JP2005270562A - Radiation computer tomograph, radiation computer tomographic system and radiation computer tomographic method using the same - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide an on-vehicle type high-performance radiation computer tomograph for a mass examination at a low cost by providing a low-cost, high-speed and high-sensitivity subject-rotating type cone beam radiation computer tomograph. <P>SOLUTION: The subject-rotating type cone beam radiation computer tomograph detects a rotating angle of the subject and controls a signal accumulation time of a one-shot exposure type radiation image sensor panel based on the rotating angle signal. This tomograph further controls the signal accumulation time of a signal accumulation type reference signal generating means. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to a radiation CT imaging apparatus that performs tomographic imaging of a subject, a radiation CT imaging system, and a radiation CT imaging method using the same, and more particularly to a cone beam radiation CT imaging apparatus that uses a two-dimensional radiation image sensor panel. It is. Furthermore, the present invention relates to a subject rotation type cone beam radiation CT imaging apparatus for imaging by rotating between an opposing radiation source and a radiation image sensor while a subject such as a living body is standing on a floor surface or in a sitting position.

  In recent years, digitization has progressed in various fields of medicine. Also in the field of X-ray diagnosis, in order to digitize an image, a two-dimensional X-ray imaging apparatus that converts incident X-rays into visible light by a scintillator (phosphor) and further captures the visible light image with an image sensor. It has been developed.

  In the field of X-ray still images, as a two-dimensional X-ray imaging apparatus, for example, a large plate using 17-inch size (43 cm × 43 cm) amorphous silicon (a-Si) is used for mammography and chest imaging. Still image pickup devices (flat panel detectors) have been made.

  In the field of X-ray animation, incident X-rays are converted into scintillators (phosphors) and I.D. I. A two-dimensional digital X-ray fluoroscopic apparatus that converts visible light with a TV camera using a CCD type image pickup device and a system that replaces the TV camera with a flat panel detector. Has also been devised.

  Furthermore, application of a flat panel detector to an X-ray CT imaging apparatus is also considered. Japanese Laid-Open Patent Publication No. 2001-194461 discloses a flat X-ray CT imaging apparatus that captures images while a pair of an X-ray tube and a detector are rotated while facing each other around a subject lying on a bed. A technique using a panel detector is disclosed (Patent Document 1).

  In contrast, a turntable is provided that can rotate a subject in an upright or sitting position for lung cancer screening for a normal healthy person. I. There has been proposed an apparatus that performs imaging while rotating with respect to a two-dimensional X-ray imaging apparatus such as a TV camera or a flat panel detector. This is disclosed in JP-A-5-42132 (Patent Document 2), JP-A 2000-217810 (Patent Document 3), and the like. In FIG. 12, reference numeral 1 denotes an X-ray CT apparatus as an example. In the figure, 2 is a scanner unit, 5 is a turntable on which a subject P is placed, 6 is an X-ray source, and 7 is I · I. A processing circuit 3 includes a preprocessing unit 8 and an image reconstruction unit 9. Reference numeral 4 denotes display means.

  With this apparatus, it is possible to start shooting quickly while keeping the subject standing, and the replacement with the next subject can be completed in a short time, and the overall shooting time can be shortened. Furthermore, since imaging is performed at most one rotation using cone beam X-rays, it is suitable as a system suitable for examination such as reduction of exposure of a subject. In addition, since a gantry for rotating the X-ray source and the X-ray detector is not necessary, the positions of the X-ray source and the X-ray detector can be freely set, and the magnification rate and measurement site during measurement can be changed. be able to. Further, an X-ray source for general imaging can be used, and a gantry is not required, so that the cost of the system can be reduced. Furthermore, since the system configuration is simple, it is convenient for in-vehicle use.

As a feature of the cone beam X-ray CT imaging apparatus described in the present invention, the detector only rotates relative to the subject on the rotation axis, and does not move in the direction of the rotation axis. Therefore, if there is a large two-dimensional detector that can capture the entire region of interest of the subject, it is possible to acquire all projection image data for reconstructing all CT images with at most one rotation.
JP 2001-194461 A Japanese Patent Laid-Open No. 5-42132 JP 2000-217810 A

  When a conventional subject rotation type cone beam X-ray CT imaging apparatus is used for medical purposes, it is desired to shorten the restraint time of the subject as a subject as much as possible. As a method for shortening the holding period of the same posture of the subject, that is, the time required for the measurement, it is conceivable to increase the rotational speed of the subject, but the time required for one rotation of the subject is about 5 seconds or more, When the subject is rotated at a higher speed than this, there is a problem that the subject turns into a so-called eye that feels abnormal.

  In addition, when the subject is rotated at a high speed, the centrifugal force applied to the subject also increases as the rotational speed increases, and there is a problem that the subject moves during measurement.

  Depending on the part to be imaged, it is necessary to hold the breath while taking a picture. In particular, when taking a picture of a part with a large body movement due to breathing such as chest photography, the breath must be held. Since the breath-holding time of about 5 seconds per rotation is relatively long, the subject feels uneasy during the examination, and body movement artifacts may occur due to unexpected movements and tremors. There was a problem that shooting might fail because it was out of view.

  If the subject moves during the measurement, there is a problem that “blur” or “blur” occurs in the three-dimensional X-ray distribution image generated by the reconstruction, and the image quality of the three-dimensional X-ray image is deteriorated. It was.

  Also, when rotating a turntable on which a subject having a considerable weight is placed, it takes time to reach a certain rotational speed and stabilize. If it does so, the rotation of one rotation or more was needed, and there existed a problem that the burden of a patient became large.

  On the other hand, if you start shooting without being stable, the rotation angle for each projection will be different, and the accuracy of reconstruction will decrease, and the signal from the reference element will vary widely, leading to poor correction accuracy and image deterioration. There was a problem.

  In general, X-ray CT imaging apparatuses have a problem that even if X-rays are exposed under the same conditions of tube voltage and tube current, the X-ray intensity actually changes for each scan at the time of imaging, or more precisely for each view. It was. Causes such as expansion / contraction of the axis of the anode due to the temperature change of the X-ray tube, movement of the axis between the bearings by moving the tube, and change of the focal position due to swinging motion while the anode is rotating, etc. Problems occur. As a countermeasure against this problem, in the conventional single-slice type and multi-slice type X-ray CT imaging apparatuses, the intensity of X-rays that do not pass through the subject among the X-rays from the X-ray tube is measured at the end of the X-ray detector. In many cases, a method is used in which a reference channel is provided and each channel value is corrected using a value measured in the reference channel to obtain data for creating an image. The line CT imaging apparatus has the same problem.

