JP5348247B2 - Radiography equipment - Google Patents

Description

  The present invention relates to a radiation imaging apparatus including a radiation source and an FPD, and in particular, captures a series of fluoroscopic images while the radiation source and the FPD move synchronously in opposite directions, and superimposes the images to superimpose a subject. The present invention relates to a radiation imaging apparatus capable of selecting either a tomographic imaging mode capable of acquiring an image or a spot imaging mode for capturing a single fluoroscopic image by irradiating a single shot of radiation.

  In the medical institution, a radiation imaging apparatus 51 that acquires a tomographic image of the subject M is provided. In such a radiation imaging apparatus 51, a series of fluoroscopic images are continuously shot while a radiation source 53 for irradiating radiation and an FPD 54 for detecting radiation move synchronously, and the series of fluoroscopic images are overlapped to obtain a tomographic image. Some are configured to acquire images (see FIG. 13). In such a radiation imaging apparatus 51, during the imaging of a series of fluoroscopic images, the radiation source 53 and the FPD 54 move so as to approach each other along the body axis direction A of the subject M, and the radiation source 53 and the FPD 54 After the positions in the body axis direction A coincide with each other, the radiation source 53 and the FPD 54 move away from each other along the body axis direction A. Such a radiation imaging apparatus is described in Patent Document 1, for example.

  An operation when the radiation imaging apparatus 51 captures a tomographic image as described above will be described. First, the radiation source 53 emits radiation intermittently while moving. That is, every time one irradiation is completed, the radiation source 53 moves along the body axis direction of the subject M, and again performs radiation irradiation. In this way, 74 perspective images are acquired, and these are superimposed. The completed image is a tomographic image in which a tomographic image obtained by cutting the subject with a certain cut surface is reflected.

  The radiation imaging apparatus 51 can also capture a radioscopic image that simply sees through the subject M by changing the setting. In acquiring such a fluoroscopic image, the radiation source 53 and the FPD 4 are moved to predetermined positions, and the fluoroscopic image is taken only once while the subject M is placed therebetween. Such imaging is called spot imaging to distinguish it from tomographic imaging.

How the radiation irradiated from the radiation source 53 is detected by the FPD 4 will be described. The radiation incident on the FPD 4 is once converted into an electric signal, amplified at a predetermined magnification, and then converted into a digital signal. This amplification factor is an analog gain. In order to obtain a radioscopic image suitable for diagnosis, it is necessary to irradiate the subject with a certain amount of radiation. Suppose that the radiation dose is extremely reduced for the purpose of suppressing radiation exposure of the subject. At this time, if the analog gain is increased by the amount corresponding to the decrease in the dose, the noise component included in the electric signal is also amplified, so that a granular false image is generated in the obtained radioscopic image. However, unless the analog gain is increased, the electrical signal is not sufficiently amplified, and the contrast of the radioscopic image is deteriorated. As described above, the conventional configuration is set so that the analog gain is not more than a certain value and the radiation dose is not less than a certain value for the purpose of preventing deterioration of the visibility of the radioscopic image.
JP 2006-271513 A

However, the conventional configuration has the following problems.
According to the conventional configuration, there is a problem that the amount of exposure of the subject is large in the tomographic imaging mode. In spot photography, radiation is irradiated only once. However, in tomography, a plurality of fluoroscopic images are taken, so that the exposure amount of the subject becomes high. If 74 fluoroscopic images are acquired for tomography, the radiation exposure amount in tomography is 74 times that in spot imaging.

  The present invention has been made in view of such circumstances, and an object thereof is to provide a radiographic imaging apparatus in which the exposure amount of a subject is suppressed in a tomographic imaging mode.

The present invention adopts the following configuration to solve the problems described above.
That is, the radiation imaging apparatus according to the present invention synchronizes the radiation source for irradiating radiation, the radiation source control means for controlling the output of the radiation source, the radiation detection means for detecting radiation, and the radiation source and the radiation detection means. A series of fluoroscopic images that are continuously shot while the radiation source and the radiation detection means are moved, a moving means that moves the image automatically, an image generation means that generates a fluoroscopic image based on a detection signal output by the radiation detection means, Or a superimposing unit that generates a tomographic image by superimposing portions in which the same portion of the subject is captured in a series of fluoroscopic images, and the radiation detecting unit includes (A) radiation incident on the radiation detecting unit. a radiation conversion means for converting into an electric signal, amplifying means constituted by an analog type amplifier amplifies at a predetermined amplification rate (B) an electrical signal, (C) analog Day The output of the amplifying means has a an A / D converting means for converting the digital data, spotted style shooting the radiation source the radiation detecting device by irradiating single radiation to obtain a single fluoroscopic image When the radiography mode is selected and the mode of imaging for obtaining the tomographic image is the tomographic imaging mode, the amplification factor in the tomographic imaging mode is larger than the amplification factor in the spot imaging mode. This single radiation output is smaller than the single radiation output of the radiation source in the spot imaging mode.

  [Operation / Effect] The radiation imaging apparatus of the present invention has a configuration in which a spot imaging mode and a tomographic imaging mode can be selected. The radiation detection means provided in the radiation imaging apparatus of the present invention converts the radiation into an electrical signal, amplifies it, and outputs it to the image generation means. In the conventional configuration, the electric signal is amplified at a predetermined amplification factor regardless of the photographing mode. According to the present invention, the amplification factor in the tomographic imaging mode is greater than the amplification factor in the spot imaging mode. Thereby, even if the radiation intensity in the tomographic imaging mode is suppressed, the contrast of the fluoroscopic image acquired in the tomographic imaging mode is high.

  In general, when the amplification factor is increased, a noise component superimposed on an electric signal is also amplified. When a fluoroscopic image is acquired by the radiation imaging apparatus of the present invention, a false image due to a noise component is likely to appear. However, the tomographic images are acquired by superimposing a series of fluoroscopic images or a portion of the series of fluoroscopic images in which the same part of the subject is reflected. When the fluoroscopic images are compared with each other, the appearance of the false image is different. Therefore, if these are superimposed, the false image is canceled. Thus, no false image appears in the tomographic image finally obtained. The reason for the appearance of the false image between the fluoroscopic images is that the noise component superimposed on the electric signal fluctuates with time, and the plurality of fluoroscopic images are acquired at different timings.

  In addition, since spot photography is performed with a low amplification factor as in the past, the fluoroscopic image in spot photography is clear and no false image appears.

