JP2004154277A - Method for power feeding to multiple radiation source x-ray ct apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a method for power feeding for securely and stably feeding power, in a small-sized multiple radiation source X-ray CT apparatus serving both as a target and an anode. <P>SOLUTION: A combination of a cathode, a grid and an anode corresponding to a target part of a subject to be photographed is selected, and power is fed to the selected cathode to generate electronic beams. At this time, the negative limited potential of the selected grid is released, and a positive potential is applied to the specific anode selected as the target to be the equal potential in the circumferential direction via a plurality of feeding points. Then, by forming a bias potential from the cathode to the specific anode, radiation of the electronic beams passing the grid from the cathode to the specific anode is permitted. <P>COPYRIGHT: (C)2004,JPO

Description

[0001]
TECHNICAL FIELD OF THE INVENTION
The present invention relates to a power supply method for a multi-source X-ray CT apparatus used for three-dimensional image diagnosis.
[0002]
[Prior art]
The multi-source high-speed X-ray CT scanner employs an electron beam control method of electrically turning on / off the generation of X-rays, thereby greatly shortening the scan time of a conventional X-ray CT scanner (1/1). (60 to 1/2000 seconds) to perform tomography of the object to be measured. Such a high-speed X-ray CT scanner is proposed in, for example, JP-A-10-295682 and JP-A-10-075944.
[0003]
A conventional multi-source X-ray CT scanner includes a plurality of detectors fixedly arranged at equal pitch intervals on a concentric circle surrounding an imaging region, and a vacuum provided further outside to surround a group of these detectors. A vessel, a plurality of X-ray generators fixedly arranged in a vacuum vessel so as to be arranged at equal pitches on concentric circles surrounding a diagnostic space, and a control device for controlling the X-ray generator . The X-ray generator is composed of 32 triode vacuum tubes densely arranged on concentric circles, each of which irradiates a fan-shaped X-ray (fan beam) toward a subject placed in an imaging region. . The X-ray generation control device has 32 pulse generation control ports corresponding to the pulse generators provided for each X-ray generator in a one-to-one correspondence, and an X-ray generator optimal for imaging based on predetermined input data. Is selected, and ON / OFF control of the power supply circuit is performed at high speed so that X-rays are emitted only from the selected X-ray generator.
[0004]
X-rays emitted from the X-ray generator pass through the subject and are detected by the detector. The detection signal is integrated in a data recording device, further processed by a data processing device, and reproduced on a display as an X-ray tomographic image.
[0005]
[Problems to be solved by the invention]
However, in the conventional apparatus, since the installation space is limited by the spatial arrangement relationship with the subject such as the inner diameter of the cylinder and the dimensions of the X-ray generation unit, the number of X-ray generation units that can be arranged is limited. There is. For this reason, there is a problem that the spatial resolution of the apparatus cannot be improved by increasing the number of the X-ray generation units without limit.
[0006]
Further, in each X-ray generation unit of the conventional apparatus, a power supply circuit is required for each of the cathode, anode, and gate (grid), so that the power supply capacity becomes enormous. In particular, when trying to increase the spatial resolution, there is a problem that not only the manufacturing cost of the power supply circuit increases but also the operation cost increases because the total number of power supply circuits is enormous.
[0007]
Further, if a large number of X-ray generators are densely arranged over the entire circumference in order to subdivide an imaging area per shot, it becomes difficult to accurately guide an electron beam to a target X-ray target.
[0008]
SUMMARY OF THE INVENTION The present invention has been made to solve the above-described problems, and it is an object of the present invention to provide a power supply method capable of supplying power reliably and stably in a small multi-source X-ray CT apparatus serving as a target and an anode. Aim.
[0009]
[Means for Solving the Problems]
A power supply method for a multi-source X-ray CT apparatus according to the present invention is provided in an annular vacuum vessel so as to be electrically insulated from the vacuum vessel and surround an object arranged along an axis. An X-ray generator having a large number of cathodes / grids / anodes densely fixedly arranged over a 360-degree circumference along a circumference centered on the axis; A plurality of X-ray detectors that are densely fixedly arranged over a 360-degree circumference along a circumference centered on the axis so as to correspond to the detector, and a plurality of feeding points that conduct to the anode and a power supply; A method for feeding a multi-source X-ray CT apparatus comprising:
A combination of a cathode / grid / anode corresponding to the imaging region of the subject is selected and an electron beam is generated by supplying power to the selected cathode, while releasing the negative regulated potential of the selected grid and selecting as a target. Applying a positive potential to the specific anode via the plurality of feeding points so as to be equal in the circumferential direction with respect to the specific anode to form a bias potential from the cathode toward the specific anode. Accordingly, emission of an electron beam from the cathode toward the specific anode through the grid is allowed.
