Abstract:
A solid state x-ray source ( 14 ) for a computed tomograph (CT) imaging system ( 10 ) is presented. X-ray source ( 14 ) has a cathode ( 58 ) which is preferably formed of a plurality of addressable elements. The cathode is positioned within a vacuum chamber ( 74 ) so that electrodes emitted thereby impinge upon anode ( 68 ) spaced apart from cathode ( 58 ). An electron beam ( 82 ) is formed and moved along the length of cathode ( 58 ). The anode ( 68 ) is disposed within a cooling block portion ( 58 ) and operatively adjacent to an x-ray transmissive window ( 66 ). The anode ( 68 ) and x-ray transmissive window ( 66 ) are disposed within an elongated channel ( 64 ) of the cooling block portion ( 56 ).

Description:
TECHNICAL FIELD 
     The present invention relates generally to computed tomograph (CT) imaging and, more particularly, to a x-ray source utilized in connection with CT systems. 
     BACKGROUND ART 
     In at least some computed tomograph (CT) imaging system configurations, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the “imaging plane”. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile. 
     In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal spot. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector. A scintillator is located adjacent the collimator, and photodiodes are positioned adjacent the scintillator. 
     Multislice CT systems are used to obtain data for an increased number of slices during a scan. Known multislice systems typically include detectors generally known as 3-D detectors. With such 3-D detectors, a plurality of detector elements form separate channels arranged in columns and rows. Each row of detectors forms a separate slice. For example, a two slice detector has two rows of detector elements, and a four slice detector has four rows of detector elements. During a multislice scan, multiple rows of detector cells are simultaneously impinged by the x-ray beam, and therefore data for several slices is obtained. 
     A system that does not require a rotating x-ray source is described in U.S. Pat. Nos. 4,521,900 and 4,521,901. In the &#39;900 patent, a large vacuum chamber is used which incorporates an electron gun and ring-shaped targets to produce x-rays. The electron beam emerges from the gun several feet away from the patient, travels a bent path to move toward the targets then hits the material to produce x-rays. The single fairly high power electron beam sweeps out a circle, a ring that surrounds the patient, to produce the “scan” effect. One drawback to such a system is that a large vacuum system to enclose the electron beam&#39;s path or trajectory is required, and further, a complicated beam deflection system is employed to accurately steer the beam. 
     Accordingly, it would be desirable to provide a CT scanner and CT scanner system that provides a x-ray source that reduces the complexity of the scanning system and does not require a rotating x-ray source. 
     SUMMARY OF THE INVENTION 
     It is therefore one object of the invention to provide a solid state x-ray tube to reduce the complexity of the x-ray tube. In one aspect of the invention, a CT system comprises a solid state x-ray source for a computed tomograph (CT) imaging system is illustrated. X-ray source has a cathode which is preferably formed of a plurality of addressable elements. The cathode is positioned within a vacuum chamber so that electrons emitted thereby impinge upon anode spaced apart from cathode. An electron beam is formed and moves along the length of cathode. The anode is disposed within a cooling block portion and operatively adjacent to an x-ray transmissive window. The anode and x-ray transmissive window are disposed within an elongated channel of the cooling block portion. 
     Advantageously, the present invention uses cold-cathode technology. The employment of cold-cathode technology allows the possibility for an electron beam source to be turned on and off very quickly with the limitation being the switching speed of the associated electronic and optical circuitry. In addition, fast electronic gating circuits may allow many of these emitting sources to be switched sequentially, thus allowing an electron beam to sweep a target. Such technology will allow the typically rotating x-ray source in a CT system to be removed which substantially removes the complexity associated therewith. For example, bearing issues, target balancing problems, and Z-axis growth problems are associated with prior known CT systems. Also, the prior known systems are complex to service. 
     Another advantage of the invention is that the use of solid state components eliminates the need for a large vacuum system and a complicated beam deflection system. Other eliminated features compared to the prior art include not requiring a rotating target, a filament heater circuit and motors, and the large support frames associated with a rotating target. 
     Another advantage of the invention is that because of the fast scan times, applications that require fast scan times such as cardiac imaging may be employed. 
     Yet another advantage of the invention is that slip rings commonly used in a rotating system may be eliminated. Slip ring connections typically introduce noise and complexity into the transmission of signals obtained from the detectors as well as transmitting power and high voltage to the x-ray source. 
     Another advantage of the invention is that because of high heat dissipation enabled with a stationary anode, the usual massive target with massive graphite backing is not required to store heat generated by electrons that come to rest in the target. Furthermore, this will greatly reduce or eliminate the need to wait for the X-ray tube to cool. 