In addition, the radiation imaging apparatus according to the present invention synchronizes the radiation source that irradiates radiation, the radiation source control means that controls the output of the radiation source, the radiation detection means that detects the radiation, and the radiation source and the radiation detection means. A series of fluoroscopic images that are continuously shot while the radiation source and the radiation detection means are moved, a moving means that moves the image automatically, an image generation means that generates a fluoroscopic image based on a detection signal output by the radiation detection means, Or a superimposing unit that generates a tomographic image by superimposing portions where the same portion of the subject is captured in a series of fluoroscopic images, and the radiation detecting unit includes (A) radiation incident on the radiation detecting unit. Radiation converting means for converting the electrical signal into an electrical signal; (B) an amplifying means configured to amplify the electrical signal at a predetermined amplification factor and comprising an analog amplifier ; and (C) analog data; A / D conversion means for converting the output of the amplification means into digital data, and (D) a radiographic source that irradiates the radiation detection means with a single shot to acquire a single fluoroscopic image spot When the imaging mode is tomographic image acquisition mode and the tomographic imaging mode is the tomographic imaging mode, the amplification factor setting means for setting the amplification factor of the amplification means according to the imaging mode selected at the time of imaging is provided. The means sets the amplification factor in the tomographic imaging mode to be larger than the amplification factor in the spot imaging mode, and the radiation source control means sets the single radiation output of the radiation source in the tomographic imaging mode to the radiation in the spot imaging mode. It is characterized in that it is set so as to be smaller than the single radiation output of the source and not to saturate the amplification of the amplification means. .

  [Operation / Effect] The radiation imaging apparatus of the present invention is a more detailed description of the above-described radiation imaging apparatus including a radiation converting means and an amplifying means. That is, in addition to the above-described two components, an amplification factor setting unit that sets the amplification factor of the amplification unit according to the photographing mode is provided. The radiation source control means is configured to set the output of the radiation source. Thereby, the amplification means and the radiation source can be reliably changed according to the imaging mode set by the operator. Further, the radiation source control means sets the single radiation output of the radiation source in the tomographic imaging mode to an intensity that does not saturate the amplification of the amplification means. This configuration clarifies the lower limit of the radiation output in the tomographic imaging mode. If the radiation output of the radiation source is low, the amplification factor of the amplification means is increased accordingly. At this time, if the radiation output is too low, the amplification of the amplification means is saturated, and the brightness of the image becomes unclear. According to the above configuration, since tomographic imaging is performed with radiation having sufficient intensity so that such a saturation phenomenon does not occur, a tomographic image suitable for diagnosis can be acquired.

  Moreover, the above-mentioned structure has shown the structure which made the radiography apparatus of this invention more concrete. If the single radiation output of the radiation source in the tomographic imaging mode is smaller than the single radiation output of the radiation source in the spot imaging mode, the radiation exposure of the subject in the tomographic imaging mode is minimized. Can do. Since the tomographic image is generated by superimposing a plurality of fluoroscopic images, there is no problem even if some fake images appear in each of the fluoroscopic images.

  Further, it is more desirable that the moving means described above moves the radiation source and the radiation detecting means synchronously in opposite directions, and the superimposing means superimposes a series of fluoroscopic images to generate a tomographic image.

  [Operation / Effect] The above-described configuration is an application of the present invention to a configuration in which a tomographic image is generated by synchronously moving a radiation source and a radiation detection means in opposite directions. Since the fluoroscopic images are superimposed when generating the tomographic image, the false image appearing in the fluoroscopic image is canceled and removed.

  Further, the moving means described above moves synchronously in the same direction as the radiation source and the radiation detection means, divides a series of fluoroscopic images into strips, generates a strip-like image, and is irradiated to the radiation detection means. The image forming apparatus further includes equiangular image generating means for connecting a plurality of the same angle images while selecting the strip-like images having the same radiation angle, and superimposing means for generating a series of identical angle images. It is more desirable to generate tomographic images by superimposing them.

  [Operation / Effect] The above-described configuration is an application of the present invention to a configuration in which a radiation source and a radiation detection means move in the same direction synchronously to generate a tomographic image. Even in such a configuration, the fluoroscopic images are superimposed when generating the tomographic image, so that the false image appearing in the fluoroscopic image is canceled and removed.

Further, the amplifying means described above is constituted by analog amplifiers, that are provided by an A / D converting means for converting the output of the amplifying means adapted analog data into digital data.

  [Operation / Effect] According to the above-described configuration, the electric signal is amplified before being digitally converted. Since the electric signal before amplification is analog data, the amplification of the electric signal can be performed more faithfully if the amplification is performed while the data is in an analog form.

  It is more desirable that an input means for inputting an operator's instruction is further provided so that the photographing mode can be changed by the operator's instruction.

  [Operation / Effect] According to the above-described configuration, the tomographic imaging mode and the spot imaging mode can be selected by an operator's instruction. If the surgeon selects one of the modes, the gain is changed accordingly. The surgeon can take a picture without worrying about the amplification factor.

  The radiation imaging apparatus of the present invention has a configuration in which a spot imaging mode and a tomographic image imaging mode can be selected. The radiation detection means provided in the radiation imaging apparatus of the present invention converts the radiation into an electrical signal, amplifies it, and outputs it to the image generation means. According to the present invention, the amplification factor in the tomographic imaging mode is greater than the amplification factor in the spot imaging mode. A tomographic image is acquired by superimposing a plurality of fluoroscopic images. When the fluoroscopic images are compared with each other, the appearance of the false image is different. Therefore, if these are superimposed, the false image is canceled out. Thus, no false image appears in the tomographic image finally obtained. In addition, since spot photography is performed with a low amplification factor as in the past, the fluoroscopic image in spot photography is clear and no false image appears.

1 is a functional block diagram illustrating a configuration of an X-ray imaging apparatus according to Embodiment 1. FIG. 1 is a cross-sectional view illustrating a configuration of an FPD according to Example 1. FIG. 1 is a schematic diagram illustrating a configuration of an FPD according to Embodiment 1. FIG. 3 is a schematic diagram illustrating a table according to Embodiment 1. FIG. FIG. 6 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the first embodiment. FIG. 6 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the first embodiment. FIG. 6 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the first embodiment. 6 is a functional block diagram illustrating a configuration of an X-ray imaging apparatus according to Embodiment 2. FIG. FIG. 10 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the second embodiment. FIG. 10 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the second embodiment. FIG. 10 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the second embodiment. FIG. 10 is a schematic diagram for explaining an operation in a tomographic imaging mode according to the second embodiment. It is a functional block diagram explaining the structure of the radiography apparatus of a conventional structure.

