Patent Publication Number: US-6912268-B2

Title: X-ray source and system having cathode with curved emission surface

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
   This is a continuation of application Ser. No. 10/124,864, filed Apr. 17, 2002, now U.S. Pat. No. 6,760,407, which is hereby incorporated by reference. 

   BACKGROUND OF THE INVENTION 
   The present invention relates generally to systems and methods that employ X-ray sources. 
   X-ray sources have found widespread application in devices such as imaging systems. X-ray imaging systems utilize an X-ray source in the form of an X-ray tube to emit an X-ray beam which is directed toward an object to be imaged. The X-ray beam and the interposed object interact to produce a response that is received by one or more detectors. The imaging system then processes the detected response signals to generate an image of the object. 
   For example, in typical computed tomography (CT) imaging systems, an X-ray tube 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 radiation beam 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 tube 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. A group of X-ray attenuation measurements, i.e. projection data, from the detector array at one gantry angle is referred to as a “view”. A “scan” of the object comprises a set of views made at different gantry angles during one revolution of the X-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two-dimensional slice taken through the object. 
   Conventional X-ray tubes comprise a vacuum vessel, a cathode assembly, and an anode assembly. The vacuum vessel is typically fabricated from glass or metal, such as stainless steel, copper or a copper alloy. The cathode assembly and the anode assembly are enclosed within the vacuum vessel. 
   To generate an X-ray beam, the cathode emits electrons which are then accelerated toward the anode, causing the electrons to impact a target zone of the anode at high velocity. The acceleration is caused by a voltage difference (typically, in the range of 20 kV to 140 kV for medical purposes, although possibly higher or lower especially for non-medical purposes) which is maintained between the cathode and anode assemblies. The X-rays emanate from a focal spot of the target zone in all directions, and a collimator is then used to direct X-rays out of the vacuum vessel in the form of an X-ray fan beam toward the patient. 
   In typical X-ray tubes, electrons are emitted from the cathode by a process known as thermionic emission. According to this process, the cathode filament (which is typically formed of a tungsten wire) is provided a current that causes resistive heating of the filament to high temperatures. At such temperatures, the electrons in the filament have sufficient energy that they do not bond to specific atoms (the energy level of the electrons places the electrons in the conduction band) and therefore are susceptible to being emitted from the cathode. A complex focusing structure is used to direct the electrons toward the focal spot. 
   A problem that is therefore encountered is that the cathode is continuously provided with electrical energy which is converted to heat energy, and it is necessary to remove the heat energy from the cathode. Removing heat energy from the cathode is difficult, however, because the cathode is located inside the vacuum vessel and therefore convection is not available as a heat transfer mechanism. Additionally, although conduction is available as a heat transfer mechanism, the large voltage differential that is maintained between the cathode and the anode results in the construction of the cathode being undesirably complex, especially when taken in combination with the complex focusing mechanism that is also provided. A more significant problem is that the heat causes the filament to move (thermal expansion) and changes the location and shape of the focal spot on the target. 
   Therefore, an improved X-ray source which reduces the need for heat transfer away from the cathode and which is relatively simple in construction would be highly advantageous. 
   BRIEF SUMMARY OF THE INVENTION 
   In a first preferred aspect, an X-ray source comprises a cold cathode and an anode. The cold cathode has a curved emission surface capable of emitting electrons. The anode is spaced apart from the cathode. The anode is capable of emitting X-rays in response to being bombarded with electrons emitted from the curved emission surface of the cathode. 
   In a second preferred aspect, an imaging system for imaging an object of interest comprises an X-ray source, a detector array, an image reconstructor, and a display. The X-ray source includes a cold cathode and an anode both of which are disposed within a housing. The cold cathode has a curved emission surface and comprises a plurality of emitters disposed on a substrate. The anode is spaced apart from the cathode, and emits X-rays in response to being bombarded with electrons emitted from the curved emission surface. 
   The detector array comprises a plurality of detector elements which receive the X-rays after the X-rays pass through the object of interest and which generate signals in response thereto. The image reconstructor is coupled to receive the signals from the detector elements, and constructs an image of the object of interest based on the signals from the detector elements. The display is coupled to the image reconstructor and displays the image of the object of interest. 
   Other principle features and advantages of the present invention will become apparent to those skilled in the art upon review of the following drawings, the detailed description, and the appended claims. 

