JP2014219248A - Radiation image detection device and operation method of the same - Google Patents

Abstract

PROBLEM TO BE SOLVED: To sample a dose signal at an appropriate sampling period in accordance with the dose of radiation.SOLUTION: An electronic cassette 12 comprises a detection panel 30 in which pixels 40 accumulating signal charge expressing an image signal by receiving an X-ray are two-dimensionally arrayed. In the detection panel 30, detection pixels 40b are provided that detect the X-ray. A signal processing circuit 51 periodically samples a dose signal according to a dose per unit time of the X-ray from the detection pixels 40b. A period appropriateness determination part 70 determines whether the sampling period SP of the dose signal to be set by a period setting part 71 is appropriate with respect to the dose. When it is determined by the period appropriateness determination part 70 that the sampling period SP is inappropriate, the period setting part 71 changes the sampling period SP.

Description

  The present invention relates to a radiological image detection apparatus having an AEC function for stopping irradiation of radiation when a cumulative dose of radiation reaches a target dose, and an operating method thereof.

  In the medical field, X-ray imaging systems using radiation, such as X-rays, are known. The X-ray imaging system includes an X-ray generation apparatus that generates X-rays and an X-ray imaging apparatus that captures X-ray images. The X-ray generator has an X-ray source that irradiates an X-ray toward a subject (patient), and a radiation source control device that controls driving of the X-ray source. The X-ray imaging apparatus performs an X-ray image detection apparatus that detects an X-ray image based on X-rays transmitted through a subject, an imaging control apparatus that controls driving of the X-ray image detection apparatus, and stores and displays X-ray images. Has a console.

  The X-ray image detection apparatus includes a detection panel (also referred to as a flat panel detector (FPD)) that detects an X-ray image as an electrical signal, and a control board. The detection panel is formed with an imaging region in which pixels that accumulate signal charges in response to X-rays are two-dimensionally arranged. The control board includes a signal processing circuit that converts signal charges read from the pixels in units of rows into image signals that form an X-ray image.

  The X-ray image detection apparatus includes an X-ray detection unit that detects X-rays transmitted through the subject and an output of the X-ray detection unit in order to obtain an X-ray image having an appropriate image quality while suppressing exposure to the subject. The dose sampling unit that periodically samples the dose signal representing the X-ray dose per unit time (X-ray intensity) and the accumulated dose of X-rays, which is the integrated value of the dose signal, has reached the target dose. Some of them include an AEC (Automatic Exposure Control) unit for determining whether or not (see, for example, Patent Document 1). When the AEC unit determines that the accumulated dose has reached the target dose, the X-ray irradiation by the X-ray source is stopped. The accumulated dose is determined by a tube current irradiation time product (so-called mAs value) which is a product of the X-ray irradiation time and the tube current given to the X-ray source. Irradiation conditions such as irradiation time and tube current have approximate recommended values depending on the subject's imaging area, sex, age, etc., such as the chest and head, but the X-ray transmittance changes depending on individual differences such as the physique of the subject. AEC is performed in order to obtain an X-ray image with more appropriate image quality.

  In Patent Document 1, a pixel in a detection panel is used as an X-ray detection unit, and a signal processing circuit is used as a dose sampling unit. More specifically, the signal processing circuit reads the charge accumulated in the parasitic capacitance generated between the pixel and the signal line for reading the signal charge from the pixel to the signal processing circuit as a dose signal at a constant sampling period. . The cumulative dose is obtained by integrating the dose signals obtained by each sampling. Patent Document 1 describes that the recommended value of the sampling period of the dose signal is a value between about 0.1 ms to 10 ms.

JP 2000-139890 A

  Incidentally, noise is constantly generated in an electric circuit such as a dose sampling unit. For this reason, in addition to the signal component according to the X-ray dose, the dose signal includes a noise component due to stationary noise in the dose sampling unit. If the S / N ratio, which is the ratio between the signal component of the dose signal and the noise component, is too low, the accuracy of determining whether or not the accumulated dose of the AEC has reached the target dose cannot be maintained. Must be maintained. Since the magnitude of the stationary noise in the dose sampling unit is constant and does not change greatly depending on the dose, the S / N ratio of the dose signal increases as the signal component increases. In order to increase the signal component of the dose signal and increase the S / N ratio, the dose signal sampling period by the dose sampling unit corresponding to the period during which the output of the X-ray detection unit is integrated may be lengthened.