3 X-ray tube (radiation source)
4 FPD (radiation detection means)
6 X-ray tube control unit (radiation source control means)
7 Synchronous movement mechanism (moving means)
11 Image generation unit (image generation means)
12 Superposition part (superposition means)
21 Console (input means)
40 Conversion layer (radiation conversion means)
44 Amplifier array (amplification means)
45 Amplifier array controller (amplification factor setting means)

  Hereinafter, the best mode for carrying out the invention in a plurality of embodiments will be described.

  Next, embodiments of the radiation imaging apparatus according to Embodiment 1 will be described with reference to the drawings. In addition, the X-ray in each Example is equivalent to the radiation of the structure of Example 1.

<Configuration of X-ray imaging apparatus>
FIG. 1 is a functional block diagram illustrating the configuration of the radiation imaging apparatus according to the first embodiment. As shown in FIG. 1, an X-ray imaging apparatus 1 according to the first embodiment includes a top plate 2 on which a subject M that is a target of X-ray tomography is placed, and a subject provided above the top plate 2. An X-ray tube 3 that irradiates M with a cone-shaped X-ray beam, and a sheet-like flat panel X-ray detector that is provided below the top plate 2 and detects a transmission X-ray image of the subject M ( (Hereinafter abbreviated as FPD) 4, and the X-ray tube 3 and the FPD 4 are connected to each other across the region of interest of the subject M in a state where the central axis of the cone-shaped X-ray beam and the center point of the FPD 4 always coincide with each other. An X-ray grid that absorbs scattered X-rays provided so as to cover an X-ray detection surface that detects X-rays of the FPD 4 and a synchronous movement control unit 8 that controls the synchronous movement mechanism 7 that moves synchronously in the opposite direction. And 5. Thus, the X-ray tube 3, the top plate 2, and the FPD 4 are arranged in the vertical direction in this order. The X-ray imaging apparatus 1 corresponds to the radiation imaging apparatus of the present invention, and the X-ray tube 3 corresponds to the radiation source of the present invention. The FPD 4 corresponds to the radiation detection means of the present invention. The synchronous movement mechanism 7 corresponds to the movement means of the present invention.

  The X-ray tube 3 is configured to repeatedly irradiate the subject M with a cone-shaped and pulsed X-ray beam according to the control of the X-ray tube control unit 6. The X-ray tube 3 is provided with a collimator that collimates the X-ray beam into a cone shape that is a pyramid. The X-ray tube 3 and the FPD 4 generate imaging systems 3 and 4 that capture X-ray fluoroscopic images. The X-ray tube control unit 6 corresponds to the radiation source control means of the present invention.

  The X-ray imaging apparatus 1 according to the first embodiment further includes a main control unit 25 that comprehensively controls the control units 6 and 8 and a display unit 22 that displays an X-ray tomographic image. The main control unit 25 is constituted by a CPU, and realizes the control units 6 and 8 and an image generation unit 11 and a superposition unit 12 described later by executing various programs. The superimposing unit 12 corresponds to the superimposing unit of the present invention, and the image generating unit 11 corresponds to the image generating unit of the present invention.

  The console 21 is used to input an operator's instruction. As a result, the operator can select whether to perform spot imaging or tomographic imaging, and can issue an instruction to start imaging. The console 21 corresponds to input means of the present invention.

  The synchronous movement mechanism 7 is configured to move the X-ray tube 3 and the FPD 4 in synchronization. The synchronous movement mechanism 7 linearly moves the X-ray tube 3 along a linear trajectory parallel to the body axis direction A of the subject M according to the control of the synchronous movement control unit 8. Moreover, during the examination, the cone-shaped X-ray beam irradiated by the X-ray tube 3 is always irradiated toward the region of interest of the subject M. The X-ray irradiation angle is determined by the X-ray tube 3. Is changed from, for example, an initial angle of −20 ° to a final angle of 20 °. Such an X-ray irradiation angle change is performed by the X-ray tube tilting mechanism 9. The synchronous movement control unit 8 controls the synchronous movement mechanism 7 and the X-ray tube tilt mechanism 9 synchronously. The body axis direction A corresponds to the moving direction of the present invention.

  In addition, the synchronous movement mechanism 7 moves the FPD 4 provided below the top plate 2 linearly along the body axis direction A of the subject M in synchronization with the linear movement of the X-ray tube 3 described above. The moving direction is opposite to the moving direction of the X-ray tube 3. That is, the cone-shaped X-ray beam whose irradiation source position and irradiation direction change as the X-ray tube 3 moves is always received by the entire surface of the X-ray detection surface of the FPD 4. As described above, in one inspection, the FPD 4 acquires, for example, 74 fluoroscopic images P while moving in synchronization with the X-ray tube 3 in the opposite directions. Specifically, the imaging systems 3 and 4 are opposed to the position indicated by the alternate long and short dash line illustrated in FIG. 1 through the position indicated by the broken line with the position of the solid line as the initial position. That is, a plurality of fluoroscopic images are taken while changing the positions of the X-ray tube 3 and the FPD 4. By the way, since the cone-shaped X-ray beam is always received by the entire surface of the X-ray detection surface of the FPD 4, the central axis of the cone-shaped X-ray beam always passes through the center point of the FPD 4 during imaging. During imaging, the center of the FPD 4 moves straight, but this movement is in the direction opposite to the movement of the X-ray tube 3. That is, the X-ray tube 3 and the FPD 4 are moved synchronously and in opposite directions along the body axis direction A.

  A configuration of the FPD 4 will be described. As shown in FIG. 2, the FPD 4 includes a conversion layer 40 made of amorphous selenium, an active matrix substrate 41 laminated on the conversion layer 40, and a planar electrode 42 for placing the conversion layer 40 in a predetermined electric field. It has. The active matrix substrate 41 is provided with a collection electrode 41 a for collecting carriers so as to be in contact with the conversion layer 40. The collection electrodes 41 a are arranged in a matrix along the plane in which the active matrix substrate 41 extends. Each of the collecting electrodes 41a is provided with a capacitor 41c for accumulating charges. Each of the capacitors 41c is provided with a transistor 41t for controlling charge extraction. Therefore, the capacitor 41c and the transistor 41t are two-dimensionally arranged as shown in FIG. The planar electrode 42 is covered with an insulating layer 39. The conversion layer 40 corresponds to the radiation conversion means of the present invention. One capacitor 41c and one transistor 41t connected thereto constitute a single X-ray detection element.