[0010]
Note that it is preferable to supply a direct current having a voltage higher than the potential of the cathode by 150 ± 5 kV to the anode through a plurality of feeding points.
[0011]
Further, it is preferable that the plurality of feeding points are distributed and arranged at four positions of azimuths 0 °, 90 °, 180 °, and 270 ° with respect to the anode.
[0012]
BEST MODE FOR CARRYING OUT THE INVENTION
Hereinafter, preferred embodiments of the present invention will be described with reference to the accompanying drawings.
[0013]
As shown in FIGS. 1 and 2, a medical X-ray CT apparatus 10 as a 150 kV multi-source X-ray CT apparatus includes a donut-shaped gantry 11 in which an X-ray generator 30 and a radiation detector 60 are built. The patient 5 as a subject is provided so as to be taken in and out of the diagnostic space 11a in the center of the vacuum chamber together with the moving bed 2. That is, the movable bed 2 is supported by the slider mechanism 3 so as to be movable in the X-axis direction along the guide rail 4.
[0014]
The gantry 11 is provided with an X-ray generator 30, a beam limiter (not shown), a radiation detector 60, an image signal digitizer (not shown), an electron gun driving circuit (not shown), and the like. The X-ray generator 30 is housed in an annular or tubular vacuum container 20. The inside of the vacuum vessel 20 is evacuated by a vacuum pump through an exhaust port (not shown). The fan-shaped X-rays 6 emitted from the X-ray generator 30 are narrowed down by a collimator (not shown), further defined by an unillustrated beam limiter to a predetermined diameter at an irradiation position, and placed in the diagnostic space 11a. After being transmitted by the radiation detector 60.
[0015]
The radiation detector 60 is densely fixed and arranged on a concentric circle surrounding the diagnostic space 11a in which the subject 5 is arranged, and includes 4086 ultra-sensitive CdTe single-crystal photoelectric conversion elements 72. It has a resolution of 5 mm. Incidentally, the imaging width of one shot is about 2 mm. The gantry 11 has an outer diameter of about 2000 mm and an inner diameter of about 800 to 1000 mm.
[0016]
As shown in FIG. 2, the X-ray generators 30 are arranged on a concentric circle outside the circumference on which the plurality of radiation detectors 60 are arranged, and are provided corresponding to the radiation detectors 60. The X-ray generator 30 and the radiation detector 60 are slightly shifted in the X-axis direction, and the X-ray 6 is inclined slightly forward with respect to the radius (Z-axis) of the gantry 11 as shown in FIG. In the direction of the light. Therefore, as shown in FIGS. 1 to 3, the X-rays 6 pass through the subject 5 placed in the diagnostic space 11 a without being blocked by the radiation detector 60 on the X-ray emission side (upper side). And is detected by the radiation detector 60 on the opposite side (downward).
[0017]
A data recording device 18 is connected to an input side of a control device 17 having a digital operation circuit. The X-ray transmission information detected by the radiation detector 60 is photoelectrically converted into a current signal proportional to the transmitted X-ray amount, and transmitted to a data recording device 18 and an image signal digitizer (not shown) via a preamplifier 15 and a main amplifier 16. Sent and recorded.
[0018]
The recorded data is further output from the data recording device 18 to the data processing device 19 and processed by the data processing device 19. The processed data is reproduced and displayed on a display (not shown) as X-ray CT image information of the subject 5.