     Another advantage of the invention is that shielding necessary for canceling the effect of the earth&#39;s magnetic field is not required. 
     Other objects and advantages of the present invention will become apparent upon the following detailed description and appended claims, and upon reference to the accompanying drawings. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a pictorial view of a CT imaging system according to the present invention. 
     FIG. 2 is a block schematic diagram of the system illustrated in FIG.  1 . 
     FIG. 3 is a cross-sectional view of a solid state x-ray tube according to the present invention. 
     FIG. 4 is a cross-sectional view of an alternative embodiment according to the present invention. 
     FIGS. 5A and 5B are diagrammatic views of a scan using multiple x-ray tubes and detectors according to the present invention. 
     FIG. 6 is a cross-sectional view of an alternative embodiment of the present invention. 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
     Referring to FIG. 1, a computed tomography (CT) imaging system  10  is shown as including a gantry  12  representative of a “third generation” CT scanner. The gantry  12  has an x-ray source  14  that projects a beam of x-rays  16  toward a detector array  18  on the opposite side of the gantry  12 . 
     The detector array  18  is formed by a plurality of detection elements  20  which together sense the projected x-rays that pass through a medical patient  22 . Each detection element  20  produces an electrical signal that represents the intensity of an impinging x-ray beam and hence, the attenuation of the beam as it passes through the patient  22 . During a scan to acquire x-ray projection data, the housing  12  and the components mounted thereon rotate about a center of gravity. 
     The operation of the x-ray source  14  is governed by a control mechanism  26  of the CT system  10 . The control mechanism  26  includes an x-ray controller  28  that provides power and timing signals to the x-ray source  14 . A data acquisition system (DAS)  32  in the control mechanism  26  samples analog data from the detection elements  20  and converts the data to digital signals for subsequent processing. An image reconstructor  34  receives sampled and digitized x-ray data from the DAS  32  and performs high speed image reconstruction. The reconstructed image is applied as an input to a computer  36  which stores the image in a mass storage device  38 . 
     The computer  36  also receives and supplies signals via a user interface or graphical user interface (GUI). Specifically, the computer  36  receives commands and scanning parameters from an operator console  40  that preferably includes a keyboard and mouse (not shown). An associated cathode ray tube display  42  allows the operator to observe the reconstructed image and other data from the computer  36 . The operator supplied commands and parameters are used by the computer  36  to provide control signals and information to the x-ray controller  28 , the DAS  32 , and a table motor controller  44  in communication with a table  46  to control operation of and movement of the system components 
     Referring now to FIGS. 3 and 4, a respective longitudinal cross-sectional view and lateral cross-sectional view of an x-ray source is illustrated. X-ray source  14  has a housing  50  that is sealed to provide a vacuum therein. Housing  50  has a support frame  52  positioned therein. Support frame  52  is preferably comprised of an insulative material such as alumina. 
     Housing  50  has a support portion  54  and a cooling block portion  56 . Support portion  54  is preferably formed from an insulative material such as a high voltage epoxy compound. Various types of compounds would be evident to those skilled in the art. Cooling block portion  56  is thermally conductive and electrically conductive. Cooling block portion  56  is preferably formed of copper. 
     Support portion  54  is generally an elongated semi-tubular shape. As illustrated, support portion  54  is u-shaped. Support portion  54  is used to position a cathode  58  for generating electrons. Cathode  58  may be supported by a cathode support portion  60  which is integrally molded with support portion  54 . Cathode support portion  60  extends a predetermined distance D from a back wall  62  of support portion  54 . The distance D may be adjusted depending on the desired characteristics of the materials used and output desired. 
     Cooling block portion  56  has an elongated channel  64  or beam opening extending therethrough. Elongated channel  64  has an x-ray transmissive window  66  disposed therein. X-ray transmissive window  66  preferably completely fills elongated channel  64 . X-ray transmissive window  66  is preferably formed from an electrically conductive material and a thermally conductive material such as a carbon-based material like graphite. Also, it is preferred that the atomic mass or “Z” of the x-ray transmissive window  66  is relatively low. Other suitable materials known to those skilled in the art include beryllium. 
     An anode  68  is formed directly and operatively adjacent to x-ray transmissive window  66 . Preferably, anode  68  is formed of a thin metallic layer  68  or foil. The thin film anode  68  is preferably formed of a high atomic weight material such as tungsten or uranium. Of course, those skilled in the art will recognize that preferably the highest atomic weight material is used but there may be a trade-off between physical dimension, strength-to-weight ratio, and x-ray production. Anode  68  may be formed as a thin film which is deposited directly onto window  66 . Because of the x-ray process, heat may be generated at anode  68  and therefore anode  68  is preferably thermally coupled to cooling block portion  56 . Anode  68  may also be formed from a relatively thin layer of tungsten or tungsten alloy (2 to 30 microns) on a copper substrate. 