  The two-dimensionally arranged transistors 41t are connected to wirings extending in a lattice form vertically and horizontally. That is, the readout electrodes α of the transistors 41t arranged in the vertical direction in FIG. 3 are all connected to one of the common amplifier electrodes Q1 to Q4, and the gates G of the transistors 41t arranged in the horizontal direction in FIG. All are connected to one of the common gate control electrodes H1 to H4. The gate control electrodes H1 to H4 are connected to the gate drive 43, and the amplifier electrodes Q1 to Q4 are connected to the amplifier array 44. The amplifier array 44 is configured by arranging analog amplifiers in an array and corresponds to the amplification means of the present invention.

  A configuration for reading out charges from each capacitor 41c will be described. It is assumed that charges are accumulated in each of the capacitors 41c in FIG. The gate drive 43 turns on the transistors 41t all at once through the gate control electrode H1. The four transistors 41t arranged side by side transmit the charge (original signal) to the amplifier array 44 through the amplifier electrodes Q1 to Q4. The original signal is analog data and corresponds to the electrical signal of the present invention.

  Next, the gate drive 43 simultaneously turns on the transistors 41t through the gate control electrode H2. In this way, the gate drive 43 sequentially turns on the gate control electrodes H1 to H4. Each time, the transistors 41t in different rows are turned on. Thus, the FPD 4 is configured to read out the electric charge accumulated in each of the capacitors 41c for each row.

  In the amplifier array 44, an amplifier that amplifies a signal is provided in each of the amplifier electrodes Q1 to Q4. The original signals input to the amplifier array 44 from the amplifier electrodes Q1 to Q4 are amplified here with a predetermined amplification factor. The amplified signal which is analog data output from the amplifier array 44 is digitally converted by the A / D converter 46 and then output to the image generation unit 11. Each amplified signal is converted into a pixel value, which is two-dimensionally arranged to generate an image. The A / D converter 46 corresponds to the A / D conversion means of the present invention. The A / D converter 46 is provided for each amplifier.

  The amplifier array 44 amplifies the original signal based on the analog gain (amplification factor) output from the amplifier array control unit 45. When the amplifier array controller 45 outputs a value of 3 as an analog gain, the amplifier array 44 triples the original signal output from the transistor 41t and outputs an amplified signal. The amplifier array control unit 45 corresponds to the amplification factor setting means of the present invention.

  The storage unit 23 stores a table T in which analog gains and parameters related to control of the X-ray tube 3 are associated (see FIG. 4). Each parameter related to the control of the X-ray tube 3 for spot imaging is associated with an analog gain for spot imaging. Each parameter related to the control of the X-ray tube 3 for tomography is associated with an analog gain for tomography. The analog gain value for tomography is higher than that for spot imaging. In the tomography, parameters related to the control of the X-ray tube 3 are defined so that X-rays with a lower output than spot imaging are output. The parameters are specifically the tube current, tube voltage, and pulse width of the X-ray tube 3 (irradiation time per pulse of the X-ray beam). By setting the parameters as described above, the exposure dose when a single X-ray pulse in tomography hits the subject is suppressed to be lower than that in spot imaging. In tomography, the subject is irradiated 74 times with X-ray pulses.

  The radiation intensity per shot of the radiation beam during tomography is smaller than that during spot imaging. Such adjustment of the radiation beam output may be realized by adjusting the pulse width of the radiation beam, or by adjusting the tube current and tube voltage of the radiation beam. Further, the pulse width, tube current, and tube voltage may be adjusted together.

  Next, the principle of obtaining a tomographic image of the X-ray imaging apparatus 1 according to the first embodiment will be described. FIG. 5 is a diagram for explaining a tomographic image acquisition method of the X-ray imaging apparatus according to the first embodiment. For example, a reference cutting surface MA parallel to the top plate 2 (horizontal with respect to the vertical direction) will be described. As shown in FIG. 5, points P and Q positioned on the reference cutting surface MA are always X-rays of the FPD 4. A series of FPDs 4 are synchronously moved in the opposite direction of the X-ray tube 3 in accordance with the irradiation direction of the cone-shaped X-ray beam 19 by the X-ray tube 3 so as to be projected onto the fixed points p and q of the detection surface. The subject image is generated by the image generation unit 11. In the series of subject images Pm, the projected images of the subject are reflected while changing the position. Then, when this series of subject images Pm is superimposed by the superimposing unit 12, images (for example, fixed points p and q) positioned on the reference cut surface MA are accumulated and imaged as an X-ray tomographic image. become. On the other hand, the point I that is not located on the reference cut surface MA is reflected as a point i in a series of subject images while changing the projection position on the FPD 4. Unlike the fixed points p and q, such a point i is blurred without forming an image when the superimposing unit 12 superimposes the fluoroscopic images. Thus, by superimposing a series of subject images Pm, an X-ray tomographic image in which only an image of the subject M located on the reference cut surface MA is reflected is obtained. In this way, when the X-ray fluoroscopic images are simply superimposed, an X-ray tomographic image at the reference cut surface MA is obtained. The position in the vertical direction of the reference cutting section MA is the reference cutting position of the present invention.

  Furthermore, by changing the setting of the overlapping portion 12, a similar tomographic image can be obtained at any cutting position horizontal to the reference cut surface MA. During shooting, the projection position of the point i moves in the FPD 4, but this moving speed increases as the separation distance between the point I before projection and the reference cut surface MA increases. By using this to superimpose a series of acquired subject images while shifting in the body axis direction A at a predetermined pitch, an X-ray tomographic image at a cutting position parallel to the reference cut surface MA is obtained. The superimposing unit 12 performs superimposition of such a series of subject images. A method for acquiring a tomographic image in this way is called filter back projection.

<Operation of X-ray imaging apparatus>
Next, the operation of the X-ray imaging apparatus according to Embodiment 1 will be described. For the purpose of effectively showing the characteristics of the configuration of the first embodiment, the tomographic imaging is first performed after the spot imaging of the subject. First, the subject is placed on the top 2. When the surgeon instructs to move the X-ray tube 3 and FPD 4 through the console 21, the X-ray tube 3 and FPD 4 move to the same position in the body axis direction of the subject M. A portion of the subject M sandwiched between the X-ray tube 3 and the FPD 4 is a region of interest of the subject M that is seen through spot imaging.