[0019]
The output side of the control device 17 is connected to the power source 14, the anode (target) 33 in the X-ray generator 30, the discharge electrode 44, and the grid electrode 45 of the gate array. When an X-ray generation command signal is sent from the data recording device 18 to the control device 17, the control device 17 controls the power supply operation from the power supply 14 to the electron gun driving circuit based on the command, and controls a plurality of grid electrodes. An object at a position suitable for a part to be imaged is selected from among 45. In response to this, an electron beam is emitted from one of the discharge electrodes 44 in the X-ray generator 30, the negative bias voltage applied to the selected grid electrode 45 is released, and the potential becomes zero, and the electron beam 6 a is changed to the grid electrode. The light passes through the hole 45 and enters the target 33. When the electron beam 6a is incident on the target 33, X-ray 6 is generated from the target 33, and the fan-shaped X-ray 6 is emitted toward the subject 5 via a collimator (not shown) attached to a window. Has become.
[0020]
As shown in FIG. 9, a pipe 81 of the exhaust mechanism 80 is introduced into the gantry 11 from the gantry side wall opening 83 and communicates with the exhaust port 82 of the vacuum tube 20. The pipe 81 communicates with a large-diameter pipe 84 via a valve 85, and the large-diameter pipe 84 communicates with a suction port of a vacuum pump (not shown). This pump (not shown) has the ability to exhaust the inside of the vacuum tube 20 until the internal pressure of the tube 20 becomes 1 × 10 ⁻⁷ to 1 × 10 ⁻⁹ Torr.
[0021]
The terminal 27 penetrates through the side wall of the gantry 11 and is introduced inside, and is electrically connected to each of the electron gun driving circuit, the gate array (grid) driving circuit, and the anode (target) driving circuit of the X-ray generator 30. This terminal 27 is connected to an external power supply 14 whose operation is controlled by the control device 17.
[0022]
A plurality of Cu electrode rods 28 are electrically connected to the terminals 27, and the tips of the Cu electrode rods 28 are pressed against the power supply points 34 a of the anode block 34.
[0023]
The terminal 27 and the Cu electrode rod 28 are insulated from surrounding members by the insulator 25. The insulator 25 is made of high-purity alumina (Al ₂ O ₃ ) having high withstand voltage characteristics, and has a withstand voltage of 150 to 200 kV.
[0024]
As shown in FIG. 8, the power supply points 34a are distributed to four positions of the anode block 34 at directions of 0 °, 90 °, 180 °, and 270 °. A high-voltage DC of 150 kV is supplied from the power supply 14 between the anode (target) 33 and the cathode 44 via these four power supply points 34a. By employing such a multi-point power supply system, the X-rays 6 having substantially the same output are emitted from each of the X-ray generators 30 at substantially the same timing.
[0025]
Next, the X-ray generator 30 will be described with reference to FIGS.
[0026]
As shown in FIG. 3, the X-ray generator 30 is housed in a vacuum vessel 20 whose entire surface is substantially covered with a shielding member 24 made of a lead plate having a thickness t4 of 3 to 5 mm. The vacuum vessel 20 is formed of an annular hollow tube made of non-magnetic stainless steel having a thickness of 3 mm. The shielding member 24 covers most of the outer surface of the vacuum vessel 20 except for the window 20a serving as an X-ray emission port, so that the X-rays 6 do not leak to directions other than the diagnostic space 11a.
[0027]
An X-ray emitting unit 50 is attached to the window 20a of the vacuum container. The X-ray emitting section 50 is also provided with a shielding member 53 made of a lead plate and a collimator (not shown). In the present embodiment, the thickness of the shielding member 53 of the X-ray emission unit 50 is set to 5 mm.
[0028]
A high-voltage introduction terminal 27 penetrates through the side wall of the gantry 11 and is introduced inside, and is electrically connected to the electron gun driving circuit, the gate array (grid) driving circuit, and the anode (target) driving circuit of the X-ray generator 30. . The high-voltage introduction terminal 27 is connected to the power supply 14 whose operation is controlled by the control device 17, and supplies a direct current of 150 kV to the Cu electrode rod 28 shown in FIG.
[0029]
The Cu electrode rod 28 is insulated and protected from surrounding members by a target current introducing insulator 25 made of ceramic having high withstand voltage characteristics. The tip of the Cu electrode rod 28 is pressed against a feed point 34 a of an anode block 34 of the X-ray generator 30. The insulator 25 is made of a ceramic such as alumina, and has a pressure resistance of 200 kV.
[0030]
The X-ray generator 30 will be described in more detail with reference to FIGS. In these figures, the shapes of the vacuum vessel and its accessories are schematically shown.