  When the X-ray dose transmitted through the subject is relatively low, such as when the tube current value set for the X-ray source is relatively low or the subject's body thickness is relatively thick, the signal component of the dose signal And the S / N ratio also decreases. Therefore, in such a case, it is necessary to lengthen the sampling period of the dose signal to increase the S / N ratio of the dose signal and to ensure the accuracy of the determination of the AEC unit.

  Conversely, when the X-ray dose is relatively high, a dose signal with a sufficiently high S / N ratio can be obtained even if the sampling period is somewhat short. In AEC, for example, when predicting the time when the cumulative dose reaches the target dose based on the time change of the cumulative dose obtained by integrating the dose signals, the time when the cumulative dose reaches the target dose is predicted as the sampling period is short. The number of times of performing the calculation increases, and the accuracy of the prediction calculation increases accordingly. Therefore, if a high S / N ratio is ensured, the sampling period should be short. Thus, there is an optimum value corresponding to the X-ray dose in the sampling period of the dose signal.

  Although the recommended value of the sampling period of the dose signal is described in Patent Document 1, it is not described how to set the sampling period in specific cases. Therefore, it is unnecessary to set a short sampling period when the X-ray dose is relatively low and a long sampling period needs to be set, or conversely a relatively high X-ray dose and a short sampling period. In some cases, a long sampling period was set.

  The present invention has been made in view of the above problems, and an object of the present invention is to provide a radiological image detection apparatus capable of sampling a dose signal at an appropriate sampling period corresponding to a radiation dose and an operation method thereof.

  In order to achieve the above object, the present invention provides a detection panel having an imaging region in which pixels that receive radiation emitted from a radiation generator and store signal charges representing an image signal are two-dimensionally arranged, and radiation. The radiation detection unit to detect, the dose sampling unit to periodically sample the dose signal according to the dose per unit time of radiation based on the output of the radiation detection unit, and the cumulative dose of radiation based on the dose signal An AEC unit that determines whether or not a dose has been reached, a cycle suitability determination unit that determines whether or not the sampling period of the dose signal by the dose sampling unit is appropriate for the dose, and a cycle suitability determination unit based on the dose signal And a cycle setting unit that changes the sampling cycle when it is determined that the sampling cycle is inappropriate.

  The cycle suitability determining unit repeatedly executes the sampling cycle suitability determination until it determines that the sampling cycle is appropriate.

  The cycle setting unit changes the sampling cycle stepwise from the initial value. More specifically, the cycle setting unit lengthens the sampling cycle stepwise or shortens it stepwise.

  The cycle suitability determining unit compares the magnitude relationship between the signal value of the dose signal and a preset threshold value to determine the suitability of the sampling cycle. More specifically, when the sampling period is increased stepwise, the first threshold is set as the threshold, and the cycle suitability determining unit determines that the sampling period is less than the first threshold when the signal value of the dose signal is below the first threshold. Judged inappropriate. On the other hand, when the sampling cycle is shortened step by step, the second threshold is set as the threshold, and the cycle suitability determination unit determines that the sampling cycle is inappropriate when the signal value of the dose signal exceeds the second threshold. To do.

  The periodic suitability determination unit compares the magnitude relationship between the signal value of the dose signal of the radiation detection unit and the threshold value in the subject region irradiated with the radiation transmitted through the subject in the imaging region.

  A plurality of radiation detection units are arranged so as to be evenly distributed over the entire surface of the imaging region. The periodic suitability determination unit recognizes a subject region irradiated with radiation that has passed through the subject out of the imaging region based on the dose signals of the plurality of radiation detection units, and the dose signal of the radiation detection unit in the subject region The magnitude relationship between the signal value and the threshold value is compared.

  It is preferable to include a mode setting unit that sets either the first change mode in which the sampling cycle is lengthened in steps or the second change mode in which the sampling period is shortened in steps. The mode setting unit includes at least one of tube current, tube voltage, subject body thickness information, and imaging region set in the radiation generator, and is set for each imaging and reaches the radiation dose to the imaging region. The first and second change modes are set based on the shooting conditions for determining the first and second change modes. The mode setting unit sets the first change mode in the case of an imaging condition in which the arrival dose is relatively high, and conversely sets the second change mode in the case of an imaging condition in which the arrival dose is relatively low.

  A setting condition changing unit for changing a setting condition when the cycle setting unit sets a sampling cycle, an initial value of the sampling cycle, a change step for defining an interval of the sampling cycle to be changed once, and an upper limit value of the sampling cycle Or it is preferable to provide the setting condition change part which changes the setting conditions containing at least 1 among a lower limit. The setting condition changing unit includes at least one of tube current, tube voltage, subject body thickness information, and imaging region set in the radiation generator, and is set for each imaging and reaches the imaging region. An initial value of the sampling period is set based on the imaging conditions for determining the dose.

