JP2011194242A - Radiographic apparatus, radiographic system, and program - Google Patents

Abstract

PROBLEM TO BE SOLVED: To obtain an accurate radiographic image with a high S/N ratio by holding the output when radiation is emitted and when radiation is not emitted within the dynamic range of an imaging system corresponding to a plurality of conditions for radiography such as a still image photography mode and a video mode.SOLUTION: The voltage to be applied from voltage control means 6 to a reading circuit unit 3 is adjusted by an arithmetic operation unit 7 so that the maximum or minimum value of an electric signal when the radiation is emitted and the minimum or maximum value of the electric signal when the radiation is not emitted are both held within the dynamic range of the reading circuit unit 3 and an A/D conversion unit 4.

Description

  The present invention relates to a radiation imaging apparatus, a radiation imaging system, and a program suitable for use in, for example, medical diagnosis and industrial nondestructive inspection. In this specification, “radiation” includes electromagnetic waves such as X-rays and γ-rays, α-rays, β-rays, and the like.

  Conventionally, radiation imaging systems installed in hospitals, clinics, etc. include an analog imaging method and a digital imaging method. In the analog imaging method, the patient is irradiated with radiation such as X-rays, and the radiation transmitted through the patient is exposed to the film. In the digital imaging system, radiation transmitted through a patient is converted into an electrical signal to obtain image data as digital data.

Digital imaging methods include CR (Computed Radiography) method and FPD (Flat Panel Detector) method. In the CR method, a radiation image is temporarily stored in a photostimulable phosphor mainly made of BaFBr, and then digital data is obtained by laser beam scanning. In the FPD method, radiation transmitted through a patient is converted into visible light by a scintillator such as Gd ₂ O ₂ S: Tb or CsI: Tl, and converted into an electric signal by a photoelectric conversion element mainly made of an amorphous silicon semiconductor. In addition, some FPD systems use, for example, amorphous selenium as a material that directly converts radiation into an electric signal without using a scintillator. The former is called an indirect type FPD, and the latter is called a direct type FPD.

  In recent years, there has been a growing demand for FPD to capture moving images such as those conventionally performed with an image intensifier (II), such as gastroscopic imaging and intraoperative angiographic imaging. The reason is as follows. I. However, while there are problems such as image distortion in the surroundings, halation problems during intense radiation irradiation, and sensitivity deterioration problems during long-term use, FPD does not have such problems, especially in recent years. This is because it is becoming possible to create it at a relatively low cost. The CR system has been widely used as a device that can be digitized since the 1980's, but it is unsuitable for moving image shooting due to its operating principle. That is, the CR system for still images and the I.D. I. The FPD method having both functions is considered to become the mainstream of digitization in the future medical care. Digitization using the FPD method not only greatly improves the workflow in the hospital and makes it easier to record and print imaging data, but also improves diagnostic efficiency by making full use of advanced image processing technology using a computer. It can be expected to contribute greatly to improvement. From the latter half of 1990, standing and recumbent type radiography apparatuses adopting the FPD method have been put on the market. Recently, X-ray imaging apparatuses capable of taking moving pictures have been proposed and have been put on the market.

  In moving image shooting (perspective shooting), in order to continuously irradiate a patient with long-term radiation compared to still image shooting, it is necessary to consider patient exposure reduction, and in comparison with still image shooting, FPD Higher sensitivity is a challenge. In still image shooting, for example, it is necessary to capture an image of a lung field portion through which radiation easily passes and a mediastinum portion through which radiation does not pass through a single shooting, and the dynamic range of FPD is emphasized as well as sensitivity. Thus, the radiation dose irradiated to acquire one image differs between still image shooting and moving image shooting.

  In Japanese Patent Laid-Open No. 2004-23654 (Patent Document 1), in the moving image shooting mode and the still image shooting mode, the gain (Gf) in the moving image shooting mode and the gain in the still image shooting mode (Gs) set in the readout circuit unit. ) Is disclosed as an example where Gf> Gs. This is because the SNR of the FPD is improved by increasing the gain at the time of moving image in consideration of the random noise generated after the gain setting circuit.

  In Japanese Patent Laid-Open No. 11-331703 (Patent Document 2), since the signal amount varies depending on the shooting mode, the S / N ratio of the detected image is set by effectively using the dynamic range of the A / D converter. For the purpose of improving, a configuration for switching the gain is disclosed. In Patent Document 2, the DSA mode and the fluoroscopic mode have a charge amount that differs by three digits, so that the signal detection amplifier is an integration amplifier and the integration capacitance is switched. It is disclosed that when it is difficult to finely adjust the capacitance value, it is possible to finely adjust the gain with the subsequent stage. In Japanese Patent Application Laid-Open No. 2004-228688, the switching of the capacitance of the integrating amplifier and the switching of the gain of the subsequent stage are described separately in terms of words, but in the sense of changing the switching signal level of the capacitance, it is considered to have the same meaning as the gain switching.

  In general, in moving image shooting using a large area FPD, pixel addition driving is performed to increase the frame rate. In Patent Document 2, the signal delay and waveform distortion of the scanning line driving signal are caused by the resistance and capacitance of the scanning line that drives the switch element. Therefore, the frame rate required for the circulatory system diagnosis system (moving image shooting) cannot be obtained by normal driving, and the frame rate is increased by turning on a plurality of scanning lines at a time (pixel addition driving).

JP 2004-23654 A Japanese Patent Laid-Open No. 11-331703

  Generally, there is a known method of changing the gain in the readout circuit section in an imaging mode in which the radiation dose, that is, the signal amount from the FPD differs depending on the imaging mode such as the moving image mode and the still image mode, or other imaging modes. ing. For example, Patent Document 2 describes that the amount of charge differs by three orders of magnitude between the DSA mode and the fluoroscopic mode, and the following problem occurs when the gain setting is increased.

(1) Since the input offset voltage and input offset current of each operational amplifier in the channel unit are different in the operational amplifier in the readout circuit section, the output in the dark time varies in the channel unit (dark offset variation).
(2) In the radiation imaging circuit unit, the dark current of the sensor, the capacitance (Cgs) between the gate electrode and the source electrode of the switch element, the leakage current between the source and drain of the switch element, and the like vary from pixel to pixel. As a result, the dark output varies in units of pixels (dark offset variation).

In particular, if a high gain is applied so as to compensate for the difference in the 3-digit charge amount as in the DSA mode and the fluoroscopic mode, the relationship between the dark offset variation (Voff) and the signal amount depends on the design of the readout circuit unit. (Vsig) is as follows.
Voff ≒ Vsig
Or
Voff> Vsig

FIG. 10 is a characteristic diagram for explaining the above problem.
As shown in the figure, when a high gain is applied, dark offset variation (Voff) increases, and error pixels that deviate from the dynamic range of the readout circuit unit are generated. Alternatively, even if the dark offset variation (Voff) is within the dynamic range of the readout circuit unit, although not shown, it may exceed the dynamic range of the subsequent AD converter. If this happens, the radiographic image information cannot be obtained correctly.

