JP2013000233A - X-ray ct apparatus - Google Patents

Abstract

PROBLEM TO BE SOLVED: To improve the photography accuracy in dual energy scanning.SOLUTION: An X-ray tube of the X-ray CT apparatus includes a first cathode, a second cathode and an anode. The first cathode is set at a first tube voltage to emit electrons. The second cathode is set at a second tube voltage different from the first tube voltage in a direction facing the first cathode to emit electrons. The anode is disposed between the first cathode and the second cathode, and includes a first irradiation surface to be irradiated with the electrons from the first cathode and a second irradiation surface to be irradiated with the electrons from the second cathode which are disposed back to back. The first irradiation surface of the anode has a first target angle relative to the irradiation with the electrons from the first cathode, and the second irradiation surface has a second target angle different from the first target angle relative to the irradiation with the electrons from the second cathode to emit X rays toward a subject.

Description

  Embodiments described herein relate generally to an X-ray CT apparatus.

  In recent years, X-ray CT (computed tomography) devices that visualize the inside and inside of a body have been used for medical diagnosis and industrial nondestructive inspection. In the diagnostic X-ray CT apparatus, a dual energy scan that switches between a low tube voltage (for example, 80 [kV]) and a high tube voltage (for example, 140 [kV]) at high speed is executed during scanning, and different energy distributions are obtained. A dual energy image with an X-ray beam having

  There is an attempt to separate a calcified tissue part and a blood vessel image by a contrast agent by analyzing a dual energy image obtained by a dual energy scan using an X-ray CT apparatus. As a method of switching the output voltage of the X-ray high-voltage power supply, two means for setting the tube voltage are provided, and a low tube voltage and a high tube voltage are set in time with the data acquisition system (DAS). Some switch the signal to change the output voltage of the high-voltage power supply.

JP 2004-363109 A

  In the dual energy scan according to the prior art, since the output voltage of the high voltage power supply is changed according to the tube voltage, it takes time to switch between the low tube voltage and the high tube voltage, and the change also becomes continuous. When the tube voltage is switched for each view, the waveform becomes like a triangular wave. Therefore, the energy distribution of the X-ray beam also changes continuously accordingly, so that the target tissue is not completely separated from the dual energy image.

  In addition, since the X-ray tube usually has one cathode and the heating is controlled by the tube current of the one cathode, the tube current cannot be switched at high speed. The tube current is determined according to the emission characteristics of the tube. Therefore, in the case of a high tube voltage, a larger tube current flows than in the case of a low tube voltage. However, switching of the tube current is not always performed appropriately with respect to switching of the tube voltage. An object of this embodiment is to improve the imaging accuracy of the dual energy scan.

  An X-ray CT apparatus according to one embodiment includes an X-ray tube and a detector. The X-ray tube has a first cathode, a second cathode, and an anode. The first cathode is set to a first tube voltage and emits electrons. The second cathode is set to a second tube voltage different from the first tube voltage and irradiates electrons in a direction opposite to the first cathode. The anode is disposed between the first cathode and the second cathode, and is irradiated with electrons from the first irradiation surface irradiated with electrons from the first cathode and the second cathode. A second irradiation surface is provided back to back. In the anode, the first irradiation surface has a first target angle with respect to the electron irradiation from the first cathode, and the second irradiation surface has an electron emission from the second cathode. A second target angle different from the first target angle with respect to the irradiation is applied, and X-rays are irradiated toward the subject. The detector detects the X-ray transmitted through the subject.

  An X-ray CT apparatus according to one embodiment includes an X-ray tube, a detector, and control means. The X-ray tube has a first cathode, a second cathode, and an anode. The first cathode is set to a first tube voltage and emits electrons. The second cathode is set to a second tube voltage different from the first tube voltage and irradiates electrons in a direction opposite to the first cathode. The anode has a first irradiation surface on which electrons are irradiated from the first cathode and a second irradiation surface on which electrons are irradiated from the second cathode, back to back. X-rays are emitted toward The detector detects the X-ray transmitted through the subject. The control means holds the X-ray tube integrally with the detector facing the detector, rotates the subject around the subject, and sets the set of the X-ray tube and the detector with respect to the top plate in the direction of the rotation center axis. The relative movement is controlled, and an integral multiple of the distance of the relative movement during the imaging interval from the imaging with one cathode to the imaging with the other cathode in continuous imaging is set as the first irradiation surface and the second Match the length between the irradiated surfaces.