[0031]
The X-ray generator 30 emits an electron beam 6 a from an electron gun 40 into a tungsten target (anode) 33 in an anode housing 31 in a vacuum atmosphere, and generates a secondary X-ray 6 from the target 33. An opening of the vacuum vessel 20, for example, the window 20a is covered with a window member 51 of the X-ray emission unit 50 via an O-ring, and the inside is kept airtight. An exhaust pipe (not shown) communicates with the vacuum vessel 20, and the inside of the vacuum vessel 20 is evacuated to a predetermined degree of vacuum (for example, 1 × 10 ⁻⁶ to 1 × 10 ⁻⁷ Torr) by a vacuum pump (not shown). I have.
[0032]
The cathode (electron gun) 40 is supported by an insulating support plate 41 while being insulated from peripheral members in the vacuum vessel 20. The cathode 40 is provided with a discharge electrode 44 and the grid electrode 45 made of CeB ₆ single crystal or SEB ₆ single crystal or LaB ₆ single crystal. A negative bias voltage (for example, -700 V) is always applied to the grid electrode 45. When this voltage is set to zero potential, the cathode electrode 44 passes through the beam passage hole 45b of the grid electrode 45 to the anode 33. The electron beam 6a is emitted toward the device. The diameter of the electron beam passage hole 45b of the grid electrode is 5 mm.
[0033]
The grid electrode 45 is made of tungsten, molybdenum, or the like on a ceramic ring substrate of silicon nitride (Si ₃ N ₄ ), silicon oxide (SiO ₂ ), silicon carbide (SiC), alumina (Al ₂ O ₃ ), sialon (SiAlON), or the like. After laminating a conductive thin film made of a high melting point metal or alloy such as tantalum, an insulating portion is formed by etching the conductive thin film into a predetermined pattern. The electron beam passage holes 45b of the grid electrode 45 are formed by machining such as drilling after pattern etching.
[0034]
As shown in FIG. 4, a pair of positive and negative filament current introduction terminals that are insulated and protected from the surroundings by insulators 42a and 42c are connected to a discharge electrode 44 via a cable 42b, and a discharge voltage of, for example, +150 kV is applied to the discharge electrode 44. Is applied.
[0035]
The discharge electrode 44 is of a so-called Cambridge type, and is made of a single crystal of a non-metallic compound of SeB ₆ or CeB ₆ or LaB ₆ having a diameter of 1 to 3 mm. Since the high-voltage electric field concentrates on the tip of the discharge electrode 44 to easily cause chipping, the corner edge portion is cut off and rounded. The lower half of the discharge electrode 44 is cut off from both sides to form flat surfaces, and a pair of positive and negative molybdenum wires are connected to each flat surface so as to press a graphite chip. The base ends of the pair of positive and negative molybdenum wires are respectively connected to positive and negative terminals embedded in the insulator, and a high voltage is applied from a power supply (not shown) via a cable 42b.
[0036]
The target 33 functions as an anode electrode for generating X-rays, and is disposed at a position where a surface on which the electron beam 6a emitted from the cathode (electron gun) 40 collides becomes a focus. This is a so-called braking X-ray in which a part of the kinetic energy is emitted as X-rays 6 when high-speed electrons enter the target atom and are blocked from moving. That is, when the electron beam 6a is incident on the target 33 from the discharge electrode 44, the X-ray 6 is emitted from the target 33 in a fan-like direction in a direction of reflection according to the angle of the surface. The target 33 is preferably formed of a rectangular plate of tungsten or a tungsten alloy, and has a thickness of 0.5 to 7.0 mm, a width of 8 to 12 mm, and a length of 30 to 50 mm. In this embodiment, the thickness of the target 33 is set to 1.0 mm. The irradiation surface of the target 33 is inclined by about 20 ° with respect to the optical axis of the electron beam 6a.
[0037]
Theoretically, secondary X-rays are generated by hitting an electron beam on a target having a thickness of several microns, but if the thickness of the target is too thin, mortality, cracking, chipping, etc. It is most preferable that the thickness of the target is about 1.0 mm because the target is easily damaged. Note that if the thickness of the target 33 is too large, the amount of heat generation increases and cooling becomes difficult. Therefore, it is preferable to make the target 33 as thin as possible.