The above problem does not occur only when the gain is set large.
In the moving image shooting mode used in cardiovascular system diagnosis, the frame rate is increased by using pixel addition driving. Since pixel addition driving drives a plurality of rows of switching elements at a time, the amount of variation in dark offset due to the dark current of the line imaging element and the capacitance (Cgs) between the gate electrode and the source electrode of the switching element is inevitably required. Become bigger. Of course, in pixel addition driving, the signal amount increases in proportion to the number of additions.

  In general, a medical radiation imaging apparatus is required to have a high S / N ratio in order to reduce patient exposure. Therefore, the dark offset output generated in the readout circuit unit and radiation imaging circuit unit is acquired once when radiation is not irradiated (non-irradiation) in the imaging process, and is subtracted from the data during radiation irradiation. Is desirable. Since the output at the time of radiation irradiation is superimposed on the output at the time of non-irradiation, that is, the offset output at the time of darkness, just obtaining the output only at the time of radiation irradiation may not always obtain correct imaging information. In particular, when the gain is set high, the radiation signal output irradiated with radiation may fall within the dynamic range even when the dark offset output is below the dynamic range. It is necessary that the output of both irradiation and non-irradiation is within the dynamic range of the system.

  The present invention has been made in view of the above-described problems, and in response to a plurality of shooting conditions such as a still image shooting mode and a moving image shooting mode, outputs during irradiation and non-irradiation output of the imaging system Fit within the range. Accordingly, it is an object to provide a radiation imaging apparatus, a radiation imaging system, and a program that can obtain a radiographic image that is accurate and has a high S / N ratio.

  In the radiation imaging apparatus of the present invention, a pixel including a radiation detection element that converts radiation into an electrical signal and a switching element that selects the radiation detection element and transfers the electrical signal is provided on a substrate. A plurality of radiation detectors arranged in parallel, a signal detector for detecting radiation, a signal reader for reading out an electrical signal from the radiation detector, and an A / D converter for A / D converting the electrical signal from the signal reader. A voltage control means for controlling a voltage applied to the signal reading means, and a maximum value of the electrical signal at the time of radiation irradiation and a minimum value of the electrical signal at the time of non-radiation, or the electrical power at the time of radiation irradiation. Both the minimum value of the signal and the maximum value of the electrical signal when no radiation is irradiated fall within the dynamic range of the signal reading means and the A / D conversion means. And a voltage adjusting means for adjusting the applied voltage, and a radiation imaging apparatus capable of imaging in a plurality of imaging modes, wherein the signal reading means receives an electrical signal from the radiation detecting means at an inverting input terminal. A non-inverting input terminal for inputting an electric signal from the first operational amplifier via a first operational amplifier that is input and supplied with a first reference voltage to the non-inverting input terminal and a capacitor connected in series to the inverting input terminal And a second operational amplifier to which a second reference voltage is supplied. The voltage adjusting means adjusts the second reference voltage as the applied voltage in accordance with each photographing mode.

  The radiation imaging system of the present invention includes a radiation source that irradiates a subject with radiation, and the radiation imaging apparatus that captures radiation from the subject.

  According to another aspect of the radiation imaging system of the present invention, a radiation source that irradiates a subject with radiation, a radiation detection element that converts radiation into an electrical signal, and switching for selecting the radiation detection element and transferring the electrical signal A plurality of pixels having a plurality of elements arranged side by side on the substrate, a radiation detection means for detecting radiation from a subject, a signal readout means for reading out an electrical signal from the radiation detection means, and the signal readout A radiation imaging apparatus for imaging radiation from a subject, comprising: A / D conversion means for A / D converting the electrical signal from the means; and voltage control means for controlling a voltage applied to the signal reading means; The maximum value of the electrical signal at the time of radiation irradiation from the radiation source and the minimum value of the electrical signal at the time of radiation non-irradiation, or the electrical signal at the time of radiation irradiation The signal reading means from the voltage control means so that both the minimum value and the maximum value of the electrical signal when no radiation is irradiated fall within the dynamic range of the signal reading means and the A / D conversion means. A radiation imaging system capable of imaging in a plurality of imaging modes, wherein the signal readout unit inputs an electrical signal from the radiation detection unit to an inverting input terminal. The first operational amplifier to which the first reference voltage is supplied to the non-inverting input terminal and the electric signal from the first operational amplifier are input to the non-inverting input terminal via the capacitor connected in series to the inverting input terminal. A second operational amplifier to which a second reference voltage is supplied; and the voltage adjusting unit adjusts the second reference voltage as the applied voltage in accordance with each photographing mode.

  The program of the present invention selects a radiation detection element that converts radiation into an electrical signal at the time of radiation irradiation and the radiation detection element and transfers the electrical signal to a computer according to each imaging mode of a plurality of imaging modes. And a maximum value of an electrical signal detected by a radiation detection device arranged in parallel on the substrate and read out by a signal reading means, and detected by the radiation detection device when no radiation is applied. The minimum value of the electric signal read by the signal reading means or the minimum value of the electric signal detected by the radiation detecting device when irradiated with radiation and read by the signal reading means and detected by the radiation detecting device when no radiation is irradiated The maximum value of the electric signal read by the signal reading means is the signal reading. The electric signal from the radiation detection device is input to the inverting input terminal so that it falls within the dynamic range of the A / D conversion means for A / D converting the electric signal from the signal reading means and the signal reading means. A first operational amplifier to which a first reference voltage is supplied to a terminal and an electric signal from the first operational amplifier through a capacitor connected in series to an inverting input terminal and a second reference to a non-inverting input terminal The second reference voltage as the applied voltage to the signal reading means from the voltage control means for controlling the applied voltage to the signal reading means including the second operational amplifier to which the voltage is supplied corresponds to each photographing mode. To execute the adjustment procedure.

  According to the present invention, in response to a plurality of shooting conditions such as a still image shooting mode and a moving image shooting mode, the output during irradiation and non-irradiation falls within the dynamic range of the imaging system, and is accurate. A radiographic image with a high / N ratio can be obtained. In the present invention, it is possible to freely perform both still image shooting and moving image shooting with a single radiation imaging apparatus. Especially in moving images, even if gain is increased or pixel addition driving is performed, A good image can be obtained. One radiation beam imaging device enables convenient radiation imaging capable of capturing both moving images and still images.