  The setting condition changing unit changes the changing step according to the determination result of the periodic suitability determining unit.

  The AEC unit predicts the time at which the accumulated dose reaches the target dose based on the dose signal obtained by sampling multiple times. The setting condition changing unit sets the upper limit value of the sampling period to a value shorter than ½ of the recommended irradiation time of radiation.

  It is preferable to provide a change mode for changing the sampling period and a fixed mode for not changing the sampling period.

  In addition, the present invention provides a detection panel having an imaging region in which pixels that receive radiation emitted from a radiation generation device and store signal charges representing an image signal are two-dimensionally arranged, a radiation detection unit that detects radiation, and A dose sampling unit that periodically samples a dose signal according to the dose per unit time of radiation based on the output of the radiation detection unit, and whether the cumulative dose of radiation has reached the target dose based on the dose signal An AEC unit that determines whether a sampling period of the dose signal by the dose sampling unit is appropriate for the dose based on the dose signal, and a cycle setting unit that sets the sampling cycle In the operation method of the radiological image detection apparatus provided, when the sampling cycle is determined to be inappropriate by the cycle suitability determination unit, sampling is performed by the cycle setting unit To change the period.

  According to the present invention, when the sampling period of the dose signal is determined to be inappropriate, the sampling period is changed, so that the dose signal can be sampled at an appropriate sampling period corresponding to the radiation dose.

1 is a schematic diagram of an X-ray imaging system. It is a figure which shows an imaging condition table. It is an external appearance perspective view which shows an electronic cassette. It is a block diagram which shows the internal structure of an electronic cassette. It is a figure which shows the example of arrangement | positioning of a detection pixel. It is a figure which shows transition of operation | movement of a detection panel, a control part, and an AEC part. It is a block diagram which shows the internal structure of an AEC part. It is a graph explaining the extrapolation calculation for calculating the remaining time until the accumulated dose of X-rays reaches a target dose. It is a figure which shows the relationship between the sampling period of a dose signal, and the charge accumulation amount of integral amplifier. It is a figure which shows the relationship between the sampling period of a dose signal, and the charge accumulation amount of integral amplifier. It is a block diagram which shows the internal structure of a control part. It is a figure which shows the example of the suitability determination of a sampling period. It is a figure which shows 1st Embodiment which lengthens a sampling period in steps. It is a flowchart which shows transition of operation | movement of an electronic cassette. It is a flowchart which shows the transition of the operation | movement of the suitability determination of a sampling period. It is a figure which shows 2nd Embodiment which shortens a sampling period in steps. It is a figure which shows the example of the suitability determination of the sampling period in 2nd Embodiment. It is a figure which shows 4th Embodiment which switches the 1st change mode which lengthens the sampling period of (A) in steps, and the 2nd change mode which shortens the sampling period of (B) in steps during dose sampling operation | movement. . It is a figure which shows changing a sampling period according to remaining time.

[First embodiment]
In FIG. 1, the X-ray imaging system 2 includes an X-ray imaging apparatus 10 and an X-ray generation apparatus 11. The X-ray imaging apparatus 10 is a portable X-ray image detection apparatus that detects X-rays transmitted through a subject (patient) H and outputs an X-ray image, and controls driving of the electronic cassette 12. The radiography control apparatus 13 is configured to include an X-ray image storage and display process console 14. The X-ray generator 11 includes an X-ray source 15 that irradiates the subject H with X-rays, and a source control device 16 that controls driving of the X-ray source 15.

  The electronic cassette 12 is detachably attached to the holder 17a of the standing position photographing stand 17 for photographing the subject H in a standing position and the holder 18a of the lying position photographing stand 18 for photographing the subject H in a standing position. It is done. The electronic cassette 12 is held by the holders 17 a and 18 a of the imaging tables 17 and 18 with the front surface 34 a (see FIG. 3) facing the X-ray source 15. The subject H is positioned by an operator such as a radiologist so that the imaging region is located between the X-ray source 15 and the electronic cassette 12. The X-ray source 15 can be set in a desired direction and position by a radiation source moving device (not shown), and the X-ray source 15 is shared by the standing position imaging stand 17 and the supine position imaging stand 18.

