JPS61244176A - Digital radiography device - Google Patents

Abstract

PURPOSE:To constitute a digital radiography device having a dynamic range which covers widely the dose of X rays, by connecting two capacitors at least in parallel to a photoelectric transducer and changing storage capacities in accordance with the dose of irradiated X rays. CONSTITUTION:Switches Sc1-Sc128 are opened in case of radiographing with the small dose of X rays. In case of radiographing with a large dose of X rays, switches Sc1-Sc128 are closed to increase capacities of capacitors connected in parallel to photoelectric transducers Pd1-Pd128. Since capacities are increased by connecting capacitors Sc1-Sc128 in this manner, the charged electric charge is not saturated though the dose of incident X rays is large, and the S/N ratio is not damaged by opening switches Sc1-Sc128 though the dose of incident X rays is small.

Description

[Detailed Description of the Invention] [Technical Field of the Invention] The present invention involves irradiating a subject with radiation and transmitting radiation (t)
The present invention relates to a digital radiography device that converts one dose into a digital signal and displays it.

, [Technical background of 1 Ming] Figure 5 shows an image intensifier (hereinafter 1.1.
It is called. ) and a single image tube to detect radiation images. That is, X
II tube 1c x 111tlj control part 21
iIl and irradiates X-rays toward the subject.

The X-rays that have passed through the subject P become a radiation image I.

■, enters 22; 1.1.22 amplifies this radiation image and outputs it as an optical image. This optical image is
3 to an image pickup tube 24 such as a vidicon. Il
The image tube 24 converts this optical image into an analog video signal,
Supplies to A/D converter.

A/D converter 7 converts this analog video signal into a digital video signal and supplies it to image memory 28. The image memory 28 temporarily stores this digital video signal. The digital video signal stored in the image memory is subjected to digital image processing such as digital subtraction in the image processing section 26. This digital image-processed digital signal is supplied to the D/A converter 10. The D/A converter 10 converts this digital signal into an analog signal and displays it on the TV monitor 29.

1.1 as above. Conventional digital radiography apparatuses using a camera and an image pickup tube have the same problems as listed below.

That is, ■1.1. Since it is a vacuum tube with a built-in electron lens, there is a limit to how large the field of view can be made due to its structure. For example, the one commonly used is 12 inches (30 cm
m) Up to about φ, the field of view is circular.

■1.1. Since the input surface is a vacuum tube with a convex or concave shape, distortion such as threading distortion occurs in the output image. Further, even if a uniformly absorbing object is photographed, the density distribution of the output image is not uniform, and so-called shading occurs particularly at the periphery of the image.

■Image tubes have a narrow dynamic range and cannot be used for both fluoroscopy and photography.

[Object of the Invention] The present invention has been made to address the above-mentioned problem, and provides a digital radiography device that has a wide dynamic range, is lightweight and compact, can secure a large field of view, and does not cause distortion in output images. The purpose is to provide

[Summary of the Invention] To achieve this object, the present invention includes (1) a radiation source that irradiates radiation toward a subject, detects the radiation that has passed through the subject, and generates a charge according to the intensity of the radiation; Radiation Il arranged in one or two dimensions
l? ! ! Controls the connection between the air converter, the first capacitor connected in parallel with each of these radiation-electric converters, the second capacitor connected in parallel with these first capacitors, and the first capacitor and the second capacitor. a connection control section; a power supply connected to one terminal of each first capacitor; a switch section provided at each other terminal of each first capacitor and closed in sequence; A digital radio comprising: an output end that outputs a current flowing from a power source; an A/D converter that converts the amount of current generated at the output end into a digital signal; and a display that displays the digital signal. Digital radiography device (3) according to claim 1, characterized in that the output of the radiographic source (2) is switched to increase in conjunction with the connection of the second capacitor (3) ftJ of the second capacitor. Digital radiography apparatus according to claim 1, characterized in that the customer base is larger than the capacity of the first capacitor (4) The connection control section is characterized in that a diode and a Waener diode are connected in parallel. A digital radiography apparatus according to claim 1.