1 is a schematic configuration diagram illustrating an X-ray imaging system according to a first embodiment of the present invention. It is a timing chart which shows an example of various operation pulses in the X-ray imaging system of a 1st embodiment. FIG. 2 is a characteristic diagram schematically showing analog outputs of various AMPs that are constituent elements of the readout circuit section in FIG. 1. It is a circuit block diagram which shows the structural example of each circuit which produces | generates each reference potential in the power supply control part in 1st this embodiment. It is a circuit block diagram which shows the other structural example of each circuit which produces | generates each reference potential in the power supply control part in 1st this embodiment. It is the schematic for demonstrating the effect by 1st Embodiment. It is a flowchart which shows the X-ray imaging method by 1st Embodiment. It is a circuit block diagram for demonstrating 3rd Embodiment. It is a schematic diagram which shows the internal structure of a personal user terminal device. It is a characteristic view for demonstrating the conventional problem.

  DESCRIPTION OF EXEMPLARY EMBODIMENTS Hereinafter, preferred embodiments in which a radiation imaging apparatus of the invention is applied to X-ray imaging will be described in detail with reference to the drawings. Note that the present invention is not limited to imaging of X-rays, and can be applied to imaging of electromagnetic waves such as γ rays, and various types of radiation such as α rays and β rays.

(First embodiment)
FIG. 1 is a schematic configuration diagram showing an X-ray imaging system according to a first embodiment of the present invention.
The X-ray imaging system includes an X-ray source 10 that irradiates a subject with X-rays, an X-ray imaging device 20 that captures X-rays transmitted through (or reflected from) the subject, and the X-ray source 10 and the X-ray imaging device 20. And a control unit (not shown) for controlling the driving of the motor.

  In the X-ray imaging apparatus 20, reference numeral 1 denotes an X-ray detection circuit unit for detecting X-rays from a subject. The X-ray detection circuit unit 1 includes a plurality of pixels arranged in a two-dimensional array, each having an X-ray detection element that converts radiation into an electrical signal and a switching element for selecting the X-ray detection element. Has been configured. Here, in this embodiment, the X-ray detection element has a photoelectric conversion element, and the switching element uses a thin film transistor (TFT).

In FIG. 1, this photoelectric conversion element uses a PIN photodiode (PD) whose main component is an amorphous silicon semiconductor. The photoelectric conversion element is provided with a phosphor as a wavelength converter (scintillator) that receives radiation such as X-rays and converts the light into a wavelength band that can be sensed by the photoelectric conversion element. This phosphor is composed mainly of any of Gd ₂ O ₂ S, Gd ₂ O ₃ , and CsI. Here, as the X-ray detection element, instead of the photoelectric conversion element, a direct conversion type X-ray detection element that absorbs radiation and directly converts it into an electric signal without using a wavelength converter may be used. good. As this direct conversion type X-ray detection element, for example, one material selected from lead iodide, mercury iodide, selenium, cadmium telluride, gallium arsenide, gallium phosphorus, zinc sulfide, and silicon is used. Configured as the main component.

  Further, instead of the PIN photodiode, an MIS sensor mainly composed of an amorphous silicon semiconductor may be used as the photoelectric conversion element. This MIS type sensor is configured by sequentially laminating a first metal layer, an insulating layer made of amorphous silicon nitride, a semiconductor layer made of hydrogenated amorphous silicon, an impurity semiconductor layer, and a second metal layer. Here, the first metal layer functions as a lower electrode, and the second metal layer functions as an upper electrode. The second metal layer is constituted by a transparent conductive layer, a metal layer, and a stacked structure thereof, and is disposed on at least a part of the impurity semiconductor layer. The insulating layer prevents passage of electrons and holes generated by the semiconductor layer. The impurity semiconductor layer is for blocking one injection of holes or electrons. For example, the impurity semiconductor layer is an impurity semiconductor layer doped with N-type for blocking hole injection.

  In the MIS type sensor, in the refresh mode, an electric field is applied to the semiconductor layer in a direction in which hole carriers are guided from the semiconductor layer to the second metal layer. On the other hand, in the photoelectric conversion mode, carriers generated by light incident on the semiconductor layer remain in the semiconductor layer, and an electric field is applied in a direction in which electrons are guided to the second metal layer. Then, holes accumulated in the semiconductor layer or electrons guided to the second metal layer by the photoelectric conversion mode are detected as an optical signal.

  A TFT is generally configured using amorphous silicon as a main material. The electrical signal photoelectrically converted in the PD is transferred to the readout circuit unit 3 by the TFT. PD and TFT form a pixel as a pair, and are arranged in a two-dimensional array. For example, when obtaining a fluoroscopic image of the human chest, the imaging region may be about 43 cm × 43 cm, and the number of pixels is about 7.2 million pixels when configured with a 160 μm pitch.

Reference numeral 2 denotes a drive circuit unit 2 for driving the X-ray detection circuit unit 1.
Reference numeral 3 denotes a readout circuit unit that reads an electric signal from the X-ray detection circuit unit 1. In the readout circuit unit 3, the first stage unit, the second stage unit, and the third stage unit are respectively cascaded by operational amplifiers (AMPs 18a to 18c). In the AMPs 18a to 18c, the reference potential is biased to the non-inverting input terminal (+ terminal). In order to increase the S / N ratio (hereinafter referred to as SNR) of the X-ray imaging apparatus, the AMP 18a is specifically designed to reduce noise. AMP1 is an integrating circuit in which capacitors C19a and 19b are connected between the inverting input terminal (− terminal) and the output terminal, and the capacitance value C1 and the capacitance value (C1 + C2) are selected based on the GAIN1 signal for the integration capacitance. Done.

  In still image shooting, since it is necessary to obtain image information with a single image of the lung field that is easy to transmit X-rays and the mediastinum and abdomen that are relatively difficult to transmit X-rays, It is necessary to irradiate large X-rays and project the shadows on one image. Therefore, the integration capacity is increased by switching on GAIN1. Conversely, in the case of moving image shooting, X-rays are irradiated for a longer time than in still image shooting, so the X-ray dose per frame (acquisition of one piece of image information) is reduced from the viewpoint of reducing patient exposure. There is a need. Therefore, in the case of moving image shooting, in order to increase the output voltage value of the AMP 18a, the integration capacity is set small.

  A capacitor C19c is connected between the output terminal of AMP1 and the inverting input terminal (−terminal) of AMP2, and capacitors 19d and 10e are connected between the inverting input terminal (−terminal) of AMP2 and the output terminal. Yes. The capacitance value C4 and the capacitance value (C4 + C5) are selected for the integral capacitance based on the signal GAIN2. When the capacitance value C4 is selected, the output of the AMP 18b is multiplied by (−C4 / C3) the output of the AMP 18a. When the capacitance value (C4 + C5) is selected, the output of the AMP 18b is multiplied by (− (C4 + C5) / C3) the output of the AMP 18a.

  Next, the output of the AMP 18b is sampled and held in the capacitor 19f by the sample and hold signal SH. Next, the signal sampled and held by the capacitor 19f is output to the A / D converter 4 via the AMP 18c by the multiplexer signal MPX.