   
     BRIEF DESCRIPTION OF THE DRAWINGS 
       FIG. 1  is a pictorial view of an imaging system; 
       FIG. 2  is a block schematic diagram of the system illustrated in  FIG. 1 ; 
       FIG. 3  is a perspective view of a casing enclosing an X-ray tube insert; 
       FIG. 4  is a sectional perspective view with the stator exploded to reveal a portion of an anode assembly of the X-ray tube insert of  FIG. 3 ; 
       FIG. 5  is a simplified schematic view of a solid state cathode of the X-ray tube of  FIG. 3 ; 
       FIG. 6  is a cross sectional view of a portion of the solid state cathode of  FIG. 5 ; 
       FIG. 7  is a flowchart of the operation of the system of  FIG. 1 ; 
       FIG. 8  is a front view of the solid state cathode of  FIG. 5 ; 
       FIG. 9  is a set of curves showing intensity profiles achievable with the solid state cathode of  FIG. 5 ; 
       FIG. 10  is a schematic view of another solid state cathode; and 
       FIG. 11  is a schematic view of an alternative CT gantry using multiple solid state cathodes. 
   

   DETAILED DESCRIPTION OF THE INVENTION 
   Referring to  FIGS. 1 and 2 , a system  10  that uses an X-ray source  14  is shown. The X-ray source  14  may be used in any application that uses X-rays. For example, in medical applications, the X-ray source may be used to implement a radiography system. In security applications, the X-ray source may be used to implement a baggage checking or other security checkpoint imaging systems. By way of example, the system  10  in  FIGS. 1-2  is a radiography system used for medical imaging, and in particular a computed tomography (CT) imaging system. 
   The CT system  10  includes a gantry  12  representative of a “third generation” CT scanner. The X-ray source  14  is an X-ray tube and is mounted to the gantry  12  and generates a beam of X-rays  16  that is projected toward a detector array  18  mounted to an opposite side of the gantry  12 . The X-ray beam  16  is collimated by a collimator (not shown) to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as an “imaging plane”. The detector array  18  is formed by detector elements  20  which together sense the projected X-rays that pass through an object of interest  22  such as a medical patient. The detector array  18  may be a single-slice detector, a multi-slice detector, or other type of detector. Each detector element  20  produces an electrical signal that represents the intensity of an impinging X-ray beam after it passes through the patient  22 . During a scan to acquire X-ray projection data, the gantry  12  and the components mounted thereon rotate about a gantry axis of rotation  24 . 
   Rotation of the gantry  12  and the operation of the X-ray tube  14  are 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 tube  14  and a gantry motor controller  30  that controls the rotational speed and position of the gantry  12 . A data acquisition system (DAS)  32  in the control mechanism  26  samples analog data from the detector elements  20  and converts the data to digital signals for subsequent processing. An image reconstructor  34  performs image reconstruction (preferably, high speed image reconstruction) based on the signals received from the detector array  18  by way of the DAS  32 . The image reconstructor  34  may be any signal processing device capable of reconstructing images based on signals received from the detector array  18 . 
   A cathode ray tube or other type of display  42  is coupled to the image reconstructor  34  by way of a computer  36 , such that the display  42  is able to receive and display the reconstructed image from the image reconstructor  34 . The computer  36  receives the reconstructed image, stores the image in a mass storage device  38 , and drives the display  42  with signals that cause the display  42  to display the reconstructed image. The images may be displayed as they are acquired or stored for later viewing. The computer  36  also receives commands and scanning parameters from an operator via console  40  that has a keyboard. The operator-supplied commands and parameters are used by the computer  36  to provide control signals and information to the DAS  32 , the X-ray controller  28  and the gantry motor controller  30 . In addition, the computer  36  operates a table motor controller  44  which controls a motorized table  46  to position the patient  22  in the gantry  12 . Particularly, the table  46  moves portions of the patient  22  along a Z-axis through gantry opening  48 . 
   The computer  36  is coupled to a communication interface  50  which connects the computer  36  to a communication network  52 . The communication network  52  may be a local area network, metropolitan area network, or wide area network that connects a group of clinics and/or hospitals. The communication network  52  may also be the Internet. The communication interface  50  is used to transmit medical images or other data acquired using the CT system  10  to other devices on the communication network  52 . The communication interface  50  may also be used to transmit data pertaining to the health and operation of the system  10 , for example, for predictive maintenance or prognostics. The communication interface  50  may also be used to receive control signals from other devices on the communication network  52  which control the system  10 . 
   It should be noted that the embodiment of  FIG. 2  is merely one possible configuration of a CT system that employs the X-ray source  14 . For example, although the X-ray controller and the image reconstructor are both shown as devices which are separate from the computer  36 , it is also possible to integrate the X-ray controller  28  and/or the image reconstructor  34  into the computer  36 . Additionally, as previously noted, the X-ray source could also be used in other applications. 