  When the surgeon gives an instruction to start spot imaging through the console 21, the X-ray imaging apparatus 1 enters the spot imaging mode. The X-ray tube control unit 6 acquires the parameters for spot imaging from the storage unit 23, and the amplifier array control unit 45 acquires the analog gain for spot imaging from the storage unit 23. Then, a high-contrast perspective image P using relatively strong radiation is generated by the image generation unit 11 and displayed on the display unit 22. Thus, the X-ray tube control unit 6 sets the output of the X-ray tube 3 according to the selection mode selected at the time of imaging, and controls the X-ray tube 3. Further, the amplifier array control unit 45 sets the amplification factor of the amplifier array 44 according to the shooting mode selected at the time of shooting. In other words, when the imaging mode is changed, the X-ray imaging apparatus 1 is changed to one suitable for the imaging mode in which the output of the X-ray tube 3 and the amplification factor of the amplifier array 44 are selected.

  Next, when the surgeon gives an instruction to start tomographic imaging of the subject through the console 21, the X-ray imaging apparatus 1 is switched to the tomographic imaging mode. The X-ray tube control unit 6 acquires tomographic parameters from the storage unit 23, and the amplifier array control unit 45 acquires tomographic analog gains from the storage unit 23. Originally, since relatively weak radiation is used, a fluoroscopic image with low contrast should be acquired. However, since the analog gain acquired by the amplifier array controller 45 is higher than that at the time of spot photographing, the original signal output from the transistor 41t is amplified with a high amplification factor, and a high-contrast perspective image is displayed. Generated by the generation unit 11. The analog gain in the tomographic imaging mode is 10 times as shown by the table T in FIG. The X-ray tube 3 and the FPD 4 acquire 74 fluoroscopic images P1 to P74 while moving in the opposite direction along the body axis direction A of the subject.

  In the fluoroscopic image P amplified at a high amplification factor, a noise component (false image) superimposed on the original signal is reflected. Specifically, as shown in FIG. 6, a granular false image is reflected in each of the fluoroscopic images P. If this is spot photography, this granular false image is not removed. However, in tomography, the false image is removed in the process of superimposing the fluoroscopic images P. Since the false image superimposed on the original signal fluctuates with time, the pattern of the granular false image is completely different among the acquired fluoroscopic images P. Such a granular false image is canceled when the superimposing unit 12 superimposes the fluoroscopic image P, and the influence of noise does not appear in the tomographic image C as shown in FIG. Thus, the tomographic image C with excellent visibility is displayed on the display unit 22. Note that the shaded area in FIG. 6 is a fluoroscopic image of the subject reflected in the fluoroscopic image P.

  In addition, the X-ray tube control unit 6 sets the single X-ray output of the X-ray tube 3 in the tomography mode so that the amplification of the amplifier array 44 is not saturated. If the X-ray output is weak, the original signal output from the X-ray detection element becomes small accordingly. In order to use this weak original signal for generating a tomographic image, it is necessary to amplify the original signal with a higher amplification factor. However, the amplification factor of the amplifier array 44 is limited. That is, if the amplifier array 44 is to be amplified with a high amplification factor, the amplification of the amplifier array 44 is saturated. Then, the amplifier array 44 outputs the same output data to the A / D converter 46 regardless of the strength of the original signal output from the detection element. That is, the fluoroscopic image of the subject M represented by the strength of the original signal output from the X-ray detection element is erased by the amplifier array 44. If such a saturation phenomenon of the amplifier array 44 occurs, the tomographic image C cannot be generated.

  Because of such circumstances, according to the configuration of the first embodiment, imaging is performed with sufficiently strong X-ray output in the tomographic imaging mode. Thereby, the saturation phenomenon of the amplifier array 44 does not occur. Therefore, the fluoroscopic image of the subject M represented by the strength of the original signal output from the detection element can be reliably sent to the image generation unit 11 without being erased by the amplifier array 44.

  Next, an X-ray imaging apparatus 30 according to the second embodiment will be described. FIG. 8 is a functional block diagram of the X-ray tomography apparatus according to the second embodiment. As shown in FIG. 8, the X-ray imaging apparatus according to the second embodiment is similar to the configuration described in the first embodiment. Therefore, the description of each part having the same configuration is omitted as appropriate. The configuration of the second embodiment is different in the manner of movement of the X-ray tube 3 and the FPD 4 and the manner of image processing of the X-ray fluoroscopic image. Since the configuration of the FPD 4 is the same as that of the first embodiment, the description thereof is omitted.

  The synchronous movement mechanism 7 moves the FPD 4 provided at the lower part of the top plate 2 linearly along the body axis direction A of the subject M in synchronization with the linear movement of the X-ray tube 3 described above. The moving direction is the same as the moving direction of the X-ray tube 3. That is, the cone-shaped X-ray beam whose irradiation source position and irradiation direction change as the X-ray tube 3 moves is always received by the entire surface of the X-ray detection surface of the FPD 4. In this way, in one inspection, the FPD 4 acquires, for example, 74 X-ray fluoroscopic images while moving in synchronization with the X-ray tube 3 in the same direction. Specifically, the X-ray tube 3 and the FPD 4 move in the same direction to the position indicated by the alternate long and short dash line through the position indicated by the broken line.

  During the examination, the cone-shaped X-ray beam irradiated by the X-ray tube 3 is always irradiated toward the region of interest of the subject M, and this X-ray irradiation angle is a series of X-ray fluoroscopic images. During shooting, it is always a constant 0 °.

  The X-ray imaging apparatus 30 further includes a same-angle image generation unit 19 that generates a later-described same-angle image at a position where the image generation unit 11 and the overlapping unit 12 are interposed.

  Next, the principle of acquiring a tomographic image of the X-ray imaging apparatus 30 according to the second embodiment will be described. The 74 X-ray fluoroscopic images acquired by the series of imaging are formed by the image generation unit 11 and then sent to the same angle image generation unit 19, where, for example, 50 same angle images are formed. . If these 50 images of the same angle are superposed by the superposing unit 12, a desired X-ray tomographic image can be acquired.

  The operation in the same angle image generation unit 19 will be described. In the same angle image generation unit 19, first, the obtained X-ray fluoroscopic image is divided along the direction orthogonal to the synchronous movement direction of the X-ray tube 3 and the FPD 4 to obtain, for example, 50 strip-shaped images. . Then, among the (74 × 50 = 3,700) strip-like images obtained from the series of X-ray fluoroscopic images, a strip-like image in which the angles of the exposed X-rays are the same is selected. Connect them together to obtain the same angle image. If each of the X-ray fluoroscopic images is divided into 50, 50 images of the same angle are acquired. Although the X-ray beam according to the present invention has a cone shape, a reconstruction method in a well-known X-ray tomography apparatus using a long X-ray beam can be applied through such a process.