  Reference numeral 4 denotes an A / D conversion unit that performs A / D conversion on the electrical signal from the readout circuit unit 3. In FIG. 1, for simplification of explanation, one pixel is shown, and one A / D conversion unit 4 is shown for one pixel. However, in reality, one A / D conversion unit 4 is used. Many pixels are A / D converted. For example, signal processing is performed in parallel up to the switch to which the MPX signal is input, and serial conversion is performed by the MPX signal. RC, DRC, and RES are reset signals for integration capacitors of AMP 18a, AMP 18b, and AMP 18c, respectively.

  Reference numeral 5 denotes a timing generation unit that applies operation pulses to the drive circuit unit 2 and the readout circuit unit 3. Here, an example of a timing chart of various operation pulses given from the timing generation unit 5 to the drive circuit unit 2 and the readout circuit unit 3 is shown in FIG. FIG. 2 shows the timing when four pixels are serially converted. Here, the timing of the operation pulses to the TFT and the operation pulses of RC, DRC, SH, RES, and MPX with respect to the X-ray input is shown. The output of the A / D conversion unit 4 is also shown in the last stage of FIG. As the operation pulse to the TFT, SIN, OE, and OPV signals are given from the timing generation unit 5 to the drive circuit unit 2, and are given from the drive circuit unit 2 to the TFT based on these signals.

  Reference numeral 6 denotes a voltage control unit that controls applied voltages to the X-ray detection circuit unit 1, the drive circuit unit 2, and the readout circuit unit 3. The voltage control unit 6 includes a constant voltage application unit 16 that applies a predetermined constant voltage to the X-ray detection circuit unit 1 and the drive circuit unit 2, and a variable voltage application unit 17 that applies a variable voltage to the readout circuit unit 3. It is prepared for. The constant voltage application unit 16 applies Vs to the PD of the X-ray detection circuit unit 1 and applies Vss and Vcom to the drive circuit unit 2. The variable voltage application unit 17 includes VREF1 at the non-inverting input terminal (+ terminal) of the AMP 18a of the readout circuit unit 3, VREF2 at the non-inverting input terminal (+ terminal) of the AMP 18b, and a non-inverting input terminal (+ terminal) of the AMP 18c. VREF3 is applied to each. Here, a power source to which VREF1 is applied is a variable power source 11, a power source to which VREF2 is applied is a variable power source 12, and a power source to which VREF3 is applied is a variable power source 13. The power source for applying the power source voltage to the readout circuit unit 3 is a variable power source 14.

  Reference numeral 7 denotes a calculation unit (voltage adjustment unit) that adjusts the timing generation unit 5 and the voltage control unit 6. The imaging data digitally converted by the A / D conversion unit 4 is input to the calculation unit 7. The calculation unit includes a memory 15, and A / D converted imaging data or calculated data is stored in the memory 15. The arithmetic unit 7 is configured to include, for example, a DSP and the like together with the memory 15 and desirably processes various data at high speed. In the present embodiment, the calculation unit 7 is described as a component of the X-ray imaging apparatus 20, but the calculation unit 7 may be an external component of the X-ray imaging apparatus 20 included in the X-ray imaging system.

  When the arithmetic unit 7 determines that at least one of the minimum value of the dark output and the maximum value of the bright output deviates from the range of the dynamic range of the readout circuit unit 3 and the A / D converter 4, the power source control Feedback to part 6. That is, the computing unit 7 computes an appropriate value that falls within the dynamic range based on the deviation, and feeds back the computed value to the voltage control means 6. As feedback items, the reference potentials VREF1 to VREF1-3 connected to the non-inverting input terminals (+ terminals) of the AMPs 18a, AMP18b, and AMP18c provided in the readout circuit unit 3 or the power supply voltage of the readout circuit unit are changed. That is, the maximum value of the electrical signal (output during light) at the time of radiation irradiation of the X-ray detection circuit unit 1 and the minimum value of the electrical signal (output at dark time) at the time of non-radiation are determined by the dynamic range of the readout circuit unit 3 When it does not fall within the dynamic range of the A / D converter 4, the reference potential and the power supply voltage are changed.

FIG. 3 is a characteristic diagram schematically showing analog outputs (denoted as AMP1, 2, 3) of AMP18a, AMP18b, and AMP18c, which are constituent elements of the readout circuit unit 3 in FIG. The vertical axis direction in FIG. 3 is intended for voltage, but the horizontal axis direction is not intended for time.
The non-inverting input terminals (+ terminals) of the AMPs 18a, AMP18b, and AMP18c are biased with the reference potential VREF1, the reference potential VREF2, and the reference potential VREF3, respectively. FIG. 3 shows outputs of reference potential VREF1 = 3V, reference potential VREF2 = 2V, and reference potential VREF3 = 3V. FIG. 3 shows an output example in the direction in which a current flows into the input of the readout circuit unit 3 due to X-ray detection (TFT is on).

The output of the AMP 18a is obtained by transferring the signal charge photoelectrically converted by the PD of the X-ray detection circuit unit 1, and integrating the charge to output a voltage. At this time, the non-inverting input terminal (+ terminal) of the AMP 18a is biased to the reference potential VREF1 (3V), and the output of the AMP 18a is output based on the VREF1 level (3V). When the capacitance between the inverting terminal and the output terminal of the AMP 18a is Cf and the signal charge is Qsig, the output voltage of the AMP 18a is as follows.
Output of AMP 18a = VREF1- (Qsig / Cf)
However, Cf is switched by GAIN1.

The non-inverting input terminal (+ terminal) of the AMP 18b is biased to the reference potential VREF2 (2V), and the output of the AMP 18b is output based on the VREF2 level (2V). The amplification factor (GAIN) of AMP2 is connected to the capacitance (C3) of the capacitor 19c connected to the output terminal of the AMP 18a and the inverting input terminal (− terminal) of the AMP 18b, and to the inverting input terminal (− terminal) and output terminal of the AMP 18b. It is determined by the capacitances (C4, C5) of the capacitors 19d, 19e.
Output of AMP18b = VREF2 + (Qsig / Cf) × GAIN
It becomes.

  GAIN is switched by GAIN2 and becomes C3 / C4 or C3 / (C4 + C5). In FIG. 3, GAIN is displayed as 1 time. The output of the AMP 18b is sampled and held by the capacitor 19f.

The non-inverting input terminal (+ terminal) of the AMP 19c is biased with the reference power source VREF3 (3V), and after resetting the capacitor 19g (capacitance: C7) in advance, the signal is transferred with the MPX signal. From the law of conservation of charge:
C7 · (AMP3-VREF3) = C6 · (VREF3-AMP2)
If C6 and C7 have the same capacity,
Output of AMP19c = 2 · VREF3-VREF2-Qsig / Cf
It becomes. In FIG. 3, C6 and C7 are displayed with the same capacity.