     FIG. 3  illustrates the X-ray tube  14  in greater detail. The X-ray tube  14  includes an anode end  54 , a cathode end  56 , and a center section  58  positioned between the anode end  54  and the cathode end  56 . The X-ray tube  14  includes an X-ray tube insert  60  which is enclosed in a fluid-filled chamber  62  within a casing  64 . Electrical connections to the X-ray tube insert  60  are provided through an anode receptacle  66  and a cathode receptacle  68 . X-rays are emitted from the X-ray tube  14  through a casing window  70  in the casing  64  at one side of the center section  58 . 
   As shown in  FIG. 4 , the X-ray tube insert  60  includes a target anode assembly  72  and a cathode assembly  74  disposed in a vacuum within a vacuum vessel  76 . The anode assembly  72  is spaced apart from the cathode assembly  74 . A stator  77  is positioned over vessel  76  adjacent to anode assembly  72 . Upon the energization of the electrical circuit connecting anode assembly  72  and the cathode assembly  74 , which produces a potential difference of, e.g., 60 kV to 140 kV, electrons are directed from the cathode assembly  74  to the anode assembly  72 . The electrons strike a focal spot within a target zone  78  of the anode assembly  72  and produce high frequency electromagnetic waves, or X-rays, and residual thermal energy. The target zone  78  emits X-rays in response to being bombarded with electrons emitted from the filament in the cathode assembly  74 . The X-rays are directed out through the casing window  70 , which allows the X-rays to be directed toward the object  22  being imaged (e.g., the patient). 
     FIGS. 5-7  show the cathode assembly  74  in greater detail. As shown in  FIG. 5 , the cathode assembly  74  comprises a cold cathode  79  having a curved surface  80  and which emits electrons to produce an electron beam  82 . In this context, the cold cathode is referred to as such because its operation does not depend on its temperature being above ambient temperature. In practice, typically, the operating temperature of a cold cathode is above ambient temperature, just not as much above ambient temperature as thermionic cathodes. 
   The surface  80  provides a focusing mechanism for the electron beam  82  and preferably has a shape that is optimized in accordance with the geometry of the beam and therefore the desired focal spot. The beam profile may have different shapes, e.g., square, round, hollow, and so on. The shape of the curved emission surface at least partially determines the size and shape of the focal spot on the target zone  78  of the anode assembly  72 . The surface  80  may be curved in two or three dimensions. The surface  80  may, for example, have a parabolic shape or the shape of a portion of a sphere. Alternatively, the surface  80  can be curved along a first axis and straight along a second axis which is orthogonal to the first axis (e.g., cylindrical), curved in two dimensions with different radii in the two directions, or a surface with a variable curvature over its area. 
   The cathode  79  is preferably formed of a monolithic semiconductor. In one embodiment, shown in  FIG. 6 , the cathode  79  is a solid state field emission array fabricated using soft-lithographic patterning on a curved substrate. In other embodiments, the cathode  79  may be fabricated of carbon nanotubes disposed in an array that forms a curved emission surface. Other arrangements could also be used. 
     FIG. 6  is an enlarged view of a portion of the curved surface  80 . The cathode is formed of a plurality of cathode emitters  84  formed on a substrate  86 . The substrate  86  has an insulating layer  90 , a cathode gate film conductor  92 , and a plurality of cones  94 . The insulating layer  90  is preferably discontinuous, i.e., with spaces therebetween. The spaces may have dimensions on the order of 1-3 microns or less. The cones  94  may, for example, be molybdenum cones emitters that are used to generate the electrons. Other materials/structures could also be used, such as Spindt emitters. The cones  94  are preferably disposed with the spaces between the insulating layer so that the cones  94  directly contact the substrate  86 . The gate film  92  may also be formed of molybdenum or other similar metal. In operation, a bias voltage is applied to the gate film  92  to establish an electric field that causes the cones  94  to emit electrons. In one embodiment, by way of example, the cones  94  each have an effective emitting area on the order of about 1×10 −15  cm 2 , such as 1.2×10 −15  cm 2 , and each cone can produce a current up to 1 mA/tip or more when the electric field at its tip is sufficiently large. According to known fabrication techniques, cone packing densities in excess of 1×10 9  cones/cm 2 . Additionally, current densities of over 2400 A/cm 2  are also achievable. Total beam current can be controlled using a low bias voltage such as 120 V DC or below, and preferably down to 20 V DC or lower between the emitters  84  and the gate film  92 . Of course, as improvements are made in soft lithographic techniques, these parameters may be improved upon. 