The block diagram of the X-ray CT apparatus of this embodiment. The block diagram of the X-ray generator of this embodiment. Schematic of the X-ray tube of this embodiment. Schematic of the X-ray tube and detector of this embodiment. FIG. 4 is a detailed enlarged view of anodes having different target angles. The figure explaining the outline | summary of the X-ray irradiation from the X-ray tube of this embodiment. The schematic diagram explaining the X-ray irradiation position and the synchronous movement control of a top plate.

An embodiment of an X-ray CT apparatus according to the present invention will be described with reference to the accompanying drawings.
FIG. 1 is a block diagram of the X-ray CT apparatus of the present embodiment. An X-ray CT apparatus 1 shown in FIG. 1 includes a scanner device 11 and an image processing device 12. The scanner device 11 of the X-ray CT apparatus 1 is usually installed in an examination room and is configured to generate X-ray transmission data relating to an imaging region of a subject (human body) O. On the other hand, the image processing apparatus 12 is usually installed in a control room adjacent to the examination room, and is configured to generate projection data based on transmission data and generate / display a reconstructed image.

  The scanner device 11 of the X-ray CT apparatus 1 includes an X-ray generator 21, a detector 22, a diaphragm 23, a DAS (data acquisition system) 24, a rotating unit 25, a controller 26, a diaphragm driving device 27, a rotational driving device 28, and a ceiling. A plate 29 and a top plate driving device (bed device) 30 are provided.

  The X-ray generator 21 is mainly composed of an X-ray tube 31, a high voltage power supply unit 32, and a FHC (filament heating circuit) unit 34 (shown in FIG. 2). The X-ray tube 31 irradiates the detector 22 with X-rays according to the tube voltage supplied from the high voltage power supply unit 32. Fan beam X-rays and cone beam X-rays are formed by X-rays emitted from the X-ray tube 31. The high voltage power supply unit 32 supplies power necessary for X-ray irradiation to the X-ray tube 31 under the control of the controller 26.

  The detector 22 is a one-dimensional array type X-ray detector having a plurality of rows of X-ray detection elements in the channel direction and one column in the slice direction (column direction). Alternatively, the detector 22 is a two-dimensional array type detector 22 (also referred to as a multi-slice detector) having a matrix, that is, a plurality of rows of X-ray detection elements in the channel direction and a plurality of columns in the slice direction. . The detector 22 detects X-rays irradiated from the X-ray tube 31 of the X-ray generator 21 and transmitted through O through the subject.

  The diaphragm 23 adjusts the irradiation range in the slice direction of the X-rays emitted from the X-ray tube 31 by the diaphragm driving device 27. That is, the X-ray irradiation range in the slice direction can be changed by adjusting the aperture of the diaphragm 23 by the diaphragm driving device 27.

  The DAS 24 amplifies the transmission data signal detected by each X-ray detection element of the detector 22 and converts it into a digital signal. The output data of the DAS 24 is supplied to the image processing device 12.

  The rotating unit 25 is housed in a gantry (not shown) of the scanner device 11 and holds the X-ray tube 31, the detector 22, the diaphragm 23, and the DAS 24 of the X-ray generator 21 as a unit. The rotating unit 25 is configured so that the X-ray tube 31, the detector 22, the diaphragm 23 and the DAS 24 can be rotated together around the subject O in a state where the X-ray tube 31 and the detector 22 face each other. Yes. As described above, the rotating unit 25 holds the X-ray tube 31 integrally with the detector 22 facing the detector 22 and rotates around the subject.

  The controller 26 includes a CPU (central processing unit) and a memory. The controller 26 controls the X-ray generator 21, DAS 24, aperture driving device 27, rotation driving device 28, top plate driving device 30, and the like based on the control signal input from the image processing device 12, and energy. A scan such as a dual energy scan is performed while switching X-rays having different distributions.