The output levels of the AMPs 18a, 18b, 18c (shown as AMP1, 2, 3 in the table) are as shown in Table 1.
That is, the output in the example of FIG. 3 is determined by the dark potential output level by the reference potential VREF2 of the AMP 18b and the reference potential VREF3 of the AMP3. The light output level is determined.

FIG. 4 is a circuit configuration diagram showing a configuration example of each circuit that generates each reference potential in the power supply control unit 6 in the present embodiment. In this example, the configuration of the variable power supplies 11 to 13 in the power supply controller 6 is shown.
The control signal from the arithmetic unit 7 is converted by the decoder 21 and selects the resistance value input to the inverting terminal of the operational amplifier 23. In the illustrated example, 16 kinds of voltages are selected by four types of resistors 22.

FIG. 5 is a circuit diagram showing another configuration example of each circuit for generating each reference potential in the power supply controller 6 in the embodiment of the present invention. In this example, the configuration of the variable power supplies 11 to 13 in the power supply control unit 6 is shown as in FIG.
In the case of selecting a resistance value as shown in FIG. 5, it is also conceivable that noise is included in the reference potential due to the thermal noise of the resistor 22 and as a result the image quality deteriorates. In FIG. 5, the original power source (Va to Vd) that forms the potential of the reference power source is input to the operational amplifier 34 via the primary low-pass filter formed by the resistor 31 and the capacitor 32, and is output to the analog multiplexer 35. . If the operational amplifier 34 used here is a low noise type, the reference potential has less noise and the image quality is improved. In FIG. 5, four types of reference potentials Va to Vd can be selected by a selection signal from the calculation unit.

FIG. 6 is a schematic diagram for explaining the effect of this embodiment.
In FIG. 6, the left side shows dark output and explicit output (X-ray output) when the image is captured in the still image capturing mode, that is, when a low gain is set. On the right side, there are shown dark output and explicit output (X-ray output) when shooting in the moving image shooting mode, that is, when a high gain is set. In the figure, what is described in a staircase pattern is intended to delimit pixels, and an arrow indicates a signal output amount by X-rays. In FIG. 6, the output for 10 pixels is shown.

  6 differs from FIG. 9 in that the minimum value of the dark offset variation (Voff) is within the dynamic range of the readout circuit unit 3 by changing the dark output level. . In this case, the light output (X-ray output) and the maximum value in the moving image photographing mode are within the dynamic range of the readout circuit unit 3. That is, both the dark output and the bright output are within the dynamic range of the readout circuit unit 3, and the radiation image can be correctly read without causing an error as shown in FIG.

  Note that the output of the AMP 18c (final stage) of the readout circuit unit 3 shown in FIG. 1 is described in the downward direction (− direction) along with the increase in X-rays. On the other hand, in FIG. 6, the increase in X-rays is described in the upward direction (+ direction), which is the opposite direction to FIG. However, whether the signal increases in the upward direction or in the downward direction is related to the signal carrier handled by the radiation imaging apparatus and the readout circuit unit 3 and the setting of each circuit. It does not influence. That is, in FIG. 6, the description has been made so that the minimum value of the dark output and the maximum value of the X-ray output fall within the dynamic range range. For example, the output increases with an increase in X-rays like the AMP 18c of FIG. It is not appropriate when it falls. In this case, it is desirable to describe the maximum value instead of the minimum value in the dark output, and similarly to the minimum value instead of the maximum value in the light output (X-ray output). The same applies to the following second and third embodiments.

  In the present embodiment, as shown in Table 1, the dark output level of the AMP 18c of the readout circuit unit 3 is related to the reference potential VREF2 supplied to the AMP 18b and the reference potential VREF3 supplied to the AMP3. Therefore, the reference potentials adjusted and changed so as to be within the dynamic range of the readout circuit unit 3 are VREF2 and VREF3.

Next, the X-ray imaging method according to the present embodiment will be described.
FIG. 7 is a flowchart showing the X-ray imaging method according to the present embodiment.
First, in the X-ray imaging system, the X-ray imaging apparatus sets an imaging mode and imaging conditions based on control by the control unit (step S1).
The shooting modes are largely classified into a still image shooting mode and a moving image shooting mode. In each mode, there are a pixel addition mode in which two or more pixels are added in the horizontal direction or the vertical direction, a pixel average mode in which the added pixels are averaged, and a pixel non-addition mode in which no pixels are added. The number of pixel additions may be two-pixel addition, three-pixel addition, four-pixel addition, or more. Examples of mode settings include a two-pixel addition mode for moving image shooting, a pixel non-addition mode for still image shooting, and a pixel non-addition mode for moving image shooting. There is also a mixed mode of moving image shooting mode and still image shooting mode.

  In the mixed mode, the moving image and the still image are repeated. In general, while monitoring a patient's continuous fluoroscopic image, a still image is obtained by turning on a photographing switch at the time of the desired still image. The shooting conditions include the parameters shown in Table 2 as an example. However, the shooting conditions are not limited to these, and a number of shooting conditions may be provided according to each mode.

Subsequently, based on the control by the control unit, the calculation unit 7 acquires the dark output in the non-irradiation state of X-rays and stores it in the memory 15 (step S2).
Subsequently, based on the control by the control unit, the calculation unit 7 determines whether or not the minimum value of the dark output is within the dynamic range of the readout circuit unit 3 and the A / D conversion unit 4 (step S3). ).

In step S3, when it is determined that the minimum value of the dark output is not within the dynamic range of at least one of the readout circuit unit 3 and the A / D conversion unit 4, the calculation unit 7 determines each reference of the power supply control unit 6. The potential is changed (step S5). In this case, the reference potential is changed so as to increase the DC level of the dark output.
On the other hand, when it is determined in step S3 that the minimum value of the dark output is within the dynamic range of both the readout circuit unit 3 and the A / D conversion unit 4, the process proceeds to step S4.

  Subsequently, in the X-ray imaging system, the X-ray source 10 irradiates the X-ray detection circuit unit 1 with X-rays based on the control by the control unit, and the calculation unit 7 acquires the light output and stores it in the memory 15. (Step S4). At this time, it is desirable that the subject (subject: usually a patient) is not in the X-ray imaging apparatus, and the entire surface of the X-ray imaging apparatus is irradiated with substantially uniform X-rays.

  Subsequently, based on control by the control unit, the calculation unit 7 determines whether or not the maximum value of the light output is within the dynamic range of the readout circuit unit 3 and the A / D conversion unit 4 (step S6). ).

In step S 6, when it is determined that the maximum value of the light output is not within the dynamic range of at least one of the readout circuit unit 3 and the A / D conversion unit 4, the calculation unit 7 determines each reference of the power supply control unit 6. The potential is changed (step S8). In this case, the reference potential is changed so as to reduce the direct output DC level. In this case, there is a risk that the minimum value of the dark output acquired previously in order to reduce the dark output deviates from the dynamic range. Therefore, after step S8, the process returns to step S2, and the minimum value of the dark output is checked again.
On the other hand, when it is determined in step S6 that the maximum value of the light output is within the dynamic range of both the readout circuit unit 3 and the A / D conversion unit 4, the process proceeds to step S7.