The image processing performed by the same angle image generation unit 19 will be described in more detail. 9, 10, and 11 are schematic diagrams illustrating image processing of the same-angle image forming unit for X-ray tomographic images according to the second embodiment. Prior to the description, as shown in FIGS. 9A to 9D, each X-ray fluoroscopic image photographed on the detection surface of the FPD 4 is moved every pitch d. as shown in FIG. 9 (e) ~ Figure _{9 (h), O 1,} O 2, ..., O I, ..., and _{O M (1 ≦ I ≦ M} ). Each time the X-ray tube 3 moves at every pitch d, the X-ray tube 3 emits X-rays intermittently. That is, X-rays are irradiated with a pulse every time it moves every pitch d. The FPD 4 also moves synchronously following the X-ray tube 3.

Specifically, when the X-ray tube 3 first irradiates X-rays at the position shown in FIG. 9A, next, the X is moved at the position shown in FIG. Irradiate the line. The X-ray fluoroscopic image O ₁ (see FIG. 9E) is obtained by detecting the X-rays in FIG. 9A by the FPD 4 and the X-rays are detected by the FPD 4 in FIG. A fluoroscopic image O ₂ (see FIG. 9F) is obtained. Similarly, when the X-ray tube 3 moves at every pitch d, X-rays are irradiated at the position shown in FIG. 9C to the (I-1) th, and the X-rays are converted into FPD 4 in FIG. 9C. Is detected, an X-ray fluoroscopic image O _I (see FIG. 9G) is obtained. Finally, the (M-1) th, 9 is irradiated with X-rays at the position (d), the FIG. 9 (d) X-ray fluoroscopic image by X-rays by FPD4 detects _{O M} (See FIG. 9 (h)). In Example 2, the imaging start position in FIG. 9A is the foot side of the subject M, the imaging end position in FIG. 9D is the head side of the subject M, and FIGS. ), The X-ray tube 3 and the FPD 4 move sequentially along the body axis direction A of the subject M.

By X-ray tube 3 is moved for each pitch d, the X-ray fluoroscopic image _{O 1, O 2, ...,} O I, ..., the _{O M} can be decomposed for each pitch d. Specifically, as shown in the enlarged view of FIG. 9 (i), the projection angle, which is the angle formed between the irradiation axis connecting the X-ray tube 3 and the FPD 4 and the body axis of the subject M, is set to θ for each pitch d. ₁ , θ ₂ ,..., Θ _J ,..., Θ _N−1 , θ _N (1 ≦ J ≦ N). Then, the image decomposed for each pitch d coincides with the strip-like images divided into the same projection angles θ ₁ , θ ₂ ,..., Θ _J ,..., Θ _N−1 , θ _N.

X-ray fluoroscopic image _{O 1} shown in FIG. 9 _{(e), O} 11, _{O 12} for each pitch _{d, ..., O 1J, ...} , O 1 (N-1), is decomposed with _{O 1N,} decomposed The strip-shaped image O ₁₁ is an image obtained by irradiation at the projection angle θ ₁ , and the decomposed strip-shaped image O ₁₂ is an image obtained by irradiation at the projection angle θ _2. The strip-shaped image O _1J is an image obtained by irradiation with the projection angle θ _J , and the finally decomposed strip-shaped image O _1N is an image obtained by irradiation with the projection angle θ _N.

Similarly, as shown in FIG. 9F, the fluoroscopic image O ₂ is decomposed into O ₂₁ , O ₂₂ ,..., O _2J ,..., O _{2 (N−1)} , O _2N for each pitch d. The decomposed strip-shaped image O ₂₁ is an image obtained by irradiation at the projection angle θ ₁ , the decomposed strip-shaped image O ₂₂ is an image obtained by irradiation at the projection angle θ ₂ , and so on. The strip-shaped image O _2J decomposed into ₂ is an image obtained by irradiation at the projection angle θ _J , and the finally-resolved strip-shaped image O _2N is an image obtained by irradiation at the projection angle θ _N. .

The (I1) th, the X-ray fluoroscopic image _{O I} shown in FIG. 9 (g), the pitch d every _{O I1, O I2, ...,} O IJ, ..., O I (N-1) , O _IN and the decomposed strip-shaped image O _I1 is an image obtained by irradiation at the projection angle θ ₁ , and the decomposed strip-shaped image O _I2 is obtained by irradiation at the projection angle θ ₂ . Hereinafter, the strip-shaped image O _IJ similarly decomposed is an image obtained by irradiation at the projection angle θ _J , and the finally-resolved strip-shaped image O _IN is irradiated at the projection angle θ _N. The obtained image is obtained.

Finally, (M1) th, the X-ray fluoroscopic image _{O M} shown in FIG. 9 (h), _O M1 every pitch _{d, O M2, ..., O} MJ, ..., O M _The strip-shaped image O _M1 decomposed and decomposed into _(N-1) and O _MN is an image obtained by irradiation at the projection angle θ ₁ , and the decomposed strip-shaped image O _M2 is irradiated at the projection angle θ ₂ . Thereafter, the strip-shaped image O _MJ similarly decomposed is an image obtained by irradiation at the projection angle θ _J , and the finally-resolved strip-shaped image O _MN is the projection angle θ. An image obtained by irradiation with _N is obtained.

Each image thus decomposed is synthesized for each of the same projection angles θ ₁ , θ ₂ ,..., Θ _J ,..., Θ _N−1 , θ _{N as shown in} FIGS. Each X-ray fluoroscopy as described above images _{O 1, O 2, ...,} O I, ..., O M is decomposed into each pitch d (i.e. the projection angle _{θ 1, θ 2, ...,} θ J, ... (Images divided by θ _N−1 and θ _N ) are shown in FIGS. 10 (a) to 10 (d), FIGS. 10 (f) to 10 (i), and FIGS. 11 (a) to 11. (D), as shown in FIGS. 11 (f) to 11 (i).

For example, in the case of the projection angle θ ₁ , the strip-shaped image O ₁₁ in the X-ray fluoroscopic image O ₁ shown in FIG. 10A and the strip-shaped image in the X-ray fluoroscopic image O ₂ shown in FIG. an image _{O 21,} ..., a strip-shaped image _{O I1} in X-ray fluoroscopic image _{O I} shown in FIG. 10 (c), ..., strip-shaped image O of the X-ray fluoroscopy in the image _{O M} shown in FIG. 10 (d) by combining the _M1, obtaining a same angle image R ₁ in the projection angle theta _1, as shown in FIG. 10 (e).