  The diaphragm driving device 27 adjusts the irradiation range of the diaphragm 23 in the X-ray slice direction under the control of the controller 26. Under the control of the controller 26, the rotation driving device 28 rotates the rotating unit 25 so that the rotating unit 25 rotates around the cavity while maintaining the positional relationship. The top plate 29 can place the subject O thereon.

  The top plate driving device 30 moves the top plate 29 along the z-axis direction under the control of the controller 26. The central portion of the rotating unit 25 has an opening, and the subject O placed on the top plate 29 of the opening is inserted. A direction parallel to the rotation center axis of the rotating unit 25 is defined as a z-axis direction, and a plane orthogonal to the z-axis direction is defined as an x-axis direction and a y-axis direction.

  Each cross section of the subject O is imaged along the z-axis. For this purpose, the rotating unit 25 and the top plate 29 are relatively moved. As described above, the rotating unit 25 is stopped in the z-axis direction by the top plate driving device 30, and this relative movement can be performed by the movement of the top plate 29. On the other hand, the top plate 29 can be stationary in the z-axis direction, and relative movement can be performed by moving the rotating unit 25 in the z-axis direction. The movement of the rotating unit 25 in the z-axis direction is executed by the rotation driving device 28. Therefore, the rotation drive device 28 or the top plate drive device 30 controls the relative movement between the rotation unit 25 and the top plate 29. Note that the rotary unit 25 may be moved by the rotary drive device 28 and the top plate 29 may be moved by the top plate drive device 30 and moved relative to each other.

  The image processing apparatus 12 of the X-ray CT apparatus 1 is configured based on a computer, and can communicate with a network N such as a hospital basic LAN (local area network). Although not shown, the image processing apparatus 12 includes basic hardware such as a CPU, a memory, an HDD (Hard Disc Drive), an input device, and a display device.

  The image processing apparatus 12 generates projection data by performing logarithmic conversion processing and correction processing (preprocessing) such as sensitivity correction on the raw data input from the DAS 24 of the scanner device 11. Further, the image processing device 12 performs scattered radiation removal processing on the preprocessed projection data. The image processing device 12 removes scattered radiation based on the value of projection data within the X-ray irradiation range, and is estimated from the size of the projection data to be subjected to scattered radiation correction or the value of the adjacent projection data. The scattered radiation correction is performed by subtracting the scattered radiation from the target projection data. The image processing device 12 generates a reconstructed image based on the corrected projection data. That is, the image processing apparatus 12 has a function as a reconstruction unit that reconstructs an image from projection data based on X-rays detected by the detector 22.

  In the X-ray CT apparatus of the present embodiment, an X-ray tube 31 and a detector 22 are integrated with a rotation / rotation (rotate / rotate) type in which the periphery of a subject is rotated, and a large number of detection elements are arrayed in a ring shape. There are various types such as a stationary / rotating type in which only the X-ray tube rotates around the subject, and any type is applicable. Here, the rotation / rotation type is described.

  In addition, the mechanism for converting incident X-rays into electric charges is based on an indirect conversion type in which X-rays are converted into light by a phosphor such as a scintillator and the light is further converted into electric charges by a photoelectric conversion element such as a photodiode. The generation of electron-hole pairs in semiconductors and their transfer to the electrode, that is, the direct conversion type utilizing a photoconductive phenomenon, is the mainstream.

  In addition, in recent years, a so-called multi-tube type X-ray CT apparatus in which a plurality of pairs of the X-ray tube 31 and the detector 22 are mounted on a rotating ring has been commercialized, and development of peripheral technology has been advanced. . The X-ray CT apparatus of the present embodiment can be applied to both a conventional single-tube type X-ray CT apparatus and a multi-tube type X-ray CT apparatus. Here, a single-tube X-ray CT apparatus has been described.