  Further, in order to rotate the turntable on which various subjects are placed with high accuracy, a complicated mechanism and a control device are required, resulting in an increase in manufacturing cost.

  Further, if 360 degrees per rotation is 5 seconds and full scan shooting is performed at 1000 projections, reading must be performed at a rate of 5 ms per projection, that is, 200 frames / second. However, driving a flat panel detector using a-Si or a-Se and a thin film transistor having an area of 17 inches (43 cm × 43 cm) or more at this speed does not provide sufficient semiconductor characteristics of a-Si or a-Se. So there was a problem that was very difficult.

  In addition, since the noise is large and the sensitivity is low, the conventional flat panel detector has a problem that the sensitivity is insufficient in a cone beam X-ray CT imaging apparatus that has a much lower X-ray dose per projection than a still image. .

  I. Sensitivity is advantageous I. In a TV camera, I.D. I. There was a particular problem. In other words, the essential problem of distortion of the image in the peripheral part, even the current maximum 17-inch size is a circular screen, so a sufficiently large area cannot be taken as a flat panel detector, the reconstruction area in cone beam X-ray CT There is a problem that the magnetic field is extremely small, and magnetism is used for electron beam control. If this is rotated, the influence of geomagnetism is complicated.

  On the other hand, in Patent Document 2, measures against fluctuations in the X-ray source, rotational shake, and body movement when the subject is rotated are not taken into consideration, and in Patent Document 3, only measures against body movement are taken into consideration. However, countermeasures against fluctuations in the X-ray source and rotational shake were not considered. Further, in Patent Document 1, no measures have been disclosed regarding the reading speed and sensitivity required for the subject rotation cone beam X-ray CT imaging apparatus.

  The present invention has been made in view of the above-described conventional problems, and an object of the present invention is to provide a low-cost, high-speed, high-sensitivity subject rotating cone beam radiation CT imaging apparatus, radiation CT imaging system, and the same. The object is to provide a radiation CT imaging method. Thereby, a high-performance radiation CT apparatus for in-vehicle use and group examination can be provided at low cost.

  In order to solve the above-described problem, the present invention provides a radiation CT imaging apparatus that detects a rotation angle of a rotation unit that rotates the subject while irradiating the subject with X-rays and a rotation angle signal, and generates a rotation angle signal. A rotation angle detection unit; a radiation image sensor panel that generates projection image data of a subject according to the rotation; and a unit that controls a signal accumulation time of the radiation image sensor panel using the rotation angle signal. Features.

  According to the present invention, in the radiation CT imaging apparatus, a rotation unit that rotates the subject around a rotation axis while irradiating the subject with radiation, and a rotation that detects a rotation angle of the rotation unit and generates a rotation angle signal. An angle detection unit, an X-ray image sensor panel that generates projection image data of a subject according to the rotation while being fixed relative to the rotation axis, and a radiation image sensor panel using the rotation angle signal And means for controlling the signal accumulation time.

  According to the present invention, in the cone beam radiation CT imaging apparatus, a rotation unit that rotates the subject while irradiating the subject with cone beam radiation, and a rotation angle that detects a rotation angle of the rotation unit and generates a rotation angle signal. And detecting means.

  In the radiation CT imaging system, the present invention detects an X-ray source that irradiates a subject with radiation, means for controlling the radiation source, rotation means for rotating the subject, and a rotation angle of the rotation means. A rotation angle detection means for generating a rotation angle signal; a radiation image sensor panel for generating projection image data of a subject in accordance with the rotation; and a signal accumulation time of the radiation image sensor panel is controlled using the rotation angle signal. And means.

  According to the present invention, in the radiation CT imaging method, the subject is rotated while irradiating the subject with continuous radiation, a rotation angle signal is detected according to the rotation, and the rotation angle signal is used to cope with the rotation angle. Projection image data to be acquired is acquired by a radiation image sensor panel that is fixed relative to the rotation axis.

  In the present invention, there is no need to wait until the rotation is stabilized, and unnecessary rotation time can be reduced. A dedicated reference element is not required. Since the panel can be used as it is, the system is simplified and the cost is reduced. Further, it is not necessary to prepare a reference element on the X-ray source side, and an X-ray source for general imaging can be used. In addition, since the rotating device and the X-ray image sensor need only be arranged in the general imaging room, the equipment cost can be reduced. Since the rotating device may use a general mechanism, cost reduction is easy. It can handle even if there is a little uneven rotation. A mechanism for accurately controlling rotation is not required. Since the system configuration is simple, maintenance is easy and it is convenient for mounting on a general examination car.

  According to the present invention, a large-sized X-ray image sensor panel using a plurality of CMOS type image sensors and a novel driving method realize a low-cost, high-speed, high-sensitivity subject rotation type cone-beam radiation imaging apparatus. be able to. As a result, a high-performance radiation CT imaging apparatus for in-vehicle use and group examination can be provided at low cost.

  Next, the best mode for carrying out the invention will be described in detail with reference to the drawings. Note that components having the same function are denoted by the same reference symbols throughout the drawings for describing the embodiments, and the repetitive description thereof will be omitted.

(First embodiment)
FIG. 1 is a diagram illustrating a schematic configuration of a first embodiment of a subject rotation type cone beam radiation CT imaging apparatus according to the present invention. In the figure, 109 is an X-ray generator (X-ray source), 108 is an X-ray source power supply, X is irradiated continuous X-rays, 107 is a subject, 104 is a signal storage type reference element, and 102 is a rotating device (rotating table). ), P is a rotating shaft of the rotating device, 103 is a rotary encoder, 101 is an X-ray image sensor panel, 105 is an imaging control device, and 106 is a display device. In the present embodiment, X-rays are used as radiation, but α rays, β rays, γ rays, and the like can also be used. This is the same in all embodiments.

  The X-ray generator 109 uses an apparatus for general imaging and is used in a continuous X-ray mode. The imaging control device 105 only sets the tube voltage, tube current, and irradiation time, and controls the start and stop of irradiation. The geometric arrangement of the X-ray generator 109, the turntable 102, and the X-ray image sensor panel 101 is set strictly. It sets so that the perpendicular drawn from the X-ray focus to the X-ray image sensor panel may pass through the rotation axis. The distance between the X-ray focal point and the rotation axis P of the turntable 102 is L, and the distance between the rotation axis P and the X-ray image sensor panel 101 is M. By changing L and M, the magnification of the projection image, the cone beam The cone angle Φ can be set.