[0038]
The cooling block 32 is made of a good conductor having excellent thermal conductivity such as copper or aluminum, holds the target 33 on its inclined surface, and protects the target 33 from thermal damage. The cooling block 32 is supported by an insulating support member 25 for insulation from the vacuum vessel.
[0039]
The coolant supply pipe 26 penetrates through the lid 23 and is introduced therein, and communicates with the internal flow path 32 a of the cooling block 32 in the anode housing 31. The refrigerant supply pipe 26 communicates with an oil supply source (not shown) so that cooling oil as a refrigerant is circulated and supplied to the internal flow path 32a of the block. Although the target 33 is heated to a high temperature by the injection of the high energy electron beam 6a, it is not consumed in a short time because it is cooled from the back side by the cooling block 32. The discharge electrode 44 is also cooled by a cooling mechanism (not shown).
[0040]
An outline of the gate array control circuit will be described.
[0041]
The anode 33, the discharge electrode 44, and the grid electrode 45 of the gate array in the X-ray generator 30 are connected to n pulse generation control ports via n pulse generators built in the control device 17, respectively. I have. When a setting mode signal is input from a mode setting indicator (not shown), the CPU of the control device 17 sends an X-ray generation command signal to the pulse generation control port according to the setting mode, and responds to the pulse generation control port receiving the command signal. A signal is transmitted to the corresponding pulse generator, and a negative bias voltage is released to the corresponding grid electrode 45 to make it a zero potential. As a result, the electron beam 6 a passes through only the beam passage hole of the corresponding grid electrode 45 and enters the anode 33, and the X-ray 6 is emitted from the anode 33.
[0042]
Next, control of the anode potential will be described.
[0043]
When the electron beam 6a is incident on the anode 33 as a target from the cathode 40 (not shown), the power source is set so that an equipotential line 90a shown in FIG. 14 is controlled. That is, the control device 17 selects a cathode / grid / anode combination suitable for photographing, supplies power to the selected cathode 40 to generate an electron beam, and releases the negative regulated potential of the selected grid 45. At the same time, a positive potential is applied to the specific anode 33 selected as the target via the plurality of feeding points so as to have the same potential in the circumferential direction through a plurality of feeding points, and a bias from the cathode 40 toward the specific anode 33 is applied. The formation of the potential allows the emission of electron beams from the cathode 40 to the specific anode 33 through the grid 45. As a result, as shown in the figure, the power supply control to the anode 33 having the same potential in the circumferential direction is performed.
[0044]
On the other hand, in a standby state in which no X-rays are generated, the control device 17 maintains the regulated potential of the grid 45 as it is, so that an equipotential line 90b shown in FIG. 10B is formed.
[0045]
Next, the X-ray emission unit 50 will be described.
[0046]
As shown in FIG. 7, the X-ray emitting unit 50 is attached to the outside of the window 20 a of the annular vacuum vessel 20 over the entire 360 °. The window member 51 of the X-ray emitting unit 50 is made of a material that easily transmits X-rays and has a small X-ray attenuation rate, such as aluminum, beryllium, an alloy thereof, or stainless steel. A notch groove 52 is provided in the X-ray passing portion of the window member 51, thereby forming a thin portion 51a. The notch groove 52 is formed by cutting approximately half the thickness of the window member 51 using a milling machine or the like over the entire 360 ° circumference. Although the thickness t2 of the thin portion 51a varies depending on the material of the window member 51, it is necessary to ensure a pressure resistance that can withstand at least the negative pressure of the vacuum vessel 20. For example, when the window member 51 is made of an aluminum plate having a thickness t1 (= 5 mm), the thickness t2 of the thin portion 51a needs to be at least 2.5 mm.
[0047]
The X-ray attenuation rate by the window member 51 is desirably suppressed to 10% or less, and most desirably 5%.
[0048]
Further, the window member 51 is mostly covered with a shielding member 53 made of a lead plate. The shielding member 53 has a beam passage 53a that allows the X-ray 6 to pass therethrough. In this embodiment, the thickness t3 of the shielding member 53 is 5 mm, and the diameter d2 of the beam passage 53a is 0.5 mm to 3.0 mm. The X-ray 6 passes through only the beam path 53a of the shielding member 53, is narrowed by a collimator (not shown), and is emitted. The emitted X-rays 6 are detected by the radiation detector 60 on the opposite side after passing through the subject 5 in the diagnostic space 11a.