[Embodiment of the Invention] Hereinafter, an embodiment of the present invention will be described with reference to the drawings. FIG. 1 is a block diagram of the same embodiment. In the figure, a high voltage generator 12 generates electric power to drive the X-ray tube 1 based on a control signal from a controller 13.

The power supplied from the high voltage generator 12 to the XSs tube 1 can be roughly divided into two types: power for irradiating weak X-rays to perform fluoroscopy, and power for performing imaging by irradiating strong X-rays. There are different types. And the power for fluoroscopy is 100% compared to the power for photography.
It is less than one-fold. The xII irradiated from the xm tube 1 passes through the subject 2 and enters an imaging plate 3, such as an intensifying screen, which converts X-ray energy into light energy.

xsia incident on the imaging plate 3 is the object 2
This corresponds to the remaining amount that could not be completely absorbed by the Xa image. This incident X-ray is converted into an analog electrical signal according to xsia by a photoelectric conversion element configured as a two-dimensional array in the detection unit 4. The electrically converted analog signal is digitally converted by the A/D converter 7 in time series and stored in the image memory 8. The image memory 8 can store data for one or several images, and the control unit 13
Data is sequentially stored in specific addresses using control signals from the controller. The arithmetic processing section 9 takes out data from the image memory 8, performs arithmetic operations on it, and returns the result to the image memory 8 again. The calculated data in the image memory 8 is transferred to the D/A converter 10.
is converted into an analog signal by This analog signal is displayed on the image monitor 11 as an X-ray image.

Next, a detection unit 4 converts the amount of light emitted by xm into an electrical signal.
will be explained in more detail with reference to the block diagram of FIG.

Element e1 including a photoelectric conversion element that converts light into electricity
, 1 to e 128, 128 are evenly arranged in a two-dimensional array. J lement e1.1~e 128°12
1281 [! It is commonly connected by one common signal line via a switch. The same 128 horizontal elements are also arranged in the vertical direction. That is, there are 128 common signal lines. Each common communication @ line is connected to the input of the multiplexer 6.

The multiplexer 6 receives the control signal GK2 from the control WJ section 13.
One of the 128 common signal lines is sequentially selected and outputted to the A/D converter 7 as an output signal of the detector 4.

In this way, the detection section outputs 128 common signals in time series from the multiplexer 6. Actually 128X12
8, 2048x2048, etc. are used, and the multiplexer 6 outputs a plurality of lines, but for the sake of simplicity, the above-mentioned example will be used for explanation.

Next, a circuit diagram of a portion of the horizontal 128 elements of the detection section 4 is shown in FIG. The anodes of the photoelectric conversion elements Pd1 to Pd128 are connected to the negative electrode of the power source E. Also capacitor c
d1 to Cd128 are photoelectric conversion elements Pd1 to Pd128
are connected in parallel. Furthermore, capacitors cd1 to Cd
Capacitor Ce1~(:,
e128 is connected in parallel with capacitors Cd1 to Cd128 via switches 3c1 to 3c12B. The switches SC1 to SC128 connected in series with the capacitors Qel to Ce12B are opened and closed by a control signal GK3 from the control section 13. For fluoroscopy with a small amount of X-rays, the switches SC1 to SC3 c128 are opened. Also, for shooting with a lot of XIHft, use switch 3.
c1 to 3c128 are used as leapfrogs, and photoelectric conversion elements Pd1 to Pd
This is done by increasing the capacitance of the capacitor connected in parallel to d128. By connecting the capacitors SC1 to SC3c128 in this way, the capacitance becomes large, so even when the amount of incident X-rays is large, the charged charges will not be saturated, and the incident
Even when the dose is small, the S/N ratio will not be impaired if the switches SC1 to Sc12g are opened. capacitor cd1
The capacitance 11Cd of ~Cd128 is the capacitance of a capacitor suitable for fluoroscopic X-rays. In addition, if the capacitance of a capacitor suitable for X-rays in the shadow is CO, the capacitors Ce1 to Ce128
The capacitance 11ce is the effective capacitance including the stray capacitance of the switch and is Ce -co -Cd tonal.