  The turntable 102 and the X-ray image sensor panel 101 are movable and fixed in place after determining their geometrical arrangement. In the cone beam X-ray CT imaging apparatus, when the cone angle is increased, an error during reconstruction is increased. Therefore, the cone angle Φ is generally set to 10 degrees or less. In that case, L becomes considerably long, and the enlargement ratio of the projection image is considerably smaller than general X-ray CT.

  The rotating device 102 is a device that continuously rotates a rotating portion based on a rotation control signal from the imaging control device 105. A holding device (not shown) for holding the subject 107 during rotation is installed in the rotating portion. Further, a rotary encoder 103 that measures the rotation angle of the rotation device 102 and outputs the rotation angle to the imaging control device 105 is provided. This is because a signal is generated every 0.36 degrees per projection when photographing 360 degrees per rotation with 1000 projections in full scan. The angle of rotation is measured and a signal for the angle is generated. As long as it is a rotation angle detection means, it is not limited to the rotary encoder 103.

  In the subject rotation type cone beam X-ray CT imaging apparatus, as described above, when 360 degrees per rotation is set to 5 seconds and full scan imaging is performed with 1000 projections, it takes 5 ms per projection, and the X-ray source is pulse-driven. Will become difficult. In particular, when an X-ray source for general imaging is used, it is difficult to irradiate stable pulsed X-rays due to poor rise and fall of the response tube current. Therefore, in this embodiment, a configuration using continuous X-rays is adopted. Further, in imaging using cone beam X-rays, it is necessary to acquire an image with no time delay in all areas during this 5 ms, and this embodiment has a structure described later in order to cope with uneven rotation. An X-ray image sensor panel and a signal storage type reference element are used to prevent this problem from occurring.

  FIG. 2 shows the system configuration of this embodiment. The imaging control device 105 includes an image data memory 203 that stores projection image data output from the X-ray image sensor panel 101, a correction data memory 204 that stores correction data, an image processing unit 202, a device control unit 207, and an accumulation time control unit. 208, an accumulation time calculation counter 209, and a rotating device control unit 206. The X-ray image sensor panel 101 includes an A / D converter 201 and a drive unit 205.

  The imaging control device 105 controls the X-ray irradiation from the X-ray generation device 109 based on the imaging conditions input from the observer, controls the field-of-view mode of the X-ray image sensor panel 101, and controls the number of pixels and the frame rate. I do. The imaging control apparatus 105 controls the accumulation time of the signal accumulation type reference element 104 and the X-ray image sensor panel 101 by the accumulation time control unit 208 based on the signal from the rotary encoder 103.

  The image processing unit 202 performs three-dimensional imaging of the subject 107 based on pre-processing such as gamma correction, image distortion correction, logarithmic conversion, and sensitivity non-uniformity correction of the X-ray image sensor panel 101, and a pre-processed projection image (projection data). Reconstructing a three-dimensional X-ray distribution image, which is a typical X-ray absorption coefficient distribution, and subjecting the three-dimensional X-ray distribution image to image processing such as well-known volume rendering processing or maximum value projection processing. A three-dimensional X-ray image that is an X-ray tomographic image or a three-dimensional two-dimensional image is generated from the line distribution image.

  Next, the structure and function of the signal storage type reference element 104 and the X-ray image sensor panel 101, the relationship between the rotary encoder 103, and the function and effect, which are features of this embodiment, will be described. FIG. 3A is an equivalent circuit diagram of a conventional general reference element. The photodiode PD (301) converts light generated by a scintillator (not shown) by X-ray irradiation into an electrical signal. A switch SW1 (305) made of a MOSFET is connected in series to the photodiode PD (301), and is further connected to the inverting terminal of the operational amplifier Amp (304).

  A feedback resistor R1 (306) is connected to the operational amplifier Amp (304) to form a current-voltage conversion circuit, and an input current is output as a voltage signal. The voltage V (303) is applied to the non-inverting terminal of the operational amplifier Amp (304) with respect to the ground (GND). When a positive read pulse enters the gate of the switch SW1 (305), the switch SW1 (305) is opened, the photodiode PD (301) is in a reverse bias state, and a constant charge is charged in the junction capacitor C1 (302). The

  Next, when the switch SW1 (305) is closed and light is incident during the accumulation period, the charged charge is discharged by the charge due to the light incidence, and the cathode potential of the photodiode PD (301) approaches the ground potential. . This discharge charge amount increases in proportion to the amount of incident light. Next, when the read pulse enters the gate of the switch SW1 (305) and the switch SW1 (305) is opened, a charge corresponding to the charge discharged within the accumulation time is supplied via the feedback resistor R1 (306). The photodiode PD301 is initialized again in the reverse bias state.

  At this time, a potential difference due to the charging current is generated at both ends of the feedback resistor R1 (306), and is output as a voltage signal from the operational amplifier Amp (304). Since this charging current corresponds to a discharge current due to light incidence, the amount of incident light is detected by this output voltage. Thus, the signal according to the X-ray dose rate at a certain time is simply used.

  FIG. 3B is an equivalent circuit of the signal storage type reference element according to the present embodiment. An integration capacitor C2 (308) and a reset switch SW2 (306) are connected between the inverting terminal and the output terminal of the operational amplifier Amp (304), thereby constituting a current integration circuit as a whole. Immediately before the read switch SW1 (305) is opened, the integration capacitor C2 of the integration circuit is discharged by an external reset pulse. Next, when the switch SW1 (305) is opened, a charge corresponding to the light output during the accumulation period is charged from the power source to the junction capacitor C1 (302) of the photodiode PD, and the potential of the photodiode PD301 is positive potential V (303. The integration capacitor C2 (308) is also charged by the charging current. Therefore, a square wave integrated waveform is obtained at the output terminal of the integrating circuit. The read switch SW 1 (305) and the reset switch SW 2 (306) are driven by a pulse signal created by the accumulation time control unit 208 based on the angle signal from the rotary encoder 103.

  The rotary encoder 103 generates a rotation angle signal every 0.36 degrees per projection when shooting 360 degrees per rotation with 1000 projections in full scan. This signal is sent to the accumulation time control unit 208, and operation timing pulses of SW1 and SW2 are created based on this signal. As a result, the output signal is accumulated in the accumulation time for one projection angle. Therefore, even if there is uneven rotation on the turntable, the X-ray dose for the time corresponding to the rotation angle per projection determined accurately can be monitored by the output of the stored signal of the signal storage type reference element 104. Can do.

  The signal storage type reference element 104 is not limited to this equivalent circuit. What is necessary is just to have a circuit capable of accumulating sensor signals within a time set by an external signal and outputting the integration amount as a signal.