[0049]
As shown in FIG. 7, in this embodiment, the diameter W1 of the window 20a of the vacuum vessel is set to 20 to 30 mm, and the width of the thin portion 51a of the window member is made substantially equal to the diameter W1 of the window 20a. The beam tilt angle θ1 between the optical axis of the X-ray 6 and the Z axis 59 (vertical axis) was set to 0.1 ° to 2.5 °. The mounting angle θ2 between the window member 51 and the Z axis 59 (vertical axis) was set to 95 ° to 105 °.
[0050]
In the shielding member 53 of the X-ray emission part, the cross-sectional shape of the X-ray emission hole 53a may be a circle, an ellipse, an ellipse, or a slit. When the exit hole is formed in a slit shape, it is preferable to select a portion through which the X-ray beam does not pass and provide a spacer in order to keep the interval between the slits constant.
[0051]
Next, the radiation detector 60 will be described.
[0052]
As shown in FIGS. 7 and 3, the radiation detector 60 is supported on the inner peripheral surface of the vacuum container 20 via a ring frame 62. The radiation detectors 60 are provided in the same number (for example, 4086) as the number of the X-ray detectors 30 in one-to-one correspondence with the X-ray detectors 30. The radiation detector 60 and the X-ray generator 30 are arranged with a slight shift in the X-axis direction, and the X-ray 6 is slightly inclined forward with respect to the radius (Z-axis) of the gantry 11 as shown in FIG. In the direction of the light. Therefore, the X-rays 6 pass through the subject 5 placed in the diagnostic space 11a without being blocked by the radiation detectors 60 on the X-ray emission side (upper side) and are on the opposite side (lower side). 60.
[0053]
The housing 61 of the radiation detector 60 is fastened to the inner peripheral surface of the ring frame 62 with bolts or the like, and houses a sensor assembly (detection unit) 70 having a CdTe photoelectric conversion element 72 inside. The shielding member 63 is attached to the inner peripheral surface of the housing 61. An entrance 63 a is formed in the shielding member 63, and the X-ray 6 is detected by the CdTe photoelectric conversion element 72 through the entrance 63 a and the opening 61 a of the housing.
[0054]
The sensor assembly (detection unit) 70 includes a CdTe photoelectric conversion element 72 and a printed circuit board 75. The CdTe photoelectric conversion element 72 is made of a single crystal of cadmium telluride having a columnar rectangular parallelepiped with a square cross section.
[0055]
The 4086 CdTe photoelectric conversion elements 72 constituting the sensor array are arranged at equal pitch intervals on the printed circuit board 75 such that the light receiving surfaces are aligned at the same height level. The printed circuit board 75 is supported by the ring frame 62 via an insulating support member 64 as shown in FIG. 7, and is connected to the other end face (opposite to the light receiving face) of the CdTe photoelectric conversion element 72 by gold wire bonding (not shown). And sealed with resin.
[0056]
Next, the cathode 40 will be described.
[0057]
As shown in FIGS. 4 and 5, the cathode 40 includes a discharge electrode 44 and a grid electrode 45 that are insulated from surrounding members. The discharge electrodes 44 are supported by the frame 47 in a state insulated from the surroundings by an insulating substrate 41 made of ceramic. On the other hand, the grid electrode 45 is supported by the frame 47 while being insulated from the surroundings by an insulating ring 48 made of ceramic. In the drawings, reference numerals 46a and 46b indicate bolts, and reference numeral 49 indicates a holding plate. The pressing plate 49 presses the peripheral portion of the insulating substrate 41 into the recess of the frame 47 so that the assembly including the discharge electrode 44 does not fall off the frame 47.
[0058]
The discharge electrode 44 is of a so-called Cambridge type, and is made of a single crystal of a non-metallic compound of SeB ₆ or LaB ₆ having a diameter of 1 to 3 mm. The lower half of the discharge electrode 44 is cut off from both sides to form flat surfaces, and a pair of positive and negative molybdenum wires are connected to each flat surface so as to press a graphite chip. The base ends of the pair of positive and negative molybdenum wires are respectively connected to positive and negative terminals embedded in the insulator, and a discharge voltage of, for example, plus 150 kV is applied from a power supply (not shown) via a cable 42b.