  FIG. 2 is a block diagram of the X-ray generator of this embodiment. As shown in FIG. 2, the X-ray tube 31 of the X-ray generator 21 has an X-ray tube envelope (insert valve) 41 whose inside is maintained at a high vacuum. Examples of the material for the insert valve 41 include metal, glass, and ceramic. The X-ray tube 31 encloses an anode (anode) 42 and a cathode (cathode) unit 43 in an insert valve 41. The cathode unit 43 includes two cathodes 43a and 43b.

  When a current is passed through the cathode 43a of the cathode unit 43 in a vacuum, the cathode 43a is heated and emits thermoelectrons. At this time, if a positive voltage is applied to the anode 42 side with respect to the cathode 43 a, the thermoelectrons emitted from the cathode 43 a are attracted by positive charges and fly toward the anode 42. As a result, electrons flow from the cathode 43a toward the anode. That is, a current flows from the anode 42 toward the cathode 43a. The cathode 43b of the cathode unit 43 operates in the same manner as the cathode 43a.

  Thus, the cathode 43a is set to the first tube voltage and radiates thermionic electrons. The cathode 43b is set to a second tube voltage different from the first tube voltage in the direction facing the cathode 43a and radiates thermoelectrons. The facing direction means that the emission direction of the thermoelectrons from the cathode 43a and the emission direction of the thermoelectrons from the cathode 43b are opposite, and includes that the direction of the vector is opposite. However, it is not limited to the case where the two directions are completely the same except that they are opposite to each other, and it means that the two directions are facing each other.

  On the other hand, the anode 42 is disposed between the cathode 43a and the cathode 43b. The emitted electron flow from both directions collides with the anode 42. Due to the collision, the anode 42 emits X-rays toward the subject.

  The FHC unit 34 includes an FHC 34a corresponding to the cathode 43a and an FHC 34b corresponding to the cathode 43b. The FHC 34a includes a heating transformer (not shown), and the secondary side of the heating transformer is electrically connected to both ends of the cathode 43a. The FHC 34b includes a heating transformer (not shown), and the secondary side of the heating transformer is a cathode. It is electrically connected to both ends of 43b. Although the case where an AC power supply is used for the FHC 34a and the FHC 34b has been described, a DC power supply may be used. The controller 26 supplies required tube currents to the cathodes 43a and 43b via the FHCs 34a and 34b, respectively, and heats the cathodes 43a and 43b, respectively.

  The high voltage power supply unit 32 includes two high voltage power supplies 32a and 32b. When the low tube voltage in dual energy scan (the tube voltage indicates the potential difference between the anode and the cathode; the same applies hereinafter) is VL and the high tube voltage is VH, the high voltage power supply 32a can output the voltage VL. The high voltage power supply 32b can output the voltage VH-VL. For example, the low tube voltage in the dual energy scan is set to 80 [kV], and the high tube voltage is set to 140 [kV]. In FIG. 2, since the anode side is grounded, −80 [kV] is output to the cathode side by the high voltage power source 32a, and −140 [kV] is output to the high voltage power sources 32a and 32b.

  The negative output of the high voltage power supply 32a and the positive output of the high voltage power supply 32b are electrically connected, and this connection point serves as an output terminal and is connected to the cathode 43a of the X-ray tube 31. The positive output of the high voltage power supply 32a is electrically connected to the anode 42 of the X-ray tube 31 and grounded. The negative output of the high voltage power supply 32b is electrically connected to the cathode 43b of the X-ray tube 31. In the X-ray generator 21, −80 [kV] is output from the high voltage power supply 32 a to the cathode 43 a of the X-ray tube 31, and −140 [kV] is output from the high voltage power supplies 32 a and 32 b to the cathode 43 b. As described above, independent tube voltages can be applied to the anode 42 and the cathodes 43a and 43b between the anode 42 and the cathode 43a and between the anode 42 and the cathode 43b.

  Here, according to the X-ray generator 21, in the dual energy scan by the scanner device 11, the thermoelectrons are simultaneously emitted from the two filaments and applied with different voltages for each view under the control of the controller 26. it can. In that case, X-rays having a low energy distribution due to a low tube voltage and X-rays having a high energy distribution due to a high tube voltage can be simultaneously generated from the anode 42 for each view.