  FIG. 5 shows an X-ray image sensor panel that realizes a large area of 408 mm × 408 mm by two-dimensionally bonding a 136 mm × 136 mm CMOS image sensor to nine image sensors 501 on a single base. Show. This is because a CMOS type image pickup device is used so that charges accumulated in a common time from all the devices can be read out with a high signal-to-noise ratio (S / N). In addition, the image sensor referred to in the present specification refers to an image sensor panel in which a plurality of pixels are two-dimensionally arranged. As will be described later, the entire surface of this CMOS image sensor panel is a pixel region, and by bonding a plurality of image sensor panels on a base, it is possible to realize an image pickup device having a large area without a joint on the image. . In principle, there is no limit to its size.

6 shows a cross-sectional view taken along the line AA ′ of FIG. The scintillator 602 is composed of Gd ₂ O ₂ , SCsI or the like using europium, terbium or the like as an activator. In this embodiment, since the large X-ray image sensor panel 101 is used, a large plate having no joints with the scintillator 602 is used. One scintillator and a plurality of image sensors 501 are integrally configured to form an X-ray image sensor panel.

Since the X-rays strike the scintillator 602 and are converted into visible light and detected by the image sensor 501, the scintillator 602 is preferably selected so that its emission wavelength matches the sensitivity of the image sensor. Furthermore, the thickness of the scintillator 602 is preferably increased in order to increase the X-ray absorption rate and increase the utilization efficiency. In addition, when Gd ₂ O ₂ is used, it is difficult to produce a large plate with a large thickness, and when a normal powder is used, it exhibits a high X-ray absorption rate and is thick. Then, since there is a defect that efficiency (DQE: detective quantum efficiency) is greatly deteriorated, in this embodiment, CsI (Tl) formed by vapor deposition on a large-sized resin substrate on which a thick film is easily formed is used. ing.

  In a medical radiation CT imaging apparatus, the pixel size may be as large as about 500 μm × 500 μm to 1 mm × 1 mm. CsI (Tl) having a thickness of 1 mm has an X-ray absorption rate of 90% or more, and the efficiency (DQE) can be maximized within the resolution range required for the X-ray CT. When the thickness is 1.5 mm or more, the DQE deteriorates sharply. Therefore, in this embodiment, CsI (Tl) having a thickness of 1 mm is used.

  The external processing substrate 605 is a substrate having a circuit for supplying a power source, a clock, and the like of the image sensor 501 and extracting and processing a signal from the image sensor 501. The flexible substrate 603 performs electrical connection between each image sensor 501 and the external processing substrate 605.

  The nine image pickup elements 501 are bonded on the base 604 so that there is substantially no gap between the image pickup elements, and the fact that there is substantially no gap is that an image formed by the nine image pickup elements. This means that there is no gap between the image sensors. The imaging device clock and the like, the input of the power source and the output of the signal from the imaging device are transmitted to the external processing substrate 605 disposed on the back side of the imaging device through the flexible substrate 603 connected to the electrode pad at the end of the imaging device. Between. The thickness of the flexible substrate 603 is sufficiently small with respect to the size, and no defect on the image occurs even if it passes through the gap between the imaging elements. Outputs from the respective image sensors are read out in parallel. High-quality images that are temporally and spatially seamless and have no time delay for each projection can be acquired at a high speed by collective exposure of the image sensor described later and this parallel reading.

  FIG. 7 shows one image sensor that constitutes the X-ray image sensor. A CMOS type image pickup device substrate of 136 mm × 136 mm is produced from a currently mainstream 8-inch wafer 704 by a CMOS process. In a medical cone-beam X-ray CT imaging apparatus, the size of a pixel may be as large as about 500 μm × 500 μm to 1 mm × 1 mm. In this embodiment, the pixel size is 500 μm × 500 μm. In the present embodiment, the nine image pickup elements are arranged, but the size of the image pickup element of the present invention can be arbitrarily selected, and as a result, the number of the image pickup elements can be appropriately changed. Preferably, the size and number of images are determined in consideration of the yield of the image sensor, the yield in the mounting process, and the like. Reference numeral 701 denotes a vertical shift register, 702 denotes a horizontal shift register, and 703 denotes an external terminal.

  FIG. 8 shows the configuration (plan view) of the image sensor of this embodiment. In this embodiment, a vertical shift register 701 and a horizontal shift register 702 are arranged in the effective area of the image sensor, and a plurality of pixels 807 are arranged two-dimensionally in the vertical and horizontal directions in the image sensor. Further, one block of the shift register that processes one line is arranged so as to be within one pitch, and these blocks are arranged to form a series of vertical shift register blocks and a horizontal shift register block. These blocks extend linearly in the vertical and horizontal directions.

  Furthermore, at least the light receiving region has the same area for all pixels. In FIG. 8, the area of one pixel circuit and the area of a light receiving region in one pixel circuit are equal between cells. In addition, it is preferable that the area of the light receiving region is the same among all the cells, but the area of the light receiving region in the cell in one line at the end of the image sensor has an internal cell in order to take a margin for slicing. It may be different from the area of the light receiving region. A bump 801 is provided on the external terminal 703, and a protective resistor 802 and a protective diode 803 for protecting the internal circuit from static electricity are connected to the bump 801. The external terminal 703 is used for connection with the flexible substrate 603 as described above.

  Note that a multiplexer is built in the image pickup device in order to speed up the operation of the image pickup device. Further, a signal is taken out from the imaging device via the external terminal 703, and there is a large stray capacitance around the external terminal 703. Therefore, by providing an amplifier in front of the external terminal 703, signal transmission characteristics can be compensated.

  FIG. 9 shows a state in which unit blocks (units for selecting and driving one row) of the vertical shift register are arranged together with one pixel circuit in one region (one cell). One pixel circuit is shown in FIG. The area of the unit block and the pixel circuit does not reflect the actual element layout because it is a schematic diagram. The vertical shift register is a simple circuit composed of a static shift register and a transfer gate for generating the reset signal ΦRES, the clamp signal ΦCL, and the selection signal ΦSEL1. Signal lines other than the reset signal ΦRES, the clamp signal ΦCL, and the selection signal ΦSEL1 are omitted.

  These are driven by signals from a clock signal line (not shown). The circuit configuration of the shift register is not limited to this, and an arbitrary circuit configuration can be adopted by various driving methods such as addition and thinning readout. However, as in the present embodiment, functional blocks are arranged together with pixel circuits in one cell, a shift register is provided in the effective area, and an imaging device in the entire effective area is realized.