[0059]
The grid electrode 45 is connected to a DC power supply via a cable (not shown). A gate voltage of minus 700 V is applied to the grid electrode 45 from this power supply. When the electron beam 6a is emitted from the discharge electrode 44 toward the anode 33, the bias applied to the grid electrode 45 is released to make the grid electrode 45 zero potential. Thus, the electron beam 6a is emitted from the discharge electrode 44 to the anode 33 through the hole 45b of the grid electrode, and the X-ray 6 is generated at the anode 33.
[0060]
The grid electrodes 45 are arranged at equal pitch intervals on a circumference with a radius of 800 mm centered on the axis of the CT apparatus. When the number of X-ray generators 30 (targets 33) is 360, the grid electrode 45 has a pitch interval of 14 mm, a width of 12 mm, and a length of 100 mm. When the number of X-ray generators 30 (targets 33) is 240, the grid electrode 45 has a pitch interval of 20.9 mm, a width of 19 mm, and a length of 100 mm. The beam passage hole 45b is formed at the longitudinal center of the grid electrode 45. The diameter of the counterbore 45a is equal to the width of the grid electrode 45, and the diameter of the beam passage hole 45b is in the range of one third to one half (30 to 50%) of the diameter of the counterbore 45a. desirable.
[0061]
Next, an outline of the operation of the above device will be described.
When the main switch of the device 10 is turned on, the cathode electrode 44 has a voltage enough to heat the cathode electrode 44 to a temperature sufficient to emit a significant amount of electrons, minus 150 kV (this can be -20 V). , A bias voltage of, for example, −50 kV (which can be −1 kV) can be applied to the grid electrode 45, and a bias voltage of +150 kV can be applied to the anode electrode 33. At the predetermined time, the data recording device 18 outputs a first X-ray generation command signal to the control device 17.
[0062]
Based on this first command, a first X-ray generation command is input to a pulse generation control port in the X-ray generation controller 17. Upon receiving this input signal, the two pulse generators simultaneously generate pulse waves, and these pulse signals are applied to the grid electrode 45 corresponding to the pulse generator. In response to the pulse generation signal, the negative bias voltage of the grid electrode 45 is released to zero potential, and the electron beam 6 a passes through the hole of the grid electrode 45 and enters the anode electrode 33.
[0063]
At the predetermined time, the command to generate the first X-ray is finished, the negative bias voltage of the grid electrode 45 is restored, and the electron beam 6a is cut off again by the grid electrode 45. During this time, the X-ray 6 is emitted from the anode electrode 33. In this manner, the operation of sequentially switching the emission positions of the X-rays 6 is repeated.
[0064]
Incidentally, the gate switching time is controlled in the range of 2.1 to 20.8 microseconds (1 / 480,000 seconds to 1 / 480,000 seconds). Further, the X-ray generation time is controlled in a range of 1.4 to 13.9 microseconds (1 / 720,000 to 1 / 72,000 second), which is two-thirds of the gate switching time.
[0065]
The generated X-ray 6 is irradiated toward the subject 5 in the diagnostic space 11a. The irradiated X-ray 6 is absorbed in accordance with the transmittance of the subject 5 and detected by the radiation detector 60 facing the X-ray.
[0066]
The X-ray transmission information detected by the radiation detector 60 is converted into a current signal proportional to the transmitted X-ray dose, then amplified by the preamplifier 15 and the main amplifier 16, and sent to the data recording device 18 as a voltage signal.
[0067]
When the detection operation by the radiation detector 60 is completed, a second X-ray generation command is input to the pulse generation control port in the control device 17, and the same detection operation as described above is performed. Then, when the X-ray transmission information obtained from all the detection operations performed sequentially is detected by the radiation detector 60, it is converted into a current signal in proportion to the transmitted X-ray dose, and is converted into a preamplifier 15, a main amplifier 16, a data recording device. The signal processing is performed in the data processing device 19 via. From the signal-processed data, X-ray CT image information of the subject 5 can be obtained.
[0068]
【The invention's effect】
As described above, according to the present invention, it is possible to reliably and stably supply power to a small multi-source X-ray CT apparatus serving both as a target and an anode.
[Brief description of the drawings]
FIG. 1 is a configuration block diagram showing an overall outline of a multi-source X-ray CT apparatus.
FIG. 2 is a schematic configuration diagram showing a multi-source X-ray CT apparatus as viewed from an X-axis direction.