     Cooling block portion  56  preferably has a plurality of cooling tubes  70  extending therethrough. Cooling tube  70  provides cooling fluid or air therethrough to reduce the temperature of cooling block portion  56  and ultimately the temperature of anode  68 . Preferably, cooling tube  70  extend substantially the length L of x-ray source  14 . Elongated channel  64  is defined in cooling block portion  56  by shoulders  72  that extend in an inward direction toward cathode  58 . As will be further described below, shoulders  72  help provide a conductive path for electrons passing through anode  68 . 
     Support portion  54  and cooling block portion  56  define a vacuum chamber  74  therein. Vacuum chamber  74  preferably extends substantially the length of support portion  54  and cooling block portion  56  short of any end wall structures. Vacuum chamber  74  is preferably actively pumped so that the vacuum is always at an optimum level. This will reduce high voltage instabilities. 
     Cathode  58  has a plurality of gating connections  76  coupled thereto. Gating connections  76  control the turning on and off of cathode  76 . High voltage input  78  is coupled to cathode  58  to provide the necessary potential for the generating of electrons. Both gating connections  76  and high voltage input  78  may be formed through support portion  54 . 
     Cathode  58  is preferably formed of an elongated array of electron emitters. Various types of emitters may be used. For example, ferro-electric emitters may be used to create an electron emission in the form of a small, relatively narrow beam width that will impinge on anode  68 . Another type of cathode that may be used is a thin film emission cathode. Such technology is similar to that used in flat panel monitors and television sets. Photo emitters may also be used for cathode  58 . Photo emitters may, for example, use compact laser diode arrays. Emission occurs according to the order in which the laser beams of sufficient power and proper wavelength “address” the emitters by raster scanning of emitters which are arranged across a face of a flat panel plane or arranged on a bar that scans or moves across the face of the device. The photo emitters may also be in the form of a line or series of smaller dimension standalone emitter batches that would emit in a pattern corresponding to the emitters that have been addressed to be emitting. In all of the embodiments, cathode  58  may be formed of a plurality of emitters  80  best shown in FIG.  3 . Cathode elements  80  are preferably addressable meaning that they may be selectively turned on and off to form the electron beam. With respect to emitters, photo emitters emit electrons when light reaches the solid state device capable of releasing the electrons into vacuum chamber  74 . Light emission from photo laser devices such as solid state lasers and the like have been controlled to within micro or nano seconds. Laser devices can produce high efficiencies of photo emission. Preferably the addressability is sequential and allows the beam formed at one end to effectively move across the cathode in a scanned manner. Light signal switching devices such as micro-machined mirrors onto a solid state monolithic substrate may also be used. Light may also be delivered using a fiber optic or free beam means. For example, a six micro amp electron beam is produced for every milliwatt of laser light at 1% quantum efficiency using a gallium arsenide laser with 780 nanometer light where circular polarization is used for polarized electron delivery. Infrared laser bars are also commercially available in the 1 watt to 10 watt power range corresponding to 6 to 60 milliamps of electrons at 1% QE. 
     Those skilled in the art will recognize that polarized elections are not required in this invention. This corresponds to the efficiency of the photo emitters. 
     In operation, the desired emitters  80  are turned on as addressed to generate the desired beam  82 . Preferably, electron beam  82  starts at one end of cathode and works its way across the cathode generating x-rays in a linear or sequentially moving manner. The electrons are released from cathode  58  and travel toward anode  68 . When the electrons impinge upon anode  68 , x-rays are released through window  66 . Heat that formed in anode  68  is thermally coupled into shoulder  72  of cooling block portion  56 . Heat may also be formed in window  66  which is also thermally conducted to cooling block  56 . Heat is removed from cooling block  56  through cooling channel  70  which may be provided with cooling fluid or air. 
     It is possible for some electrons to travel through anode  68  and enter window  66 . Because window  66  is preferably electrically conductive, electrons entering window  66  are electrically conducted to cooling block  56  and may re-enter beam  82  through shoulder  72 . This is illustrated as stray electron paths  84 . The paths  84  complete the electrical loop back to anode  68 . 
     Preferably, the length of cathode  58 , the length of anode  68 , the length of window  66 , and the length of elongated channel  64  are all substantially the same and are preferably just short of or about length L. 