Subsequently, in the X-ray imaging system, the X-ray source 10 irradiates the subject located in the X-ray detection circuit unit 1 with X-rays and performs X-ray imaging based on control by the control unit (step S7).
When the X-ray imaging is video recording, X-ray imaging for a long time is performed as compared with still image shooting. On the other hand, in the case of still image shooting, one or several X-ray irradiations are performed. In the mixed mode of moving image and still image, the operation is repeated as moving image → still image → moving image →.

  In the above description, the process of changing the reference potential is included in both the light output and the dark output, but there may be a case where there is no appropriate reference potential that satisfies both the light output and the dark output. . In that case, since there is no optimal solution in the imaging conditions, another imaging condition (for example, reducing the X-ray dose, reducing the gain, etc.) is selected, and the above process is performed again. This portion is omitted in the flowchart of FIG.

  As described above, in the present embodiment, the calculation unit 7 is connected to the power control unit 6 so that both the light-time output and the dark-time output are within the dynamic range of the readout circuit unit 3 and the A / D conversion unit 4. The signal reference potential to the readout circuit unit 3 is changed. Here, the dynamic range of the A / D conversion unit 4 is preferably the same as the dynamic range of the readout circuit unit 3, but they may be different from each other.

  In the present embodiment, the circuit configuration in which the reference potential is connected to the non-inverting input terminal of the operational amplifier as illustrated in FIG. 1 has been described as an example. However, the circuit configuration is not particularly limited to this, Any circuit configuration example that changes the DC level of the dark output or the bright output may be used.

Further, although the charge integration type operational amplifier configuration is shown in FIG. 1, a voltage amplification type operational amplifier configuration such as an inverting amplifier circuit or a non-inverting amplifier circuit using a resistor may be used.
In FIG. 1, the timing generation unit 2, the A / D conversion unit 4, the power supply control unit 6, and the calculation unit 7 are described separately. For example, these components are mounted on one printed circuit board. It may be done.

  As described above, according to the present embodiment, in response to a plurality of photographing conditions such as the still image photographing mode and the moving image photographing mode, the output at the time of X-ray irradiation and the non-irradiation of the X-ray is output. Therefore, it is possible to obtain an accurate X-ray image having a high S / N ratio. In the present embodiment, it is possible to freely perform both still image shooting and moving image shooting with one X-ray imaging device 20, and in particular for moving images, even if the gain is increased, pixel addition driving is performed. However, a good image can be obtained. One X-ray imaging apparatus 20 can provide an easy-to-use X-ray imaging system capable of capturing both moving images and still images.

(Second Embodiment)
In the present embodiment, a case will be described in which VREF1 of the variable power supply 11 is changed instead of VREF2 and VREF3 in the X-ray imaging system of FIG. That is, the calculation unit 7 determines that the maximum value or minimum value of the output when X-ray irradiation is performed and the minimum value or maximum value of the output when X-ray irradiation is not performed are both of the readout circuit unit 3 and the A / D conversion unit 4. VREF1 is changed so as to be within the range of the dynamic range.

  In the case of the configuration shown in FIG. 1, the reference potential VREF1 becomes the output of the AMP 18a as shown in FIG. 3, but the final dark output level does not appear as a function form of the VREF1 as shown in the AMP 18c. For example, in the moving image shooting mode, it is necessary to reduce the X-ray dose from the viewpoint of reducing the exposure to the patient, and in order to secure the signal amount per frame, the integration capacity of the AMP 18a of the readout circuit unit 3 is reduced. , Increase the gain of the AMP 18b.

  On the other hand, in the case of the still image shooting mode, the integral capacity of the AMP 18a is increased in order to capture images under more X-ray conditions than in the moving image capturing mode in order to capture shadows in the lung field, mediastinum, and abdomen. Thus, the dark output variation and the light output (X-ray output) in the AMP 18a differ depending on the difference in integration capacitance, and the dynamic range of the readout circuit unit 3 or the A / D conversion unit 4 may be out of range. In such a case, VREF1, which is the reference potential of AMP1, is changed / adjusted.

  However, in the configuration example shown in FIG. 1, in normal operation, the output of AMP1 is output through AMP 18b and AMP 18c, and therefore, the output of only AMP 18a cannot be input to the calculation unit for calculation.

  In the present embodiment, for example, when the output cannot be obtained directly by the A / D converter 4 as in the case of the AMP 18a, the optimum VREF1 is calculated in advance. That is, when the X-ray imaging apparatus 20 is shipped from the factory or installed at the customer, the optimum VREF1 is calculated by inspection under each of several imaging conditions, and the value of VREF1 is calculated at the time of actual imaging. Adapt. Thus, the optimum value of VREF1 calculated by the inspection is stored in the memory 15 of the calculation unit 7.

  Table 2 below shows an example of imaging conditions. Imaging conditions include an X-ray irradiation condition and a condition of the X-ray imaging apparatus 20. The former includes tube voltage (V), tube current (I), X-ray pulse width (T), distance from the X-ray tube to the X-ray detection circuit unit 1 (D), and the thickness of the filter that cuts low energy components. (Th) and material. The latter includes the bias power supply (Vs) of the X-ray imaging apparatus 20, the on-voltage (Vcom) of the TFT, the initial stage integration capacitance (Cf) of the readout circuit unit 3, the gain (G) of the readout circuit unit 3, and the like.

  As described above, according to the present embodiment, the value of VREF1 corresponding to the imaging condition is determined in advance by inspection. As a result, the output during X-ray irradiation and non-X-ray irradiation is kept within the dynamic range of the imaging system in correspondence with a plurality of shooting conditions such as still image shooting mode and moving image shooting mode, and the S / An X-ray image having a high N ratio can be obtained. In the present embodiment, it is possible to freely perform both still image shooting and moving image shooting with one X-ray imaging device 20, and in particular for moving images, even if the gain is increased, pixel addition driving is performed. However, a good image can be obtained. One X-ray imaging apparatus 20 can provide an easy-to-use X-ray imaging system capable of capturing both moving images and still images.

(Third embodiment)
In this embodiment, in the X-ray imaging system of FIG. 1, the minimum value of the dark output and the maximum value of the bright output are functioned with each imaging parameter constituting the imaging condition.
The minimum value of the dark time output and the maximum value of the light time output which are converted into functions are stored in the memory 15 of the calculation unit 7. The calculation unit 7 then sets the reference potential so that the minimum value of the dark output and the maximum value of the light output that are functionalized fall within the dynamic range of the readout circuit unit 3 and the A / D conversion unit 4. To change.