Also, according to the X-ray generator 21, under the control of the controller 26, the tube currents of the cathodes 43a and 43b are independently controlled via the two FHCs 34a and 34b, respectively. The tube current of the cathode 43b can be set separately. Therefore, as shown in FIG. 2, when the tube voltage (80 [kV]) of the cathode 43a is lower than the tube voltage (140 [kV]) of the cathode 43b, the FHC 34a controls the cathode 43b under the control of the controller 26. The tube current of the cathode 43a is set so as to be larger than the tube current.

  A control signal (switching signal) of the controller 26 for the X-ray generator 21 is also sent to the DAS 24. The image processing apparatus 12 visualizes differences in constituent elements of the subject O from data obtained by detecting X-rays transmitted through the subject O with different energies. For example, an image of a blood vessel formed by a calcified tissue and a contrast agent Can be separated to some extent.

  FIG. 3 is a schematic view of the X-ray tube of the present embodiment. FIG. 3 shows a configuration in the case where the anode 42 is disposed between the cathode 43a and the cathode 43b arranged to face each other. The anode 42 is grounded. Thermal electrons are emitted to the anode 42 from the cathodes 43a and 43b facing each other with the anode 42 interposed therebetween. As a result, the irradiated surfaces 42a and 42b of the anode 42 facing the cathodes 43a and 43b are irradiated with thermoelectrons, and X-rays in the range of the irradiation fields 130 and 140 are irradiated toward the subject.

  The irradiation surface 42a and the irradiation surface 42b are separated by a focal position difference D. The focal position difference D is a length between a position (focal position) where thermoelectrons are irradiated on the irradiation surface 42a and a position (focus position) where thermoelectrons are irradiated on the irradiation surface 42b. The focal position difference D is the width of the anode 42 in the electron flow direction.

  The irradiation surface 42a and the irradiation surface 42b are not perpendicular to the electron flow direction, but are inclined toward the X-ray irradiation direction. The inclination of the irradiation surface 42a and the irradiation surface 42b from the electron flow direction is the target angle. The X-ray generation positions of the irradiation surface 42a and the irradiation surface 42b are separated from each other by the focal position difference D. After the generation, the irradiation fields spread as shown in the figure. The generated X-rays are irradiated from the X-ray tube 31 through the diaphragm 23.

  X-rays are alternately generated from the irradiation surface 42a and the irradiation surface 42b and detected, and a dual energy scan can be realized by generating an image based on the detection result of each timing. It can also be simultaneous. By photographing at the same time or close timing, the accuracy can be improved when realizing dual energy subtraction.

  Since different tube voltages can be applied to the cathode 43a and the cathode 43b with respect to the anode 42, X-rays having different energies can be output from one X-ray tube 31 and detected. Since it is possible to simultaneously expose two different conditions of X-rays, a dual energy scan can be realized. Further, by simultaneously outputting two energy X-rays, dual energy subtraction is possible from the two energy X-rays detected by the detector 22.

  In the normal dual energy scan, instead of using a single anode and cathode, the tube voltage and tube current must be changed at the switching timing, which makes it difficult to control. These two X-ray conditions can be simultaneously exposed and detected by one X-ray tube 31. By realizing the dual energy scan with the above-described configuration, the difference in the constituent elements of the subject can be visualized, and for example, a calcified tissue part and a blood vessel image by a contrast agent can be sufficiently separated.

  Moreover, since the amount of X-rays to be irradiated can be kept constant even when the tube voltage is changed, it is possible to reduce the X-ray exposure of the subject while obtaining sufficient diagnostic image quality. With such a configuration, the X-ray CT apparatus can reduce the X-ray exposure of the subject while obtaining sufficient diagnostic image quality.

  FIG. 4 is a schematic view of an X-ray tube with different target angles. As in the case of FIG. 3, the anode 42 has an irradiation surface 42a on which thermal electrons are irradiated from the cathode 43a and an irradiation surface 42b on which thermal electrons are irradiated from the cathode 43b. The irradiation surface 42a has a first target angle with respect to the irradiation of thermoelectrons from the cathode 43a, and the irradiation surface 42b has a second target angle different from the first target angle with respect to the irradiation of thermoelectrons from the cathode 43b. Having a target angle.

