reference to a particularly preferred embodiment, it will 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 transmission structure having a proximal end and a distal end, the optical transmission structure adapted to transmit optical radiation incident on the proximal end to the distal end;

B. a light source comprising means for generating a beam of optical radiation directed towards said proximal end of said optical transmission structure;

C. a radiation generator assembly coupled to the detector assembly and comprising:

a. an electron source for emitting electrons in response to optical radiation transmitted to said distal end of said optical transmission structure, the electron source comprising a thermionic cathode having an electron emitting surface; and

b. a target element comprising at least one radiation emitting material adapted to emit therapeutic radiation in response to accelerated electrons incident from said electron source; and

D. means for providing an accelerating voltage between the electron source and the target element to form an accelerating electric field for accelerating electrons emitted from the electron source towards the target element;

wherein the light delivery structure is adapted to direct a beam of optical radiation delivered therethrough to impinge on a surface of the thermionic cathode; and

the transmitted beam of optical radiation has a power level sufficient to heat at least a portion of the surface to an electron emission temperature, resulting in thermionic emission of electrons from the surface.

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

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

4. A therapeutic radiation source according to claim 3, wherein said metallic material comprises 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, said target assembly being 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 1nA to about 1 mA.

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

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

9. A therapeutic radiation source according to claim 8, wherein said first terminal of said power supply is electrically connected to said electron emitting surface of said thermionic cathode and said second terminal of said power supply is electrically connected to said target element, thereby generating an electric field for accelerating electrons emitted from said electron emitting 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 comprises selectively operable control means for selectively controlling the amplitude of said output voltage.

12. A therapeutic radiation source according to claim 6, further comprising 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 comprises a metal base coated with an oxide.

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

15. A therapeutic radiation source according to claim 1, wherein said light transmission structure comprises a fiber optic cable.

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

17. A therapeutic radiation source according to claim 16, wherein the means for generating an accelerating voltage comprises a power supply having a first terminal and a second terminal, the power supply being electrically connected to the radiation generator assembly via the first and second terminals.

18. A therapeutic radiation source according to claim 16, wherein said electrically conductive conduit is adapted to connect said second terminal of said power source 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 adapted to electrically connect said first terminal of said power source with said thermionic cathode.

20. A therapeutic radiation source according to claim 1, further comprising a substantially rigid capsule, said electron source and said target element being disposed within said capsule, and said capsule defining 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 comprises x-rays.

22. A therapeutic radiation source comprising:

A. a radiation generator assembly, the radiation generator assembly comprising:

a. an electron source for emitting electrons to generate an electron beam along a beam path, said electron source comprising a thermionic cathode having an electron emitting surface; and

b. a target element positioned in the beam path, the target element being spaced apart from and opposing the electron emission surface, the target element comprising at least one radiation emitting element adapted to emit therapeutic radiation in response to accelerated electrons incident from the electron beam; and

c. a substantially rigid capsule, said electron source and said target element being disposed within said capsule, and said capsule defining 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 detector assembly coupled to the radiation generator assembly and including an optical transmission structure having a proximal end and a distal end; said light delivery structure being adapted to deliver laser radiation generated by said source and incident on said proximal end to said distal end and to direct a beam of said delivered laser radiation to said electron emission surface illuminating said thermionic cathode, said beam of laser radiation having a power level sufficient to heat at least a portion of said surface to an electron emission temperature, thereby causing thermionic emission of electrons from said surface;

D. one or more reflector elements disposed at predetermined locations along the inner surface of the capsule, the one or more reflector elements operable to reflect incident laser radiation not absorbed by the thermionic cathode back to the thermionic cathode.

23. A therapeutic radiation source according to claim 1, wherein the power required to heat said electron emitting surface of said cathode to generate an electron beam forming a current of about 100 microamperes ranges between about 0.1 watts and about 3.0 watts.

24. A therapeutic radiation source according to claim 1, wherein said target element is spaced apart from and opposite said electron emitting 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 optic cable and an inner surface of said flexible catheter.