Similarly, in the case of the projection angle θ ₂ , the strip-like image O ₁₂ in the X-ray fluoroscopic image O ₁ shown in FIG. 10 (f) and the strip in the X-ray fluoroscopic image O ₂ shown in FIG. 10 (g). and Jo image _{O 22,} ..., a strip-shaped image _{O I2} in X-ray fluoroscopic image _{O I} shown in FIG. 10 (h), ..., strip-shaped image of the X-ray fluoroscopy in the image _{O M} that shown in FIG. 10 (i) by combining the O _M2, obtain a same angle image _{R 2} in the projection angle theta ₂ as shown in FIG. 10 (j).

(J-1) First, in the case of the projection angle θ _J , the strip-shaped image O _1J in the X-ray fluoroscopic image O ₁ shown in FIG. 11A and the X-ray fluoroscopy shown in FIG. A strip-shaped image O _2J in the image O ₂ ,..., A strip-shaped image O _IJ in the X-ray fluoroscopic image O _I shown in FIG. 11C, and a X-ray fluoroscopic image O shown in FIG. by combining the strip-shaped image O _MJ in _M, obtain the same angle image R _J of the projection angle theta _J as shown in FIG. 11 (e).

Finally, in the case of the (N−1) th projection angle θ _N , the strip-shaped image O _1N in the X-ray fluoroscopic image O ₁ shown in FIG. A strip-shaped image O _2N in the X-ray fluoroscopic image O ₂ shown in FIG. 11 and a strip-shaped image O _IN in the X-ray fluoroscopic image O _I shown in FIG. by combining the strip-shaped image O _MN in X-ray fluoroscopic image O _M, obtain the same angle image R _N of the projection angle theta _N as shown in FIG. 11 (j).

In summary, the angle image generation unit 19, the same projection angle theta ₁ of each image is _{decomposed, θ 2, ..., θ J} , ..., θ N-1, by combining each theta _N, FIG. 10 (e), FIG. 10 (j), FIG. 11 (e), the projection angle theta _1, as shown in FIG. _{11 (j), θ 2,} ..., θ J, ..., θ N-1, per theta _N same angle image _{R 1, R 2, ...,} obtained R J, ..., and _{R N.}

Superposed section 12, the synthesized same angle image _{R 1, R 2, ...,} R J, ..., to obtain a tomographic image by performing a reconstruction process based on the _{R N.} The reconstruction process may be performed using a well-known filtered back projection (FBP) (also called “filtered back projection method”).

<Operation of X-ray imaging apparatus>
Next, the operation of the X-ray imaging apparatus according to Embodiment 2 will be described. For the purpose of effectively showing the characteristics of the configuration of the first embodiment, the tomographic imaging is first performed after the spot imaging of the subject. First, the subject is placed on the top 2. When the surgeon instructs to move the X-ray tube 3 and FPD 4 through the console 21, the X-ray tube 3 and FPD 4 move to the same position in the body axis direction of the subject M. A portion of the subject M sandwiched between the X-ray tube 3 and the FPD 4 is a region of interest of the subject M that is seen through spot imaging.

  When the surgeon gives an instruction to start spot imaging through the console 21, the X-ray imaging apparatus 1 enters the spot imaging mode. The X-ray tube control unit 6 acquires the parameters for spot imaging from the storage unit 23, and the amplifier array control unit 45 acquires the analog gain for spot imaging from the storage unit 23. Then, a high-contrast perspective image P using relatively strong radiation is generated by the image generation unit 11 and displayed on the display unit 22.

  Next, when the surgeon gives an instruction to start tomographic imaging of the subject through the console 21, the X-ray imaging apparatus 1 is switched to the tomographic imaging mode. The X-ray tube control unit 6 acquires tomographic parameters from the storage unit 23, and the amplifier array control unit 45 acquires tomographic analog gains from the storage unit 23. Originally, since relatively weak radiation is used, a fluoroscopic image with low contrast should be acquired. However, since the analog gain acquired by the amplifier array controller 45 is higher than that at the time of spot photographing, the original signal output from the transistor 41t is amplified with a high amplification factor, and a high-contrast perspective image is displayed. Generated by the generation unit 11. The analog gain in the tomographic imaging mode is 10 times as shown by the table T in FIG. While the X-ray tube 3 and the FPD 4 move at the same speed in the same direction along the body axis direction A of the subject, a plurality of same-angle images are acquired.

  The same-angle image R amplified at a high amplification factor includes a noise component (false image) superimposed on the original signal. Specifically, as shown in FIG. 12, a granular false image is reflected in each of the same-angle images R. If this is spot photography, this granular false image is not removed. However, in tomography, the false image is removed in the process of superimposing the fluoroscopic images P. Since the false image superimposed on the original signal fluctuates over time, the granular false image patterns are all different between the acquired same-angle images R. Such a granular false image is canceled when the superimposing unit 12 superimposes the same-angle image R, and the influence of noise does not appear in the tomographic image C. Thus, the tomographic image C with excellent visibility is displayed on the display unit 22. Note that the shaded area in FIG. 12 is a fluoroscopic image of the subject reflected in the same-angle image R.

  The single X-ray output of the X-ray tube 3 in the tomography mode of the second embodiment is set to have a strength that does not saturate the amplification of the amplifier array 44 as in the configuration of the first embodiment.

  As described above, the X-ray imaging apparatus 1 according to the first embodiment has a configuration in which the spot imaging mode and the tomographic imaging mode can be selected. The FPD 4 provided in the X-ray imaging apparatus 1 according to the first embodiment converts the X-rays into original signals, amplifies them, and outputs them to the image generation unit 11. In the conventional configuration, the original signal is amplified with a predetermined analog gain regardless of the shooting mode. On the other hand, according to the first embodiment, the analog gain in the tomographic imaging mode is larger than the analog gain in the spot imaging mode. Thereby, even if the radiation dose in the tomographic imaging mode is suppressed, the contrast of the fluoroscopic image P acquired in the tomographic imaging mode is high.