The cathodes of the photoelectric conversion elements Pdl-Pd128 are connected to the inputs of the respective preamplifiers A1-A128, and are also connected to the switches Sd1-5d128 on the photoelectric conversion element side. When this switch is closed, the electric charge of the power supply voltage E is charged to the capacitor. Preamplifier A1-A1
28 inputs the potential of the cathode of each Pd1 to P (1128 at high impedance, amplifies it, and outputs it at low impedance. The outputs of the preamplifiers A1 to A128 are connected to the switches Sal to 3a128 on each output side. ing.

The shift register F is controlled by the switches Sd1 to 5d128 and Sa1 to 5a1 by the control signal CK1 from the control unit 13.
Controls the opening and closing of 28. A signal corresponding to the accumulated charge under the control of this shift register F is outputted to the common signal line in time series.

Next, the operation of this embodiment will be explained. First, in the case of fluoroscopy, the switches 3c1 to Sc128 are in an open state.

Each of the capacitors Cdl to Cd128 is supplied with electric charge by the power supply voltage E, and is charged to the voltage between the terminals. The X-ray tube 1 irradiates the subject with weak continuous X-rays with a tube current of several IIIA. The X-rays transmitted after being attenuated by the object are converted into light by the imaging plate 3 and passed through the diode P.
The light enters d1 to pd128, respectively, and generates charges according to their respective intensities. These charges are transferred to capacitor C
It flows into d1 to Qd128 and cancels out the pre-charged charge. That is, the charge proportional to the intensity of incident X-rays decreases.

During transmission, the switches 3a1 to 3a128 are sequentially closed synchronously by the shift register F. When the switch Sal is closed, a current corresponding to the charge canceled by xi is supplied from the power supply voltage E to the capacitor Cd1.

This current is detected at the output terminal OUT and output to the A/D converter 7. While this switch Sal is turned on, the capacitor Cd1 is charged to the initial state. Next, the switch Sa2 is closed, and xam incident on the diode Pd2 is similarly detected. In this way, while the X-rays are being irradiated, the xam incident on each of the diodes Pd1 to Pd128 is sequentially and synchronously detected.

Further, during photographing, all the switches Sc1 to Sc128 are turned on. And as in fluoroscopy, capacitors cd1~C
Charge is supplied to d128 and the capacitors Ce1 to Ce128 by the power supply voltage E, and the voltage between the terminals is charged. X-ray tube 1 has a strong xI with a tube'FR flow of several hundred 11A! A pulse is emitted to the subject 2. xtm transmitted through object 2
is converted into a current by the diodes Pd1 to Pd128 in the same way as in fluoroscopy. These currents are connected to capacitor cd
1 to Cd128 and Ce1 to Ce128 are canceled. Switches Sa1-3 a128 are sequentially closed only once for one X-ray pulse, unlike in the case of fluoroscopy. And each switch Sa1-3 a12
8 is closed, an output corresponding to the intensity of the incident xwi is obtained at the output end 0LJT, similar to fluoroscopy.

The time during which each of the switches Sa1 to Sa128 is closed is required to be long enough to charge the capacitor to its initial state. Therefore, the capacitance of the capacitor is small and cannot store a large amount of charge. In the case of fluoroscopy, the switching time is shorter than in the case of imaging.

Next, a second embodiment of the detection section 4 will be explained with reference to the circuit diagram shown in FIG. The difference between the first embodiment and the second embodiment is that diodes and Zener diodes are used instead of the switches 3c1 to 3c128 to automatically switch the capacitance. The diodes D1 to D128 and the Zener diodes 71 to 7128 are connected in parallel, respectively. Capacitors Ce1 to Ce128 are connected in series with these. These series connected capacitors Cdl~C
(Connected in parallel to N28. Furthermore, capacitor Cd
1 to Cd128 are connected in parallel to the photoelectric conversion elements Pd1 to Pd128.