  That is, the irradiation surface 42a side is set as the first target angle. The first target angle is formed by the surface orthogonal to the irradiation direction of the thermal electrons from the cathode 43a (the traveling direction of the thermal electrons) and the irradiation surface 42a. On the other hand, the irradiation surface 42b side is set as the second target angle. The second target angle is formed by a surface orthogonal to the irradiation direction of the thermal electrons from the cathode 43b (the traveling direction of the thermal electrons) and the irradiation surface 42b.

  The irradiation surface 42a and the irradiation surface 42b face the traveling direction of the accelerated thermoelectrons, respectively, but the directions are not completely coincident and are slightly inclined in the X-ray irradiation direction. This inclination is the target angle (angle). In other words, the target angle of each irradiation surface is equal to the angle formed between each irradiation surface and a surface perpendicular to the traveling direction of the thermoelectrons.

  Since the target angles are different, the maximum irradiation field and the maximum X-ray output are different and can be used properly even with the same focal size. Therefore, the X-ray irradiation field can be changed depending on the application. Further, by arranging detectors for detecting respective X-rays, X-rays under different conditions can be detected. Moreover, it is possible to simultaneously detect X-rays having different energies with higher accuracy by placing a grit.

  FIG. 5 is a detailed enlarged view of anodes having different target angles. Thermal electrons are incident on the irradiation surface 42a from the left side. X-rays are generated by incidence on the irradiation surface 42a, an irradiation field 130 is formed, and the subject is irradiated. On the other hand, thermoelectrons enter the irradiation surface 42b from the right side. X-rays are generated by incidence on the irradiation surface 42b, an irradiation field 140 is formed, and the subject is irradiated.

  Here, as shown in FIG. 5, the target angle formed on the irradiation surface 42a is θ1, and the target angle formed on the irradiation surface 42b is θ2. The first target angle θ1 and the second target angle θ2 are different. Specifically, the shape of the anode 42 can be set such that the first target angle θ1 is larger than the second target angle θ2, for example, θ1 = 12 degrees and θ2 = 8 degrees. The first target angle θ1 and the second target angle θ2 are each formed between 8 degrees and 12 degrees. Conversely, the anode 42 may have a shape in which the second target angle θ2 is larger than the first target angle θ1.

  With this asymmetric shape, one of the dual energy scans, for example, high-energy thermionic electrons are irradiated onto the target surface at a small target angle, while the other low-energy thermionic electrons are irradiated at a large target angle. The target surface can be irradiated.

  The portion where the electron current collides to generate X-rays is the focus of the target. The smaller the focus, the smaller the penumbra and the higher the resolution of the X-ray image. One factor that determines the focal spot size is the target angle. One side of the effective focal spot size can be reduced as the angle of the anode becomes smaller. On the other hand, as the focal point becomes smaller, the amount of heat generated per unit area of the focal point increases with the generation of X-rays, so that the focal point is required to have an appropriate size. By making one target angle small while making the other target angle relatively large, it is possible to balance the two.

  FIG. 6 is an explanatory diagram for explaining the outline of X-ray irradiation from the X-ray tube of the present embodiment. As described with reference to FIGS. 3 to 5, X-rays are irradiated from the irradiation surfaces 42 a and 42 b of the anode 42, respectively. X-rays in the range of the irradiation fields 130 and 140 are directed to the subject O, respectively. Then, the light passes through the subject O and is detected by the detector 22.

  As shown in the drawing, the X-ray irradiation field 130 irradiated from the irradiation surface with a large target angle is formed larger than the X-ray irradiation field 140 irradiated from the irradiation surface with a small target angle. The intensity of the X-rays of the large irradiation field 130 is reduced, and the intensity of the X-rays of the small irradiation field 140 is increased. While the X-ray output is stronger when the target angle is smaller, the irradiation field can be widened by increasing the target angle. In this case, since it becomes an asymmetrical shape, it is possible to use only one irradiation surface depending on the application without using it as a dual energy subtraction.