  In general, when the analog gain is increased, a noise component superimposed on the original signal is also amplified. When the fluoroscopic image P is acquired by the X-ray imaging apparatus 1 according to the first embodiment, a false image due to a noise component is likely to appear as described with reference to FIG. However, the tomographic image C is acquired by superimposing a plurality of fluoroscopic images P and the same angle image R. When the perspective image P and the same-angle image R are compared, the appearance of the false image is different. Therefore, if these are superimposed, the false image is canceled out. Thus, no false image appears in the finally acquired tomographic image C. The reason for the appearance of the false image between the fluoroscopic images is that the noise component superimposed on the original signal fluctuates with time, and the plurality of fluoroscopic images P are acquired at different timings. Since the strip-like images constituting the same-angle image R are also acquired at different timings, the false image appearing in the strip-like image is also used.

  In addition, spot imaging is performed with a low analog gain as in the past, and the intensity of radiation is sufficiently strong, so that the fluoroscopic image P in spot imaging is clear without showing a false image.

  Further, the single radiation output of the X-ray tube 3 in the tomographic imaging mode is smaller than the single radiation output of the X-ray tube 3 in the spot imaging mode. Thereby, the radiation exposure of the subject M in the tomographic imaging mode can be suppressed to the minimum. Since the tomographic image C is generated by superimposing a plurality of fluoroscopic images P, there is no problem even if some false images appear in each of the fluoroscopic images P.

  Further, according to the above-described configuration, the electric signal is amplified before being digitally converted. Since the electric signal before amplification is analog data, the amplification of the original signal can be performed more faithfully if the amplification is performed while the data is in an analog form. Moreover, once digital processing is performed, data processing after the image generation unit 11 is simplified, and deterioration due to data transfer can be prevented as much as possible.

  Further, according to the above-described configuration, the tomographic image photographing mode and the spot photographing mode can be selected according to the operator's instruction. If the surgeon selects one of the modes, the analog gain is changed accordingly. The surgeon can shoot without worrying about analog gain.

  Further, the single X-ray output of the X-ray tube 3 in the tomographic imaging mode is set by the X-ray tube control unit 6 to such an intensity that the amplification of the amplifier array 44 is not saturated. This configuration clarifies the lower limit of the X-ray output in the tomographic imaging mode. When the X-ray output of the X-ray tube 3 is low, the amplification factor of the amplifier array 44 is increased accordingly. At this time, if the X-ray output is too low, the amplification of the amplifier array 44 is saturated and the brightness of the image becomes unclear. According to the configuration described above, since tomographic imaging is performed with X-rays having sufficient intensity so that such a saturation phenomenon does not occur, a tomographic image C suitable for diagnosis can be acquired.

  The present invention is not limited to the configurations of the above-described embodiments, and the following modifications can be made.

  (1) Although each embodiment described above is a medical device, the present invention can also be applied to industrial and nuclear devices.

  (2) The X-ray referred to in each of the above-described embodiments is an example of radiation in the present invention. Therefore, the present invention can be applied to radiation other than X-rays.

  As described above, the present invention is suitable for a medical radiographic apparatus.

Claims (5)

A radiation source that emits radiation;
Radiation source control means for controlling the output of the radiation source;
Radiation detection means for detecting radiation;
Moving means for moving the radiation source and the radiation detecting means synchronously;
Image generating means for generating a fluoroscopic image based on a detection signal output by the radiation detecting means;
A tomographic image is generated by superimposing a series of fluoroscopic images continuously taken while the radiation source and the radiation detecting means are moved, or a portion of the series of fluoroscopic images in which the same part of the subject is reflected And a superimposing means for
The radiation detection means includes
(A) radiation conversion means for converting the radiation incident on the radiation detection means into an electrical signal;
(B) amplifying means configured to amplify the electric signal with a predetermined amplification factor and to be configured by an analog amplifier ;
(C) A / D conversion means for converting the output of the amplification means, which is analog data, into digital data ,
When the radiation source irradiates the radiation detection means with a single radiation to acquire a single fluoroscopic image as a spot imaging mode and the imaging mode for acquiring the tomographic image as a tomographic imaging mode ,
The amplification factor in the tomographic imaging mode is greater than the amplification factor in the spot imaging mode,
The radiation imaging apparatus according to claim 1, wherein a single radiation output of the radiation source in the tomographic imaging mode is smaller than a single radiation output of the radiation source in the spot imaging mode. A radiation source that emits radiation;
Radiation source control means for controlling the output of the radiation source;
Radiation detection means for detecting radiation;
Moving means for moving the radiation source and the radiation detecting means synchronously;
Image generating means for generating a fluoroscopic image based on a detection signal output by the radiation detecting means;
A tomographic image is generated by superimposing a series of fluoroscopic images continuously taken while the radiation source and the radiation detecting means are moved, or a portion of the series of fluoroscopic images in which the same part of the subject is reflected And a superimposing means for
The radiation detection means includes
(A) radiation conversion means for converting the radiation incident on the radiation detection means into an electrical signal;
(B) amplifying means configured to amplify the electric signal with a predetermined amplification factor and to be configured by an analog amplifier ;
(C) A / D conversion means for converting the output of the amplification means, which is analog data, into digital data;
(D) An imaging mode in which the radiation source irradiates the radiation detecting unit with a single radiation to acquire a single fluoroscopic image is a spot imaging mode, and an imaging mode for acquiring the tomographic image is a tomographic imaging mode. And an amplification factor setting means for setting the amplification factor of the amplification means according to the photographing mode selected at the time of photographing,
The amplification factor setting means sets the amplification factor in the tomographic imaging mode to be larger than the amplification factor in the spot imaging mode,
The radiation source control means has a single radiation output of the radiation source in the tomographic imaging mode that is smaller than a single radiation output of the radiation source in the spot imaging mode and strong enough that the amplification of the amplification means is not saturated. A radiation imaging apparatus, characterized in that the setting is made as described above. The radiographic apparatus according to claim 1 or 2,
The moving means moves the radiation source and the radiation detecting means synchronously in opposite directions,
The radiographic apparatus characterized in that the superimposing means generates a tomographic image by superimposing a series of fluoroscopic images. The radiographic apparatus according to claim 1 or 2,
The moving means is moved synchronously in the same direction as the radiation source and the radiation detecting means,
A series of fluoroscopic images is divided into strips to generate strip-like images, which are joined together while selecting strip-like images with the same angle of radiation applied to the radiation detection means. It further comprises equiangular image generation means for generating a plurality of images,
The radiographic apparatus characterized in that the superimposing means generates a tomographic image by superimposing a series of images at the same angle. The radiation imaging apparatus according to any one of claims 1 to 4 ,
It further comprises an input means for inputting an instruction of the surgeon,
A radiographic apparatus characterized in that the mode of imaging can be changed by an instruction from an operator.

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