FIG. 3 is an internal perspective sectional view showing a multi-source X-ray CT apparatus according to an embodiment of the present invention.
FIG. 4 is a sectional view of a main part of a multi-source X-ray CT apparatus according to an embodiment of the present invention.
FIG. 5 is a cross-sectional view of a main part of the multi-source X-ray CT apparatus shown in the direction of arrow VV in FIG. 4;
FIG. 6 is an enlarged sectional view showing an X-ray generation unit.
FIG. 7 is an enlarged sectional view showing an X-ray emitting unit and a detecting unit.
FIG. 8 is a schematic front view showing a power supply point and a mounting position of a vacuum exhaust port.
FIG. 9 is an internal perspective cross-sectional view showing an exhaust passage communicating with a vacuum tube.
10A is an equipotential diagram of the anode at the time of X-ray emission, and FIG. 10B is an equipotential diagram of the anode in a standby state.
FIG. 11 is an internal perspective sectional view showing a multi-source X-ray CT apparatus according to another embodiment of the present invention.
[Explanation of symbols]
5 ... subject,
6 ... X-ray,
6a ... electron beam,
10 X-ray diagnostic imaging device,
11: vacuum chamber, 11a: diagnostic space,
14. Power supply,
15 detection circuit (preamplifier unit)
15a: operational amplifier, 15b: buffer amplifier,
15c: capacitor, 15d: reset switch,
16: Main amplifier unit
17 ... Control device,
18. Data recording device,
19 ... data processing device,
20: vacuum vessel (annular tube),
20a ... window (radiation exit),
21 ... chamber wall,
22 ... insulating support material,
23 ... lid,
24 ... Shielding material (lead plate),
25 ... Target current introduction insulator (ceramic),
26 ... refrigerant supply pipe,
27 ... High voltage introduction terminal,
28 ... electrode rod (Cu),
29 ... exhaust pipe,
30 X-ray generator,
31 ... Anode housing (Al or Cu)
32: cooling block, 32a: refrigerant channel (internal channel),
33 ... Anode (target, anode electrode)
34 ... Anode block,
34a ... feed point,
35 ... refrigerant channel,
36 ... refrigerant supply path,
37 ... irradiation window,
38 ... holding part,
40 ... cathode (electron gun),
42 ... filament current introduction insulator,
44 ... discharge electrode (cathode electrode; CeB ₆ , LaB ₆ ),
45 ... Grid electrode,
50 ... X-ray emission unit,
51 ... window member (Al),
53 ... Shielding material (lead plate),
60 ... radiation detector,
61 ... housing, 62 ... ring frame,
63: shielding plate (lead plate), 63a: entrance, 64: insulating support,
70: Sensor assembly (detection unit)
80 ... Vacuum exhaust unit,
81, 84 ... exhaust passage,
82 ... exhaust port,
83 ... opening,
90a, 90b ... equipotential lines.

Claims (3)

An annular vacuum vessel is provided in a state of being electrically insulated from the vacuum vessel, and surrounds a subject arranged along the axis 360 degrees around a circumference centered on the axis so as to surround the subject. An X-ray generator having a large number of cathodes / grids / anodes arranged in a densely fixed manner over an object, and a circumference around the axis center corresponding to the X-ray generator with a subject interposed therebetween. A plurality of X-ray detectors arranged densely and fixedly over an entire circumference of 360 °, and a plurality of feeding points electrically connected to the anode and a power source.
Select a cathode / grid / anode combination corresponding to the imaging region of the subject,
While powering the selected cathode to produce an electron beam,
Along with releasing the negative regulated potential of the selected grid, a positive potential is applied to the specific anode selected as the target so that the potential becomes equal in the circumferential direction through the plurality of feeding points to the specific anode. Forming a bias potential from the cathode toward the specific anode to allow the electron beam to pass from the cathode through the grid and toward the specific anode. Power supply method for X-ray CT apparatus. The method according to claim 1, wherein a direct current having a voltage higher than a potential of the cathode by 150 ± 5 kV is supplied to the anode through the plurality of feeding points. 3. The device according to claim 1, wherein the plurality of power supply points are arranged at four positions of directions of 0 °, 90 °, 180 °, and 270 ° with respect to the anode. 4. Method.

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