Further, instead of a shift register, an n-to- ²ⁿ decoder can be used as the scanning circuit. In this case, by connecting the output of the counter that increments sequentially to the input of the decoder, it becomes possible to scan in the same way as the shift register. Thus, an image of an arbitrary region can be obtained by random scanning. The common processing circuit arranged in each area (cell) in the effective area means a circuit that collectively processes a plurality of signals such as a final signal output amplifier, a serial / parallel conversion multiplexer, a buffer, and various gate circuits. .

  FIG. 10 shows a one-pixel circuit. Sensitivity (dynamic range) is realized by mode switching by performing kTC correction in the pixel in the photoelectric conversion unit and further providing sensitivity switching means in the pixel. In order to increase sensitivity, the capacitance of the photodiode PD is made as small as possible. At this time, the dynamic range is reduced. Therefore, a capacitor C1 is provided in parallel with the photodiode PD in order to ensure a dynamic range.

M1 is a changeover switch for switching the sensitivity. Photodiode capacitance C _PD to accumulate charges is designed to the minimum capacity to the maximum sensitivity at the time of shooting. M2 is a reset MOS transistor (reset switch) for discharging charges accumulated in the photodiode capacitor _CPD , M3 is a selection MOS transistor (selection switch) for selecting the pixel amplifier 1, and M4 functions as a source follower. This is an amplification MOS transistor (pixel amplifier 1). By making these circuit elements for each pixel on the same wafer, the capacity of the gate portion of M4 can be reduced as much as possible, and an improvement in sensitivity can be realized.

  Conventionally, a source follower is not suitable for a signal amplifier for CT from the viewpoint of dynamic range, but in this embodiment, a source follower configuration can be adopted by introducing sensitivity switching in the photodiode section. By using a source follower for each pixel, the sensitivity can be greatly improved, and an X-ray image sensor panel suitable for a cone beam X-ray CT imaging apparatus can be realized. Furthermore, since each pixel has a source follower, non-destructive readout is possible. With this non-destructive read function, various data can be read independently of the main read. For example, it is possible to provide a function of monitoring the amount of charge accumulated in the photodiode for each pixel, read this nondestructively outside, and switch the sensitivity before reaching saturation.

A clamp circuit is provided after the pixel amplifier 1. This clamp circuit removes kTC noise generated in the photoelectric conversion unit. C _CL is a clamp capacitor, and M5 is a clamp switch. Noise removal can be performed by the following operation. The electrode of the clamp capacitor C _CL which turns on the switch M5 is in pixel amplifier M7 side at a fixed potential.

When the photodiode PD is reset by the reset switch M2 in this state, the noise component is accumulated in the electrode of the clamp capacitor _{CCL on} the amplification MOS transistor M4 (pixel amplifier 1) side. When the signal charge is accumulated in the photodiode PD after the switch M5 is turned off, the electrode potential of the clamp capacitor _CCL on the amplification MOS transistor M4 (pixel amplifier 1) side becomes a noise component from the photodiode signal (including the noise component). And the potential of the amplifier amplifying MOS transistor M7 (pixel amplifier 2) of the clamp capacitor _CCL fluctuates as the noise component is removed. Thus, the clamp capacitor C _CL holds the signal from which the noise component has been removed.

  A sample hold circuit is provided after the clamp circuit. M6 is a selection MOS transistor (selection switch) for selecting the pixel amplifier 2. M8 is a sample MOS transistor switch constituting a sample hold circuit for storing optical signals, and CH1 is a hold capacitor. M10 is an amplification MOS transistor (pixel amplifier 3) that functions as a source follower, and M9 is a selection MOS transistor (selection switch) for selecting the pixel amplifier 3. M11 is a sample MOS transistor switch constituting a sample and hold circuit for accumulating noise signals, and CH2 is a hold capacitor. Further, M13 is an amplification MOS transistor (pixel amplifier 3) that functions as a source follower, and M12 is a selection MOS transistor (selection switch) for selecting the pixel amplifier 3.

  In general, in an amplification type imaging device such as a CMOS type imaging device, in order to improve a signal-to-noise ratio (S / N) at the time of reading, an amplification means (in-pixel amplifier) is provided inside to increase a signal gain. Yes. In the imaging device of this embodiment, a source follower of a MOS transistor used as an amplifying unit is used. In general, the threshold value Vth of a MOS transistor tends to vary. This variation is inherent in the design and manufacture of the device and is malicious in that it varies from pixel to pixel and from device to device. In particular, in a large-size imaging device as used in the present embodiment, the variation in the device tends to be large. In addition, when a plurality of image sensors are used, the variation between the elements is large. This variation appears as a fixed output variation, so-called fixed pattern noise (FPN), or a non-uniform background image.

  In addition, 1 / f noise (flicker noise) and thermal noise are likely to occur in the MOS transistor, and this is random noise, so that a random background image is generated. In terms of device design, if the channel length of the MOS transistor is L and the channel width is W, the thermal noise is proportional to (L / W) · 1/2 and the 1 / f noise is inversely proportional to L · W. In order to reduce the noise of the transistor, the channel length L should be minimized and the channel width W should be set large. However, when the channel width W of the source follower as an amplifier which becomes a large noise source is particularly set large, the gate-drain gap is reduced. This increases the parasitic capacitance of the device, lowers the gain, and lowers the sensitivity, which is difficult to implement.

  In this embodiment, a PMOS transistor with essentially low 1 / f noise is used as at least a source follower. As a result, the size can be reduced to about 1/10 that of an NMOS transistor. Further, even if the X-rays that have passed through the scintillator directly hit the transistor, the PMOS transistor is more preferable because the X-ray durability is stronger (leakage current increase and threshold Vth fluctuation is smaller) than the NMOS transistor.

  Further, since the threshold value Vth changes exponentially with temperature, even if each source follower has a temperature difference of 1 ° C. or less during photographing, it appears as an output fluctuation. In the case of an X-ray image sensor composed of a plurality of image sensors as in the present embodiment, if the temperature dependence differs for each image sensor, this slight variation also causes a correction error. In the case of X-ray CT imaging, if there is a correction error even with one image sensor, a ring artifact that is very conspicuous on the image occurs, which becomes a problem. For this reason, the two source followers of the sample and hold circuit have an arrangement structure in which the variation in the threshold value Vth is as small as possible in terms of layout as will be described later, and a mechanism that does not cause a temperature difference during operation.

  Therefore, as described above, the sample signal hold circuit for the optical signal and the noise signal is provided in the pixel, and the optical signal and the noise signal are stored independently of the exposure and are simultaneously output from the sample hold circuit (each column 2 Line output).