  The light output during X-ray irradiation (X-ray output) and the dark output without X-ray irradiation are functions of the imaging parameters described above. For example, the X-ray output is proportional to the X-ray dose such as the tube current (I) and the X-ray pulse width (T) and the gain (G) of the readout circuit unit. The X-ray output is inversely proportional to Cf and inversely proportional to the square of D. The X-ray output has parameters that can be expressed as such a simple function and parameters that are not so complex. The latter can be approximated as a function by conducting an experiment.

  On the other hand, it goes without saying that the dark output is irrelevant to the X-ray condition. However, if it does not become a simple function form with respect to Vs and gain G, an experiment is performed to approximately display the function. Keep it.

  Furthermore, the function approximation for all the pixels constituting the X-ray detection circuit unit 1 of the X-ray imaging apparatus 20 has many millions of pixels. Therefore, it cannot be said to be realistic. Of these pixels, only the pixel with the smallest dark output or the largest light output needs to be functionalized, so that the output of all pixels in between is within the dynamic range. Fits in. However, it should be noted that depending on the shooting conditions, the pixel with the darkest output or the pixel with the brightest output is not always the same pixel when the shooting condition is changed. Necessary pixel coordinate information may be stored in the memory 15 depending on the photographing conditions.

  As described above, according to the present embodiment, the minimum value of the dark output and the maximum value of the bright output are functioned by each shooting parameter constituting the shooting condition. As a result, corresponding to a plurality of shooting conditions such as still image shooting mode and moving image shooting mode, the output at the time of radiation irradiation and at the time of non-irradiation is kept within the dynamic range of the imaging system, and the S / N ratio is accurate. A high radiographic image can be obtained. In the present embodiment, it is possible to freely perform both still image shooting and moving image shooting with one X-ray imaging device 20, and in particular for moving images, even if the gain is increased, pixel addition driving is performed. However, a good image can be obtained. One X-ray imaging apparatus 20 can provide an easy-to-use X-ray imaging system capable of capturing both moving images and still images.

(Fourth embodiment)
In the present embodiment, a case where the power supply voltage of the readout circuit unit 3 is changed by the variable power supply 14 in the X-ray imaging system of FIG. 1 will be described.

  The dynamic range of the readout circuit unit 3 depends on the power supply voltage and is determined by the circuit configuration of the input stage and the output stage. The dynamic range of the operational amplifier is about 1 V smaller than the power supply voltage. Recently, a so-called rail-to-rail operational amplifier capable of amplitude in the range of the power supply voltage has been designed. By using this method, the same dynamic range as the power supply voltage is secured.

FIG. 8 is a circuit configuration diagram for explaining the present embodiment, and is a circuit configuration diagram of the variable power source 14 for changing the power source voltage of the readout circuit unit 7 by a control signal from the arithmetic unit 7.
The difference from FIG. 4 is that the MOS transistor 41 is connected to the operational amplifier 23. In the circuit configuration as shown in FIG. 4 for forming the reference potential, almost no current flows because the input impedance of the operational amplifier is generally high. However, since the power source of the readout circuit unit 3 is composed of a large number of operational amplifiers, it is necessary to have a capability of flowing a large amount of current. The MOS transistor 41 is a so-called power MOS transistor. The voltage source connected to the drain terminal (D) needs to have a capability of flowing a necessary current to the readout circuit unit 3. In the circuit configuration example shown in FIG. 8, 16 power supply voltages can be selected.

  The calculation unit 7 can change the dynamic range of the read circuit unit 3 by controlling the variable power supply 14 of the power supply control unit 6 to change the power supply voltage of the read circuit unit 3. In the present embodiment, the power supply voltage of the readout circuit unit 3 is changed so that the minimum value of the dark output and the maximum value of the bright output (X-ray output) are within the dynamic range.

  As described above, according to the present embodiment, the dynamic range is changed by adjusting the power supply voltage of the readout circuit unit 3. As a result, corresponding to a plurality of shooting conditions such as still image shooting mode and moving image shooting mode, the output at the time of radiation irradiation and at the time of non-irradiation is kept within the dynamic range of the imaging system, and the S / N ratio is accurate. A high radiographic image can be obtained. In the present embodiment, it is possible to freely perform both still image shooting and moving image shooting with one X-ray imaging device 20, and in particular for moving images, even if the gain is increased, pixel addition driving is performed. However, a good image can be obtained. One X-ray imaging apparatus 20 can provide an easy-to-use X-ray imaging system capable of capturing both moving images and still images.

(Other embodiments to which the present invention is applied)
The functions of the constituent elements (the control unit and the calculation unit (excluding the memory 15), etc.) constituting the X-ray imaging system according to the above-described embodiment are performed by the operation of a program stored in the RAM or ROM of the computer. realizable. Similarly, each step of the X-ray imaging method (steps S1 to S8 in FIG. 7) can be realized by operating a program stored in a RAM or ROM of a computer. This program and a computer-readable storage medium storing the program are included in the present invention.

  Specifically, the program is recorded on a recording medium such as a CD-ROM or provided to a computer via various transmission media. As a recording medium for recording the program, besides a CD-ROM, a flexible disk, a hard disk, a magnetic tape, a magneto-optical disk, a nonvolatile memory card, or the like can be used. On the other hand, as the program transmission medium, a communication medium in a computer network system for propagating and supplying program information as a carrier wave can be used. Here, the computer network is a WAN such as a LAN or the Internet, a wireless communication network, or the like, and the communication medium is a wired line such as an optical fiber or a wireless line.

  Further, the program included in the present invention is not limited to the one in which the functions of the above-described embodiments are realized by the computer executing the supplied program. For example, such a program is also included in the present invention when the function of the above-described embodiment is realized in cooperation with an OS (operating system) or other application software running on the computer. Further, when all or part of the processing of the supplied program is performed by the function expansion board or function expansion unit of the computer and the functions of the above-described embodiment are realized, the program is also included in the present invention.

  For example, FIG. 9 is a schematic diagram illustrating an internal configuration of a personal user terminal device. In FIG. 9, reference numeral 1200 denotes a personal computer (PC) having a CPU 1201. The PC 1200 executes device control software stored in the ROM 1202 or the hard disk (HD) 1211 or supplied from the flexible disk drive (FD) 1212. The PC 1200 generally controls each device connected to the system bus 1204.

  By the program stored in the CPU 1201, the ROM 1202 or the hard disk (HD) 1211 of the PC 1200, the procedure of steps S1 to S8 in FIG.

  A RAM 1203 functions as a main memory, work area, and the like for the CPU 1201. A keyboard controller (KBC) 1205 controls instruction input from a keyboard (KB) 1209, a device (not shown), or the like.