     FIG. 7  is a flowchart showing an overview of the operation of the system of FIG.  1 . At step  102 , an X-ray beam is generated at the X-ray source  14 . To generate the X-ray beam, a first electric field is applied between the gate film  92  and the emitter cones  94 . The first electric field causes the electrons to be emitted from the emitter cones  94 . The first electric field may be produced by applying a low bias voltage (&lt;50 V) to the gate film  92 . A second electric field is applied between the anode assembly  72  and the cathode  79 . The second electric field causes the electrons to accelerate towards the target zone  78  of the anode assembly  72 . The second electric field may be generated using a voltage in the range of 1 kilovolt to 1000 kilovolts, depending on the application as detailed below. At step  104 , after the X-ray beam passes through at least a portion of the patient or other object of interest  22 , the X-ray beam is detected at the detector array  18 . Then, at step  106 , the image reconstructor  34  constructs an image of a portion of the patient  22  based on data collected during the detecting step  104 . Finally, at step  108 , the image of the portion of the patient  22  or other object of interest is displayed to an operator. 
   As shown in  FIG. 8 , the emitters  84  are disposed in a two-dimensional array. For simplicity, only some of the emitters are shown in FIG.  8 . Preferably, the emitters  84  are arranged in groups with the gate film  92  for each group being electrically isolated from the gate film  92  of each of the remaining groups. In this way, each of the groups of emitters  84  is individually addressable using control lines  96 . Although a group size of one could be used, larger group sizes are preferred in order to simplify construction of the cathode  79 . 
   The emitters  84  are controlled by the X-ray controller  28 . The addressability of the emitters  84  allows a number of features to be implemented by providing different control signals to different ones of the groups of emitters  84 . 
   For example, the X-ray controller  28  is operative to adjust the control signals to the cathode  79  to control the size and shape of the focal spot. The beam shape and size is varied by turning on or off various ones or groups of the emitter  84 . Additionally, the X-ray controller  28  is operative to adjust the control signals to the cathode  79  to control the intensity distribution of the focal spot. Thus, as shown in  FIG. 8 , the focal spot is characterized by an intensity distribution which describes intensity (or current density distribution) of electron bombardment as a function of position ( FIG. 8  shows this for one dimension). Curve  112  shows a typical distribution achievable with a filament; curve  114  shows a gaussian distribution achievable with the cathode  79 ; and curve  116  shows a uniform distribution achievable with the cathode  79 . It is possible to dynamically adjust the focal spot size, shape, and/or intensity distribution of the emitter array depending on which elements are activated and/or the amount of power provided to each element. This can be used to address variabilities in the emitter array associated with manufacturing processes, and to otherwise optimize the beam profile. The current density distribution can also be adjusted as necessary to minimize the heating effects on the target zone  78  of the anode assembly  72 . 
   Additionally, the X-ray controller  28  is operative to adjust the control signals to the cathode  79  as a function of feedback information received by the X-ray controller  28  pertaining to the operation of the imaging system  10 . This allows feedback to be used to maintain the electron beam intensity, size and/or shape to a given specification. The feedback information is acquired during a calibration phase during an initialization procedure for the imaging system  10 . Alternatively, it is also possible to collect such feedback information during normal operation of the system  10 . Such feedback is usable to correct for short and long-term changes in the X-ray source  14 . The ability to control the emitters  84  in this manner allows a smaller, well-defined focal spot to be achieved, thereby improving image quality. 
   Additionally, the X-ray controller  28  is operative to adjust the control signals to the cathode  79  to separately energize multiple groups of the emitters  84  (which may be overlapping). For example, a first set of emitters  84  may be operative to emit a first electron beam having a first focal spot with a first shape, and a second set of emitters may be operative to emit a second electron beam having a second focal spot with a second shape. This allows two different focal spots with different shapes to be produced. This is useful where it is desirable to use the same imaging system  10  for different types of scanning procedures requiring different beam characteristics. 
   Additionally, the X-ray controller  28  is operative to pulse the control signals to the cathode  79  so as to cause the X-rays emitted from the anode to form an X-ray beam that pulsates. The beam current can be switched on and off quickly due to the low (e.g., 50 V or less) bias voltage and low capacitance of the device. Thus, it can be used in applications that require the X-ray beam to have a time structure. For example, in medical applications, when the portion of the patient  22  to be imaged includes a heart, it may be desirable to synchronize activation and deactivation of the cathode  79  to beating of the heart. This may be done, for example, by monitoring an electrocardiograph signal produced in response to beating of the heart. Generally, the electrocardiograph signal is periodic with each cycle corresponding to cycles of the heart. The cathode  79  may then be activated during the same portion of each of the cycles of the heart. Thus, by gating the scan using the ECG signal, the X-ray beam can be turned off except when the patient&#39;s heart is at a predetermined phase of its cycle, thereby reducing the patient&#39;s exposure to X-rays. 