  As described above, in cone beam X-ray CT imaging using continuous X-rays, each projection image needs to be acquired by driving the entire screen at the same time and the same accumulation time. Therefore, a structure is provided in which a memory is provided in each pixel. Projection image signals having the same time and the same accumulation time are stored in the in-pixel memory, and while the next projection image is acquired, the stored projection data can be read at high speed by parallel reading. As described above, the sample-and-hold circuit functions as means (in-pixel memory) for storing the image signal independently of the exposure. Combined with the parallel readout of the image sensor described above, cone beam X-ray CT imaging in a large area that could not be realized conventionally by these structures and functions can be realized.

  Further, this circuit has a function of removing noise. Since the optical signal and the noise signal are taken into the sample hold circuit from the pixel amplifier 1 with a very fast time difference, a large 1 / f noise can be ignored at a low frequency.

  Further, this circuit is used to remove thermal noise, 1 / f noise, and FPN in the pixel amplifier. The variation of the two sample-and-hold circuit elements is to arrange the capacitor as close as possible in the pixel, and the output source follower is a cross arrangement used in the normal MOS circuit layout to reduce the variation of the threshold Vth as much as possible. It is reduced as much as possible. As described above, the sample and hold circuit according to the present embodiment functions as an accumulation unit for each pixel for collective exposure, also functions as a unit for noise removal, and further eliminates output fluctuations between image sensors due to temperature changes. Is given.

  As a feature of the present invention, each projection image acquisition time is controlled using an angle signal output from the rotary encoder 103. The read switch and the reset switch are driven by a pulse signal created by the accumulation time control unit 208 based on the angle signal from the rotary encoder 103.

  The angle signal output from the rotary encoder 103 is sent to the accumulation time control unit 208, and a batch reset timing and a batch exposure operation timing pulse are generated based on this signal. By controlling the operation timing and the operation timing of the sample hold circuit, a projection image signal is accumulated and output so as not to cause a time delay in each projection image, and during an accumulation time for one projection angle. As a result, even if there is uneven rotation on the turntable, it is possible to acquire a projection image for a time corresponding to the rotation angle per projection determined accurately. This storage time is completely synchronized with the storage time of the signal storage type reference element 104 described above.

  FIG. 11 is a timing chart showing the operation timing of the pixel unit according to the present embodiment. FIG. 11A shows the output with respect to Δθ of the encoder, FIG. 11B shows the accumulation time, FIG. 11C shows ΦSH1, FIG. 11D shows ΦRES, FIG. 11E shows ΦCL, and FIG. Is ΦSH2, FIG. 11 (g) is a clock, and FIG. 11 (h) is an accumulation time of a signal accumulation type reference element.

  Next, the circuit operation will be described with reference to FIG. First, photoelectric conversion is performed by the photodiode PD. The exposure is a batch exposure and is performed at the same timing and period for all the pixels of each image sensor. Therefore, there is no time shift of the image between the image pickup devices and between the scanning lines.

  First, as shown in FIG. 11A, a signal corresponding to the start angle of the Nth projection is sent from the rotary encoder 103 to the accumulation time control unit 208.

  As shown in FIG. 11C, the optical signal from which noise accumulated in the previous projection is removed by setting the signal ΦSH1 to the high level for all the pixels simultaneously from the accumulation time control unit 208 and turning on the sample switch M8. Are collectively transferred to the capacitor CH1 through the pixel amplifier 2 (M7).

Further, the signal ΦRES a high level at all pixels as shown in FIG. 11 (d), the photodiode capacitance C _PD is reset by turning on the reset switch M2. Signal accumulation corresponding to the Nth projection starts from the time when the reset is completed. Since the signal ΦSH1 is controlled by the accumulation time control unit 208 and other signals are determined according to ΦSH1, signal accumulation in the photodiode is controlled for each projection.

At the same time, a high level signal ΦCL as shown in FIG. 11 (e), set to the reference voltage clamp capacitor C _CL by turning on the selection switch M3 of the pixel amplifier 4, the clamp switch M5.

  At the same time, as shown in FIG. 11 (f), the noise signal when the signal ΦSH2 is set to the high level and the selection switch M6 and the sample switch M11 of the pixel amplifier 6 are set to the reference voltage is turned on for all the pixels at the same time. Transfer to capacity CH2. Next, the signal ΦSH2 is set to the low level for all the pixels at the same time, and the transfer holding of the optical signal and the noise signal to the sample hold circuit is finished.

  Next, the signal ΦSEL2 is set to the high level for each row by the signal input to the shift register VSR, and the selection switches M9 and M12 are turned on to turn on the source composed of the load current source and the pixel amplifiers 3 and 4 (M10 and M13). The follower circuit is set in an operating state. Thus, the optical signal and the noise signal held in the hold capacitors CH1 and CH2 are simultaneously transferred to the noise signal output line and the optical signal output line through the pixel amplifiers 3 and 4.

  As shown in FIG. 11B, the accumulation time T1 of the X-ray image sensor panel 101 is reached by the above operation. In this way, control is performed for each frame by the encoder output.

  Further, as shown in FIG. 11 (h), the accumulation time of the signal accumulation type reference element 104 is similarly controlled from the accumulation time control unit 208 by SW1 and SW2, and the same accumulation time T1 as that of the X-ray image sensor panel 101 is set for each projection. Are accumulated and output.

  By repeating the above operation, the accumulation time of the X-ray image sensor panel 101 and the accumulation time of the signal storage type reference element 104 corresponding to each projection are controlled by the output of the encoder.

  Next, a procedure for collecting projection images for one rotation of the subject in the subject rotation type cone beam X-ray CT imaging apparatus of the present embodiment will be described. First, the observer fixes the subject 107 to the holding device. Next, when the observer gives an instruction to start measurement from a console (not shown), measurement of a projection image (projection data) is started according to a control signal from the imaging control device 105, and the rotation device 102 starts to rotate. At this time, the rotation angle is output from the rotary encoder 103 to the imaging control device 105.

  When the imaging control device 105 detects that the rotation angle of the rotation device 102 has reached a predetermined angle, the X-ray generation device 109 immediately irradiates continuous X-rays. Along with the X-ray irradiation from the X-ray generator 109, the imaging control device 105 controls the X-ray image sensor panel 101 to acquire a projection image of the subject 107. The data is output to the imaging control device 105 as projection data that is a digitized projection image. The imaging control device 105 collects the projection image captured by the X-ray image sensor panel 101 together with the rotation angle of the rotation device 102, that is, the projection angle, stores the projection image in the image data memory 203, and continuously performs it for each predetermined angle. The shooting of the rotation projection image is completed.