  Reference numeral 1206 denotes a CRT controller (CRTC), which controls display on a CRT display (CRT) 1210. Reference numeral 1207 denotes a disk controller (DKC). The DKC 1207 controls access to a hard disk (HD) 1211 and a flexible disk (FD) 1212 that store a boot program, a plurality of applications, an editing file, a user file, a network management program, and the like. Here, the boot program is a startup program: a program for starting execution (operation) of hardware and software of a personal computer.

  Reference numeral 1208 denotes a network interface card (NIC) that exchanges data bidirectionally with a network printer, another network device, or another PC via the LAN 1220.

DESCRIPTION OF SYMBOLS 1 X-ray detection circuit part 2 Drive circuit part 3 Reading circuit part 4 A / D conversion part 5 Timing generation part 6 Power supply control part 7 Calculation part 10 X-ray source 20 X-ray imaging devices 11-14 Variable power supply 15 Memory 16 Constant voltage Application unit 17 Variable voltage application unit

Claims (8)

A plurality of pixels each having a radiation detection element that converts radiation into an electrical signal and a switching element that selects the radiation detection element and transfers the electrical signal are arranged in parallel on the substrate. Radiation detecting means for detecting the signal, signal reading means for reading out an electrical signal from the radiation detecting means,
A / D conversion means for A / D converting an electric signal from the signal reading means;
Voltage control means for controlling the voltage applied to the signal reading means;
The maximum value of the electrical signal at the time of radiation irradiation and the minimum value of the electrical signal at the time of non-irradiation, or the minimum value of the electrical signal at the time of radiation irradiation and the maximum value of the electrical signal at the time of non-radiation However, voltage adjusting means for adjusting the applied voltage so that both fall within the dynamic range of the signal reading means and the A / D conversion means,
A radiation imaging apparatus capable of imaging in a plurality of imaging modes,
The signal reading means includes a first operational amplifier in which an electric signal from the radiation detection means is input to an inverting input terminal and a first reference voltage is supplied to a non-inverting input terminal, and a capacitor connected in series to the inverting input terminal Including a second operational amplifier that receives an electric signal from the first operational amplifier via the first operational amplifier and is supplied with a second reference voltage to a non-inverting input terminal;
The radiation imaging apparatus, wherein the voltage adjusting unit adjusts the second reference voltage as the applied voltage in accordance with each imaging mode. The maximum value of the electrical signal at the time of radiation irradiation and the minimum value of the electrical signal at the time of non-irradiation, or the minimum value of the electrical signal at the time of radiation irradiation and the maximum of the electrical signal at the time of non-radiation. The values are pre-functioned in relation to the shooting conditions so that both fall within the range of the dynamic range,
2. The radiation imaging apparatus according to claim 1, wherein the voltage adjustment unit selects the second reference voltage suitable for the imaging condition of the imaging at the time of imaging based on the function.   The parameters of the imaging conditions are the tube voltage of the radiation source, the tube current of the radiation source, the irradiation pulse time, the distance between the radiation source and the radiation detection means, the thickness and material of the radiation filter, and applied to the radiation detection means. The radiation imaging apparatus according to claim 2, wherein the voltage condition is an integration capacitance of the signal reading unit, and a detection gain of the signal reading unit. A radiation source for irradiating the subject with radiation,
A radiation imaging system comprising: the radiation imaging apparatus according to claim 1, which captures radiation from a subject. A radiation source for irradiating the subject with radiation,
A plurality of pixels having a radiation detection element for converting radiation into an electrical signal and a switching element for selecting the radiation detection element and transferring the electrical signal are arranged in parallel on a substrate, Radiation detection means for detecting radiation from the radiation detection means, signal readout means for reading out electrical signals from the radiation detection means, A / D conversion means for A / D converting the electrical signals from the signal readout means, and the signals A radiation control apparatus that controls the voltage applied to the reading means, and that captures radiation from the subject;
The maximum value of the electrical signal at the time of radiation irradiation from the radiation source and the minimum value of the electrical signal at the time of no radiation irradiation, or the minimum value of the electrical signal at the time of radiation irradiation and the electricity at the time of no radiation irradiation. Voltage adjustment for adjusting the applied voltage from the voltage control means to the signal reading means so that the maximum value of the signal falls within the dynamic range of the signal reading means and the A / D conversion means. Means,
A radiation imaging system capable of imaging in a plurality of imaging modes,
The signal reading means includes a first operational amplifier in which an electric signal from the radiation detection means is input to an inverting input terminal and a first reference voltage is supplied to a non-inverting input terminal, and a capacitor connected in series to the inverting input terminal Including a second operational amplifier that receives an electric signal from the first operational amplifier via the first operational amplifier and is supplied with a second reference voltage to a non-inverting input terminal;
The radiation adjusting system, wherein the voltage adjusting unit adjusts the second reference voltage as the applied voltage in accordance with each imaging mode. The maximum value of the electrical signal at the time of radiation irradiation and the minimum value of the electrical signal at the time of non-irradiation, or the minimum value of the electrical signal at the time of radiation irradiation and the maximum of the electrical signal at the time of non-radiation. The values are pre-functioned in relation to the shooting conditions so that both fall within the range of the dynamic range,
The radiation imaging system according to claim 5, wherein the voltage adjusting unit selects the second reference voltage suitable for the imaging condition of the imaging at the time of imaging based on the function.   The parameters of the imaging conditions are the tube voltage of the radiation source, the tube current of the radiation source, the irradiation pulse time, the distance between the radiation source and the radiation detection means, the thickness and material of the radiation filter, and applied to the radiation detection means. The radiation imaging system according to claim 6, wherein the voltage condition is an integration capacity of the signal reading unit, and a detection gain of the signal reading unit.   The computer has a radiation detection element that converts radiation into an electrical signal at the time of radiation irradiation and a switching element that selects the radiation detection element and transfers the electrical signal according to each imaging mode of the plurality of imaging modes. Thus, the maximum value of the electrical signal detected by the radiation detecting device arranged in parallel on the substrate and read by the signal reading means, and detected by the radiation detecting device when no radiation is irradiated and read by the signal reading means. The minimum value of the electrical signal, or the minimum value of the electrical signal detected by the radiation detection device during radiation irradiation and read out by the signal reading means, and detected by the radiation detection device during radiation non-irradiation and read out by the signal readout means. The maximum value of the electric signal is the signal reading means and the signal reading. The electrical signal from the radiation detection device is input to the inverting input terminal and the first signal is input to the non-inverting input terminal so that the electrical signal from the stage is within the dynamic range of the A / D conversion means for A / D converting. An electric signal from the first operational amplifier is input via a first operational amplifier to which a reference voltage is supplied and a capacitor connected in series to the inverting input terminal, and the second reference voltage is supplied to the non-inverting input terminal. A procedure for adjusting the second reference voltage as the applied voltage to the signal reading means from a voltage control means for controlling the applied voltage to the signal reading means including a second operational amplifier according to each photographing mode. A program to be executed.

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