   Additionally, the X-ray controller  28  is operative to control the control signals to the cathode  79  so as to cause the focal spot to wobble back and forth between multiple positions. This is sometimes useful in connection with techniques that use focal spot wobble to eliminate artifacts in the acquired image, currently implemented using multi-filament X-ray sources, magnetic deflection coils or electrostatic deflection plates. 
   In addition to the above-mentioned features, the preferred embodiment of the X-ray source  14  is also relatively simple in construction. The curved geometry eliminates the need for a complicated focusing cup and eliminates strong sensitivity to positional errors and mechanical tolerances. There is also less structure due to reduced need for a heat sink. The curved surface of the cathode  79  combines the focusing and electron emission structures into the same structure. By the use of solid state components, a large vacuum system and complicated beam deflection system is not required. 
   Referring now to  FIG. 10 , another embodiment of a preferred X-ray source  122  that has a curved emission surface  124  is illustrated. In  FIG. 10 , the emission surface  124  has the shape of a portion of a cylinder. This results in a line-focus beam that is focused to a well-defined shape and has a smooth, uniform distribution shape. Again, this geometry eliminates the complicated focusing cup and has the other benefits previously mentioned. 
   Referring now to  FIG. 11 , an interior view of an alternative gantry  132  for the system  10  is illustrated. A series of cold cathode X-ray sources  134  disposed in a ring about the gantry  132  is used to generate respective X-rays, each of which impinges on a corresponding detector array  136 . In  FIG. 11 , for simplicity, only a partial ring of X-ray sources  134  is shown, however, the series of X-ray sources  134  preferably extends around the entire circumference of the gantry  132 . Likewise, for simplicity, only a single detector array  136  is shown. Preferably, however, a series of detector arrays  136  extends around the circumference of the gantry  132 . The detector arrays  136  may be displaced from the X-ray sources  134  along the Z-axis. With this arrangement, rather than have the gantry rotate, each of the X-ray sources is activated sequentially. Thus, the X-ray controller  28  sequentially activates the X-ray sources  134  in a manner that simulates rotation of a single X-ray source about the object of interest. Thus, by avoiding the need for a rotating gantry, the complexity of the computed tomography system is substantially reduced. A rotating anode target, filament heaters, motors and large complex support frames are eliminated. Such a system is also easier to service and, due to its reduced complexity, suffers less downtime in the field. The gantry (along with the X-ray sources and detectors) remains stationary and the patient  22  is imaged without gantry rotation. 
   The X-ray system  10  is particularly suited for medical imaging applications. Medical applications typically accelerate electrons toward the anode assembly  72  by applying an electric field produced with a voltage potential between about 1 kilovolt and 1000 kilovolts and more specifically between about 30 kilovolts and about 160 kilovolts. For example, in mammography and dental applications, a voltage potential of between about 20 kilovolts to 60 kilovolts is used. Cardiography and angiography systems typically use between about 80 to 120 kilovolts. Computed tomography systems typically use between about 80 to 140 kilovolts. 
   Other applications exist for curved surface cathodes. For example, another application is an electron gun that produces hollow beams. Hollow beams are used in gyro-klystron microwave tubes and in wake-field accelerator electron injectors. In each case, a thin shell cylindrical beam is used. A curved surface field emission array with a donut-shaped active area may be used to produce such a beam. Preferably, the curvature is set to produce the correct beam shape in conjunction with the focusing properties of the entire electron gun. Again, the beam area can be moved, changed, or wobbled to meet the needs of the application. Yet another application is electron beam lithography. Electron beam lithography has been proposed as a possible method for fabricating next generation semiconductor chips with features smaller than 0.13 micrometers. Using a field emitter array, the pattern to be projected onto the silicon wafer can be made at the FEA surface by allowing only certain areas to be active. The individual beamlets are transported to the substrate through a focusing structure. Other applications microwave and RF tubes (klystron, gyrotron, and so on), RF electron guns and other electron guns, scanning electron microscopes and other scanning microprobe applications. 
   While the embodiments illustrated in the Figures and described above are presently preferred, it should be understood that these embodiments are offered by way of example only. The invention is not limited to a particular embodiment, but extends to various modifications, combinations, and permutations that nevertheless fall within the scope and spirit of the appended claims.