  The difference in target angle formed on the irradiation surfaces 42a and 42b is not relatively large as shown in FIG. However, as shown in FIG. 6, when the X-rays irradiated from both sides pass through the subject O and reach the detector 22, there is a difference in the difference in the spread of the irradiation field 130 and the irradiation field 140. Therefore, the X-ray ranges detected by the detector 22 are different. Thus, since both irradiation fields differ and X-ray intensity | strength also differs, each X-ray from irradiation surface 42a, 42b can be used for a different use.

  FIG. 7 is a schematic diagram for explaining the synchronous movement control of the X-ray irradiation position and the top board. The dual energy scan by the X-ray CT apparatus using a single anode has already been described. Here, in order to improve the accuracy of the dual energy scan, a configuration for synchronizing the irradiation timing and the irradiation position from the two irradiation surfaces will be described as follows. Here, a case where the target angles of the two irradiation surfaces are the same will be described.

  In the helical scan, as described with reference to FIG. 1, the rotating unit 25 and the top plate 29 are relatively moved along the body axis. The rotating unit 25 moves relatively to the left in FIG. 7, and the top plate 29 moves relatively to the right in FIG. The relative movement may be performed by either movement of the rotating unit 25 or the top plate 29, but in any case, the relative movement will be described. In order to show the order of movement, v and w are added as reference numerals, but both are the same except that the position has moved.

  First, the X-ray 210 is irradiated from the anode 42v to the subject O in which the rotating unit 25v and the top plate 29v are arranged at a certain timing during the helical scan. Then, an image is generated based on the irradiated X-ray 210. The relative movement is performed until the next timing, and the rotation unit 25w and the top plate 29w are arranged. With this arrangement, the subject O is irradiated with X-rays 220 from the anode 42w. Then, an image is generated based on the irradiated X-ray 220.

  Here, the X-ray 210 and the X-ray 220 are irradiated from the same anode 42, but are irradiated from different positions separated by the focal position difference D as described with reference to FIG. Referring to FIG. 3, X-rays 210 are irradiated from, for example, the irradiation surface 42a, and X-rays 220 are irradiated, for example, from the irradiation surface 42b. The positions of the irradiation surface 42a and the irradiation surface 42b differ by the focal position difference D as described with reference to FIG. Since the irradiation surface 42a and the irradiation surface 42b are at different positions, the position of the generated X-ray is also shifted accordingly. Therefore, the irradiation position of the subject O is shifted by this distance. In order to correct such a shift, the rotating unit 25 and the top plate 29 are relatively moved as described above. That is, the relative position between the rotating unit 25 and the top plate 29 is moved by this focal position difference D.

  The rotation driving device 28 or the top plate driving device 30 controls the relative movement speed, thereby setting the relative position by the focal position difference D between the irradiation timing of the X-ray 210 and the irradiation timing of the X-ray 220. Move. That is, if the time from the irradiation of the X-ray 210 to the irradiation of the X-ray 220 is t, the rotating unit 25 and the top plate 29 are relatively moved by the distance of the focal position difference D during this time t. That is, the relative movement speed is D / t. Thereby, the irradiation position of the X-ray 210 from the anode 42v and the irradiation position of the X-ray 220 from the anode 42w are relatively coincident. That is, the same position of the subject O is irradiated by the X-ray 210 and the X-ray 220.

  Thus, the rotating unit 25 controls the relative movement of the set of the X-ray tube 31 and the detector 22 with respect to the top plate in the direction of the rotation center axis in addition to the above-described rotation operation. By this relative movement, the rotating unit 25 determines the relative movement distance between the irradiation surface 42a and the irradiation surface 42b between the imaging intervals from the imaging with one cathode to the imaging with the other cathode during continuous imaging. Match the length.

  Since the above-described synchronization processing is processing for irradiating the same position with two X-rays, the same purpose may be achieved by different movement distances. For example, in the above-described example, relative movement during time t Although the distance is D, it may be D / 2. The relative movement distance is 0, D / 2, D, 3D / 2. Of these, the distance between the X-rays becomes D by combining the relative movement distances of 0 and D, and those of D / 2 and 3D / 2, respectively. Can be realized. The relative movement distance has been described from D to D / 2, but the same applies to D / 3, D / 4.