The reading operation in the storage mode shown in FIG. 2 is performed as follows.

Once switches Sd1-S (1128 and Sa1-5
After closing a128 and charging the capacitors Cd1 to Cd128, the switch is opened. The voltage between the terminals of the capacitor at this time is 20 (V) when the power supply voltage E is 20 (V).
), :] The voltage between the terminals of the capacitors Ce1 to ce128 becomes 19.3 (V) due to the voltage drop (0.7V) of the diodes D1 to D128. The charges canceled by the incident X-rays are the 20
It starts with a charge of (V). Zener die t-t'Z1~
If one voltage of Z128 (D') is 12 (V), the voltage of capacitors Cd1 to Cd128 is 7°3 (■)
Since the Zener diodes 71 to 7128 do not conduct until the voltage drops to , the voltage of the capacitors Ce1 to ce128 does not drop. However, the voltage of capacitors Cd1 to Cd128 is 7.
3 (V) or less, the voltage between the terminals of the Zener diode is higher than the Zener voltage (then the Zener diodes 71 to 7128 become conductive, and the capacitors Cel to Ce12
Current flows from capacitors Cdl to Cd128 from capacitors Cdl to Cd128, and the voltages of capacitors ce1 to Ce128 drop. In this way, in the second embodiment, when the amount of charge to be canceled increases, the capacitors CC11 to Ce128 are also automatically used. That is, by using the Zener diodes 71 to 7128 and the diodes D1 to [) 128, the capacitance can be automatically switched.

[Effects of the Invention] As described above, according to the present invention, by connecting at least two capacitors in parallel to a photoelectric conversion element and changing the storage capacity in accordance with the amount of X-ray irradiation, a wide range of xams can be supported. It is possible to provide a digital radiography apparatus having a dynamic range of .

[Brief explanation of the drawing]

FIG. 1 is a block diagram of an embodiment of a digital radiography apparatus according to the present invention, FIG. 2 is a block diagram of a detection section of the same embodiment, and FIG. 3 is a circuit diagram of a first embodiment of the elements of the detection section. 4 is a circuit diagram of a second embodiment of the elements of the detection section, and FIG. 5 is a diagram showing the overall configuration of a conventional digital radiography apparatus. 1... X-ray tube, 4...
- Detection parts pd1 to Pd128... photoelectric exchange element, c
dl~C6128... Capacitor C1~Ce128:I capacitor, S dl~
S d128...Switch 3c1 to S c128...Switch, A1 to A1
28... Preamplifier D1-D128... Diode, Z1-Z128.
...Zener diode patent attorney Kensuke Norichika (and one other person) Figure 1 Figure 2 CK. Figure 3 CK3

Claims (4)

[Claims] (1) A radiation source that irradiates radiation toward a subject; a radiation-electric conversion unit arranged in one or two dimensions that detects the radiation that has passed through the subject and generates a charge according to its intensity; A first capacitor connected in parallel with each of these radiation electric conversion units, a second capacitor connected in parallel with these first capacitors, a connection control unit that controls the connection between the first capacitor and the second capacitor, and each A power supply connected to one terminal of the first capacitor, a switch section provided at each other terminal of each first capacitor and closed in sequence, and outputting a current flowing from the power supply to the capacitor when the switch section is closed. What is claimed is: 1. A digital radiography apparatus comprising: an output terminal for generating a signal, an A/D conversion section for converting the amount of current generated at the output terminal into a digital signal, and a display section for displaying the digital signal. 2. Digital radiography apparatus according to claim 1, characterized in that: (2) the output of the radiation source is switched to increase in conjunction with the connection of the second capacitor. (3) The digital radiography apparatus according to claim 1, wherein the capacitance of the second capacitor is larger than the capacitance of the first capacitor. (4) The digital radiography apparatus according to claim 1, wherein the connection control section connects a diode and a Zener diode in parallel.