  When the imaging (collection) of projection images for all the circumferences is completed, the imaging control device 105 ends the rotation of the rotating device 102. In the image processing apparatus 202, gamma correction, image distortion correction, logarithmic conversion, and X-rays of each projection image are performed using the projection data stored in the image data memory 203 by using the correction data already stored in the correction data memory 204. Pre-processing such as sensitivity correction of the image sensor panel is performed. Furthermore, a three-dimensional X-ray distribution image is reconstructed based on the projection image. Furthermore, a known volume rendering process or maximum value projection process is performed on the three-dimensional X-ray distribution image, and a three-dimensional X-ray image that is a three-dimensional two-dimensional image is obtained from the three-dimensional X-ray distribution image. The three-dimensional X-ray image is generated and displayed on the display screen 106.

  In the present embodiment, the light receiving areas are uniformly sized between the image sensors and between the image sensors, and the centers of gravity are arranged at an equal pitch. Since variations in sensitivity within the imaging device and variations in the center of gravity of the light receiving area do not occur, a substantially seamless image can be obtained even with a configuration in which a plurality of sheets are pasted. In addition, since no dead space is generated around the image sensor, the entire surface of the image sensor is an effective area.

  By arranging these imaging elements so that there is substantially no gap, an imaging device having a large area can be formed. Furthermore, with the circuit configuration as described above, it is possible to obtain a large-area image that is substantially seamless in terms of time and space. Since the size of the pixel is sufficiently large, a sufficiently large aperture ratio can be realized even if a shift register is arranged in the effective pixel region or a circuit such as a sample hold is arranged in the pixel, so that there is no problem. Furthermore, since each image sensor has a structure for reading signals in parallel, high-speed reading can be performed.

  In this embodiment, since the shift register is arranged in the effective region, the X-ray that has passed through the scintillator 602 directly hits the shift register. However, by using a static shift register as the shift register, it is not affected by the X-ray. I have to. The shift register circuit is used for sequentially transferring pulse signals. That is, in principle, the static type is relatively less susceptible to the influence of X-rays, so that it can be used in a place where X-rays directly hit as in this embodiment. Therefore, if a static shift register is used, an imaging apparatus with reduced X-ray damage and errors and improved reliability can be realized.

  Furthermore, in the present embodiment, since a CMOS type image sensor is used as the image sensor, the power consumption is low, which is suitable for the case of configuring a large area image pickup apparatus. In addition, the sample-and-hold circuit provided in each pixel can realize full-screen batch exposure, FPN noise removal, and artifact reduction due to further temperature change. It is possible to realize an X-ray image sensor panel with a large area capable of performing CT imaging of all the lungs of a human body with high sensitivity over 200 frames / sec, high sensitivity similar to fluoroscopic imaging. Furthermore, artifacts due to rotational angle fluctuations can be eliminated by controlling the entire screen batch exposure and the accumulation time for each projection.

(Second embodiment)
In the second embodiment, an ionization dosimeter is used as a signal storage type reference element. In general, an ionization dosimeter (area dosimeter) is used to measure the exposure of a patient during fluoroscopy, attached to an X-ray fluoroscope. An area dosimeter is attached to the movable diaphragm of the X-ray generator. Other than the area dosimeter is the same as in the first embodiment.

  In the present embodiment, the area dosimeter is installed so as to cover the front surface of the operation diaphragm of the X-ray generator. The value measured by the area dosimeter is the total dose of X-rays emitted from the range limited by the operating diaphragm among the X-rays emitted from the X-ray generator. In the first embodiment, the signal storage type reference element is installed at a position where it does not become a shadow on the subject. Since the area dosimeter almost transmits X-rays, there is no problem even if it is installed so as to cover the front surface of the operation diaphragm of the X-ray generator. As a result, the whole (average) of the X-rays irradiated to the subject can be monitored. Therefore, when the dose correction of the projection image is performed using this value, the accuracy is improved as compared with the case of the first embodiment.

  Further, not only an area dosimeter but also an X-ray transmission type sensor can be used in the same manner. A conventional flat panel detector using a-Si on a glass substrate can also be used. That is, a flat panel detector divided into a plurality of pixels is installed so as to cover the front surface of the operating aperture. Since the pixel division at this time may be, for example, nine divisions in the range limited by the operating aperture, even a conventional flat panel detector can be used for X-ray CT imaging. When the dose correction of the projection image is performed using the X-ray dose values measured by dividing into nine, further optimized correction can be performed.

(Third embodiment)
In the third embodiment, a part of the X-ray image sensor panel is used as a reference element. The operation of the X-ray image sensor panel 101 is the same as that of the first embodiment. As shown in FIG. 4, the X-ray fluoroscopic image includes a portion where there is no projection image of the subject and a nest portion. A region 401 surrounded by a dotted line in the void portion is used as a reference signal detection unit. This region is determined by irradiating a very weak X-ray before the main scan to detect the nest-extracted portion. After determining the reference signal detection unit, the collimator for removing scattered radiation from the subject is set at the corresponding location on the X-ray image sensor panel. You may also use a grid. In this embodiment, since no separate reference detection element is required, the image of the subject and the reference signal can be simultaneously acquired as one image data. This simplifies the system.

(Fourth embodiment)
A fourth embodiment will be described with reference to FIG. This embodiment relates to a correction method. Acquisition of subject image data is the same as in the first embodiment. In this embodiment, the accumulation time calculation counter 209 in FIG. 2 is further used to monitor the accumulation time control pulse generated by the accumulation time control unit 208 at the time of photographing the subject, and the number of clocks during that time (FIG. 11 (g)). Is obtained to obtain T1 and T2 (FIG. 11B) of FIG. This is obtained for all projections and stored separately in the storage means. Data from the reference element is also stored.

  Usually, before photographing a subject, the sensitivity of the X-ray image sensor panel is measured in the absence of the subject. At this time, since it has nothing to do with the rotating device 102, the accumulation time is a constant value (5 msec), and 1000 pieces of data are acquired. This measurement consists of acquisition of dark image (including FPN) data of the X-ray image sensor panel 101 acquired without irradiation of X-rays and white image data when X-rays are irradiated. In order to correct the X-ray transmission image of the subject, average data of these dark image data and white image data is prepared.

  In the past, this 5 msec data had to be used for correction as it is, but in this embodiment, this average data is corrected by the storage time data stored for each projection, and the stored signal data of the reference element is further stored. Use to correct. Using the average data corrected in this way, white image correction and black image correction are performed for each projection. As a result, the subject image can be corrected correctly even if the X-ray source fluctuates and rotation is uneven. As a result, a high-quality cross-sectional image without artifacts can be obtained. Others are the same as in the first embodiment.