     Referring now to FIGS. 5A and 5B, an interior view of a gantry  12  is illustrated. In FIG. 5A, a first, second and third x-ray source  14 A,  14 B, and  14 C are used to generate respective x-rays  16 A,  16 B, and  16 C. Each of x-rays  16 A,  16 B, and  16 C impinge upon a corresponding detector  88 A,  88 B, and  88 C. By using a device without a rotating gantry, each of the x-ray sources  14 A- 14 B, x-rays  16 A- 16 B, and detectors  88 A- 88 C are relatively fixed. In such an embodiment, the beams merely scan the length of the cathode without actually physically moving the x-ray source or detectors. In such a manner, the non-rotating complexity in prior known systems is substantially reduced. Also, such systems are believed to be substantially faster in the generation of an image. 
     Referring now to FIGS. 5A and 5B, five x-ray sources,  14 ′A- 14 ′E are illustrated generating x-rays  16 ′A,  16 ′B,  16 ′C,  16 ′D, and  16 ′E toward detectors  88 ′A,  88 ′B,  88 ′C,  88 ′D, and  88 ′E in a segmented manner. Of course, those skilled in the art would recognize that a continuously tube formed according to the teachings herein could also be used. A heart  90  is used to illustrate that a CT system formed according to the present invention may be sized to be tailored for the organ or body part to be imaged. By the use of solid state components in the present invention, a large vacuum system and complicated beam deflection system is not required. A rotating anode target, filament heaters, motors and large complex support frames are also eliminated from the design. Such a system is also easier to service and is predicted to reduce downtime in the field. 
     Faster scan times because of the ease in scanning the beam may also be acquired. This allows for such imaging as cardiac imaging. Power levels are reduced because the system may be positioned closer to the patient. Because intensity falls off inversely with respect to the square of the distance from the patient. For example, it is predicted that by using the teachings of the present invention the diameter of the CT system may be reduced by 20% while the required current level may be reduced by 36%. By building small units tailored to the particular applications, such as brain scans or cardiac scans, better resolution, faster patient throughput and lower cost to specialist treatment centers may be provided. 
     The cold cathode technology is particularly useful for instant startup, long life and low power consumption with rapid switch on/switch off. 
     By providing the multiple beams as illustrated in FIGS. 5A and 5B, the temperatures under a given electron beam may be reduced and yet produce an overall higher system imaging power. The fixed position of the x-ray source allows the beams to be switched on and off in rapid succession to eliminate problems with x-ray scanner normally associated with CT systems. 
     In alternative embodiments, the anode may be formed in various manners, including an application of metal such as tungsten on a layer of copper. Other such anode assemblies may include a sandwich-type target using alternating layers of tungsten or rhenium and another material such as graphite. The size of the layers may be approximately 1 to 5 microns for the tungsten and can be sized to minimize the temperature of the focal spot and maximize x-ray output depending on the particular application. The graphite layers or other suitable materials will allow the passage of electrons and x-rays. 
     The present invention allows conventional convection cooling due to the stationary anode. The large size of the target ring associated with such a device has a large surface area and thus the heat transfer coefficient does not need to be extremely high. Also, the beam fan angle may be decreased in particular applications to decrease the focal spot temperature. 
     Referring now to FIG. 6, an alternative embodiment of a cylindrical tube around a center line  91  is shown in cross section. In this embodiment, cathode  92  is positioned at an angle relative to anode  94 . That is, an electron beam  96  from cathode  92  hits anode  92  at a predetermined angle range of 15-60 degrees. In this illustration, 20 degrees is used. Cathode  92  may have a tungsten coil  100 . However, cathode  92  may be also be formed of a cone-shaped field emitter, a hollow cylinder emitter, a carbon nanotube emitter, a photo emitter or other type of emitter known to those in the art. The type of emitter may depend on the particular system application or performance requirements. Anode  94  may, for example, be a patch of tungsten or rhenium. The back-scattered electrons  104  strike part of anode  94  as well as copper cooling plate  106 . Lines  108  indicate X-rays generated from anode  94 . Cooling block  106  has an X-ray transmissive window  110 , preferably formed of beryllium (Be) and coolant channels  112 . The anode  94  and cathode  92  are separated in space and potential by insulator  114  which forms a portion of a housing  116  together with cooling block  106 . The anode  94  in this embodiment is operatively coupled to window  110  but is separated therefrom, in contrast to the previous embodiment. 
     While the invention has been described in connection with one or more embodiments, it should be understood that the invention is not limited to those embodiments. On the contrary, the invention is intended to cover all alternatives, modifications, and equivalents, as may be included within the spirit and scope of the appended claims.