26. A therapeutic radiation source comprising:

A. a radiation generator assembly, the radiation generator assembly comprising:

a. an electron source for emitting electrons to generate an electron beam along a beam path, said electron source comprising a thermionic cathode having an electron emitting surface, and

b. a target positioned in the beam path, the target comprising means for emitting therapeutic radiation in response to accelerated electrons incident from the electron beam; wherein the thermionic cathode comprises a spiral-shaped conductive element;

B. a source of optical radiation; and

C. an optical transmission structure having an initial end and a terminal end, the optical transmission structure being adapted to transmit optical radiation generated by the source and incident on the initial end to the terminal end; wherein the light-transmitting structure is adapted to direct a beam of the transmitted optical radiation onto a surface of the thermionic cathode; and wherein the beam of optical radiation has a power level sufficient to heat at least a portion of the surface to an electron emission temperature, thereby causing thermionic emission of electrons from the surface.

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

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

29. A therapeutic radiation source according to claim 26, wherein said helical 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 adjacent turn.

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

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

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

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

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

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

37. A therapeutic radiation source according to claim 29, wherein said helically shaped electrically conductive coil has a length of from about 2mm to about 7 mm.

38. A therapeutic radiation source according to claim 26, wherein the power required to heat said electron emitting surface of said cathode to generate an electron beam forming a current of about 2 microamperes ranges between about 0.1 watts and about 1.0 watts.

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

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

41. A therapeutic radiation source according to claim 26, wherein the power loss due to thermal conduction is less than 0.2 watts.

42. A therapeutic radiation source comprising:

A. a radiation generator assembly, the radiation generator assembly comprising:

b. a target positioned in the beam path, the target comprising means for emitting therapeutic radiation in response to accelerated electrons incident from the electron beam; and

c. a substantially rigid housing enclosing said thermionic cathode and said target, said housing defining 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. an optical delivery structure having an initial end and a terminal end, the optical delivery structure adapted to deliver optical radiation generated by the source and incident on the initial end to the terminal end, the optical delivery structure adapted to direct a beam of the delivered optical radiation onto a surface of the thermionic cathode;

wherein the beam of the optical radiation has a power level sufficient to heat at least a portion of the surface to an electron emission temperature, thereby causing thermionic emission of electrons from the surface; and

the thermionic cathode includes 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 space between each adjacent turn of said electrically conductive element is substantially eliminated, thereby substantially reducing heat loss from said thermionic cathode due to heat conduction.

44. A therapeutic radiation source according to claim 26, further comprising means for generating an accelerating electric field for accelerating electrons emitted from said electron source towards said target.

45. A therapeutic radiation source comprising:

A. a light source;

B. a detector assembly comprising an optical delivery structure having a proximal end and a distal end, the optical delivery structure adapted to deliver optical radiation incident on the proximal end to the distal end, the optical delivery structure adapted to direct a beam of optical radiation delivered therethrough to impinge on a surface of the thermionic cathode;

a. an electron source for generating an electron beam along a beam path in response to optical radiation transmitted to said distal end of said light-transmitting structure, said electron source comprising a thermionic cathode having an electron-emitting surface; and

b. a target element positioned in the beam path, the target element comprising at least one x-ray emitting material adapted to emit x-rays in response to accelerated electrons incident from the electron source; and

c. a substantially rigid capsule enclosing said electron source and said target element, said capsule defining 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 to form an accelerating electric field for accelerating electrons emitted from said electron source towards said target element, said accelerating voltage at said source having a preselected maximum value relative to a reference potential at said target element;

wherein the transmitted beam of optical radiation has a power level sufficient to heat at least a portion of the surface to an electron emission temperature, thereby causing thermionic emission of electrons from the surface; and

the inner surface of the capsule is coated with a semiconductive coating to provide a substantially smooth voltage gradient within the capsule between the preselected maximum and a 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 from an ion-implanted metal.

48. A therapeutic radiation source according to claim 45, wherein the surface resistivity of said inner surface of said capsule produces lower power dissipation.

49. A therapeutic radiation source according to claim 45, wherein the dielectric constant of the inner surface of the capsule is sufficiently small to prevent spiking within the accelerating electric field.

50. A therapeutic radiation source according to claim 45, wherein the inner surface of said capsule has a secondary electron emission coefficient of 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 shield said thermionic cathode from electrical arcing at the triple junction.

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

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

54. A therapeutic radiation source according to claim 45, wherein said detector assembly comprises 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 emission of electrons impinging on the inner surface of said capsule.

56. A therapeutic radiation source comprising:

A. a light source;

the inner surface of the capsule is coated with a semiconductive coating to provide a substantially smooth voltage gradient within the capsule between the power source and the target element.

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

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