1 is a schematic diagram showing a first embodiment of a radiation CT imaging apparatus according to the present invention. 1 is a system block diagram showing a first embodiment of a radiation CT imaging apparatus according to the present invention. 2A and 2B are diagrams illustrating a reference element according to the present invention, in which FIG. 1A is an equivalent circuit diagram of a conventional reference element, and FIG. 2B is an equivalent circuit diagram of a signal storage type reference element according to the present invention. It is a figure explaining the 3rd Embodiment of this invention. It is a figure which shows the X-ray image sensor panel of 1st Embodiment by this invention. It is sectional drawing in the AA 'line of FIG. It is a top view which shows the image pick-up element of 1st Embodiment by this invention, and the wafer used as the origin. FIG. 2 is a plan view showing an array of pixels and an array of scanning circuits in the imaging apparatus according to the first embodiment of the present invention. It is a figure which shows the relationship between the 1 pixel circuit in the image sensor of 1st Embodiment by this invention, and the unit block of a shift register. 1 is a one-pixel circuit diagram of an image sensor according to a first embodiment of the present invention. It is a figure which shows the timing of the drive method of 1st Embodiment by this invention. It is a conceptual diagram which shows the structure of the conventional radiography system.

Explanation of symbols

101 X-ray image sensor panel 102 Rotating device (rotary base)
Reference Signs List 103 rotary encoder 104 signal storage type reference element 105 imaging control device 106 display device 107 subject 108 X-ray source power supply 109 X-ray generator (X-ray source)
201 A / D converter 202 drive unit 203 image data memory 204 correction data memory 205 drive unit 206 rotation control unit 207 device control unit 208 accumulation time control unit 209 accumulation time calculation counter 301 photodiode PD
302 Junction capacitance C1
303 Voltage V
304 operational amplifier Amp
305 Read switch SW1
306 Reset switch 308 Integral capacitance C2
401 Reference signal detection unit 501 Image sensor 602 Scintillator 603 Flexible substrate 604 Base 605 External processing substrate 701 Vertical shift register 702 Horizontal shift register 703 External terminal 801 Bump 802 Protection resistor 803 Protection diode 807 Pixel

Claims (14)

In radiation CT imaging apparatus,
Rotating means for rotating the subject while irradiating the subject with radiation;
A rotation angle detection means for detecting a rotation angle of the rotation means and generating a rotation angle signal;
A radiation image sensor panel for generating projection image data of a subject according to the rotation;
Means for controlling the signal accumulation time of the radiation image sensor panel using the rotation angle signal;
A radiation CT imaging apparatus characterized by comprising: 2. The radiation CT imaging apparatus according to claim 1, further comprising: a signal storage type reference signal generation unit; and a unit for controlling a signal storage time of the reference signal generation unit using the rotation angle signal. CT imaging device. 3. The radiation CT imaging apparatus according to claim 2, wherein the radiation is continuous radiation emitted from a radiation source, and the signal accumulation type reference signal generation unit is disposed between a radiation image sensor panel and the radiation source, The radiation image sensor panel has a plurality of image sensors, each image sensor has a plurality of pixels, and signals of each of the image sensors and the signal storage type reference signal generating means using the rotation angle signal. A radiation CT imaging apparatus comprising means for controlling an accumulation time. 4. The radiation CT imaging apparatus according to claim 3, wherein the signal storage type reference signal generating means is disposed on a front surface of the radiation source covering the radiation irradiation unit. 3. The radiation CT imaging apparatus according to claim 2, wherein the radiation is continuous X-ray, the radiation image sensor panel includes a plurality of image sensors, each image sensor includes a plurality of pixels, and the signal storage type. The radiation CT imaging apparatus, wherein the reference signal generating means is a part of pixels of the image sensor, and has means for controlling a signal accumulation time of the pixels of the image sensor using the rotation angle signal. 2. The radiation CT imaging apparatus according to claim 1, wherein means for calculating signal accumulation time data corresponding to the signal accumulation time, means for storing white image data, means for storing FPN image data, and the white image A radiation CT imaging apparatus comprising: data, means for performing white correction and FPN correction of each projection image data using the FPN image data and the signal accumulation time data. 7. The radiation CT imaging apparatus according to claim 6, further comprising means for correcting each projection image data using the signal data of the signal storage type reference signal generating means. In radiation CT imaging apparatus,
Rotating means for rotating the subject around a rotation axis while irradiating the subject with radiation;
A rotation angle detection means for detecting a rotation angle of the rotation means and generating a rotation angle signal;
An X-ray image sensor panel that generates projection image data of a subject according to the rotation while being fixed relative to the rotation axis;
A radiation CT imaging apparatus comprising: means for controlling a signal accumulation time of the radiation image sensor panel using the rotation angle signal. 9. The radiation CT imaging apparatus according to claim 8, further comprising: a signal storage type reference signal generation unit; and a unit for controlling a signal storage time of the reference signal generation unit using the rotation angle signal. Radiation CT imaging device. Rotating means for rotating the subject while irradiating the subject with cone beam radiation;
A rotation angle detection means for detecting a rotation angle of the rotation means and generating a rotation angle signal;
A cone beam radiation CT imaging apparatus comprising: In radiation CT imaging system,
A radiation source for irradiating the subject with radiation;
Means for controlling the radiation source;
Rotating means for rotating the subject;
A rotation angle detection means for detecting a rotation angle of the rotation means and generating a rotation angle signal;
A radiation image sensor panel for generating projection image data of a subject according to the rotation;
Means for controlling the signal accumulation time of the radiation image sensor panel using the rotation angle signal;
A radiation CT imaging system comprising: In the radiation CT imaging method,
Rotate the subject while irradiating the subject with continuous radiation,
A rotation angle signal is detected according to the rotation,
A radiation CT imaging method, wherein projection image data corresponding to the rotation angle is acquired by a radiation image sensor panel while being relatively fixed with respect to the rotation axis using the rotation angle signal. 13. The radiation CT imaging method according to claim 12, wherein a reference signal accumulation time and a signal accumulation time of the radiation image sensor panel subjected to batch exposure are controlled using the rotation angle signal. 14. The radiation CT imaging method according to claim 13, wherein signal accumulation time data corresponding to the signal accumulation time is calculated, white image data is obtained, FPN image data is obtained, the white image data, and the FPN image. Radiation wherein each projection image data is subjected to white correction and FPN correction using the data and the signal storage time data, and each projection image data is corrected using the signal data of the signal storage type reference signal generating means. CT imaging method.

Images