  As described above, the rotating unit 25 calculates an integer multiple of the relative movement distance between the imaging interval from imaging with one cathode to imaging with the other cathode during continuous imaging between the irradiation surface 42a and the irradiation surface 42b. It can also be matched to the length of

  In particular, in the case of dual energy subtraction, if there is a difference between the two images for which the difference is taken, the resolution will be lowered. However, by synchronizing as described above, the shooting positions can be matched more accurately. Thereby, the imaging resolution in the case of dual energy subtraction can be improved. That is, since the shift width of the focal position is relatively moved according to the pitch width of the helical scan, the resolution can be improved.

  Although several embodiments of the present invention have been described, these embodiments are presented by way of example and are not intended to limit the scope of the invention. These novel embodiments can be implemented in various other forms, and various omissions, replacements, and changes can be made without departing from the scope of the invention. These embodiments and modifications thereof are included in the scope and gist of the invention and are included in the equivalent scope described in the claims.

12 Image processing device 21 X-ray generator 24 DAS
31 X-ray tube 32 High voltage power supply units 32a, 32b, 32c, 32d High voltage power supply 34 FHC units 34a, 34b FHC
41 Insert valve 42 Anode 42a, 42b Irradiation surface 43 Cathode unit 43a, 43b Cathode 130, 140 Irradiation field

Claims (7)

A first cathode that is set to a first tube voltage and irradiates electrons;
A second cathode that is set to a second tube voltage different from the first tube voltage and irradiates electrons in a direction opposite to the first cathode;
A first irradiation surface disposed between the first cathode and the second cathode and irradiated with electrons from the first cathode; and a second irradiation surface irradiated with electrons from the second cathode. The first irradiation surface has a first target angle with respect to the electron irradiation from the first cathode, and the second irradiation surface has the second cathode. An anode for irradiating the subject with X-rays having a second target angle different from the first target angle with respect to irradiation of electrons from
An X-ray tube having
A detector for detecting the X-ray transmitted through the subject;
An X-ray CT apparatus comprising:   The first target angle is an angle formed by a surface orthogonal to an electron irradiation direction from the first cathode and the first irradiation surface, and the second target angle is the second target angle. 2. The X-ray CT apparatus according to claim 1, wherein the X-ray CT apparatus has an angle formed by a surface perpendicular to an irradiation direction of electrons from the cathode and the second irradiation surface.   The X-ray CT apparatus according to claim 1, wherein each of the first target angle and the second target angle is between 8 degrees and 12 degrees. A first cathode that is set to a first tube voltage and irradiates electrons;
A second cathode that is set to a second tube voltage different from the first tube voltage and irradiates electrons in a direction opposite to the first cathode;
A first irradiation surface on which electrons are irradiated from the first cathode and a second irradiation surface on which electrons are irradiated from the second cathode are back to back, and the respective irradiation surfaces are directed toward the subject. An anode that emits X-rays;
An X-ray tube having
A detector for detecting the X-ray transmitted through the subject;
The X-ray tube is held integrally with the detector facing the detector, rotated around the subject, and the relative movement of the set of the X-ray tube and the detector with respect to the top plate is controlled in the rotation center axis direction. Then, an integer multiple of the distance of the relative movement during the imaging interval from imaging with one cathode to imaging with the other cathode during continuous imaging is calculated between the first irradiation surface and the second irradiation surface. An X-ray CT apparatus comprising control means for matching the length of the X-ray CT.   The control means determines the distance of the relative movement between the imaging intervals from imaging with one cathode to imaging with the other cathode during continuous imaging, between the first irradiation surface and the second irradiation surface. The X-ray CT apparatus according to claim 4, wherein the X-ray CT apparatus matches the length of the X-ray CT.   The X-ray CT apparatus according to claim 1, further comprising a reconstructing unit that reconstructs an image from projection data based on X-rays detected by the detector. .   The X-ray CT apparatus according to claim 1, further comprising a power source that applies different independent tube voltages between each of the two cathodes and the anode.

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