  The console 14 includes an input device 14a such as a keyboard and a mouse, a display 14b, a storage device 14c, and the like. The input device 14a accepts various operation instructions from an operator such as shooting conditions. The display 14b displays an X-ray image detected by the electronic cassette 12 and various operation screens such as an imaging condition input screen. The storage device 14c is, for example, a hard disk drive, and stores X-ray images from the electronic cassette 12 and various information necessary for X-ray imaging. The X-ray image can be stored in an image storage server (not shown) connected to the console 14 via a network.

  The console 14 accepts an input of an examination order including information such as a photographing part such as the sex, age, head, chest, abdomen, hand, and finger of the subject H, and displays the examination order on the display 14b. The examination order is input from an external system (not shown) that manages subject information and examination information related to radiation examination, such as a hospital information system (HIS) and a radiation information system (RIS). The The inspection order can be manually input by an operator via the input device 14a.

  The X-ray source 15 includes an X-ray tube 15a that generates X-rays, and an irradiation field limiter (also referred to as a collimator) 15b that limits the irradiation field of the X-rays generated from the X-ray tube 15a to the subject H. . The X-ray tube 15a is provided with a filament, a target, a grid electrode, etc. (all not shown). A voltage (tube voltage) is applied between the cathode filament and the anode target. The filament generates thermoelectrons according to the tube voltage. The generated thermoelectrons are emitted toward the target. The target emits X-rays when thermal electrons from the filament collide. The grid electrode is disposed between the filament and the target, and changes the flow rate (tube current) of thermoelectrons from the filament to the target in accordance with the applied voltage.

  The irradiation field limiter 15b is configured such that four lead plates that shield X-rays are arranged on each side of a rectangle, and a rectangular irradiation opening that transmits X-rays is formed at the center. By changing the position of each lead plate, the size of the irradiation opening changes, and thereby the X-ray irradiation field is changed.

  The radiation source control device 16 controls the tube voltage, the tube current, and the X-ray irradiation time by controlling the operation of the voltage generator 16a that generates the tube voltage and the voltage applied to the grid electrode. And a communication unit 16c for exchanging synchronization signals with the electronic cassette 12 via the imaging control device 13.

  The synchronization signal is a signal for synchronizing the X-ray irradiation start / stop timing of the X-ray source 15 and the imaging start / stop timing of the electronic cassette 12, and includes an irradiation start request signal, an irradiation permission signal, and an irradiation stop signal. There is. The irradiation start request signal is a signal for inquiring whether or not to start X-ray irradiation, and is transmitted from the communication unit 16c to the electronic cassette 12 before the X-ray irradiation is started. The irradiation permission signal is transmitted from the electronic cassette 12 toward the communication unit 16c when preparation for receiving X-ray irradiation is completed in the electronic cassette 12. The irradiation stop signal is transmitted from the imaging control device 13 toward the communication unit 16c when it is determined by the AEC function of the electronic cassette 12 that the accumulated dose of X-rays has reached the target dose. The communication unit 16c receives the irradiation permission signal and the irradiation stop signal, and inputs these signals to the control unit 16b.

  The radiation source control device 16 includes an operation panel 19 for an operator to input X-ray irradiation conditions, and an irradiation switch 20 for instructing start of warm-up and X-ray irradiation to the X-ray source 15. It is connected.

  The irradiation conditions input through the operation panel 19 include tube voltage, tube current, and X-ray irradiation time. The control unit 16b controls the tube voltage generated by the voltage generation unit 16a, the voltage applied to the grid, and the time for applying the tube voltage, that is, the X-ray irradiation time, based on the irradiation conditions input from the operation panel 19. .

  The irradiation switch 20 is a two-stage pressing type. The irradiation switch 20 generates a warm-up instruction signal when pressed to the first stage (half-pressed), and generates an irradiation start instruction signal when pressed to the second stage (fully pressed). These signals are input to the control unit 16b.

  When the warm-up instruction signal is input, the control unit 16b operates the voltage generation unit 16a to start warming up the X-ray source 15. Specifically, a predetermined voltage is applied to the filament to preheat the filament, and at the same time, rotation of the target is started. At this time, a voltage is applied to the grid electrode such that the thermoelectrons generated by the filament reach the target and no X-rays are emitted. Warm-up is completed when preheating of the filament is completed and the target reaches a specified rotational speed. When the warm-up instruction signal is input to the control unit 16b, the communication unit 16c transmits an irradiation start request signal.

  When the irradiation permission signal from the electronic cassette 12 is received by the communication unit 16c and the irradiation start instruction signal is input, the control unit 16b operates the voltage generation unit 16a to irradiate the X-ray by the X-ray source 15. Let it begin. Specifically, after applying a tube voltage set in the irradiation condition to the target, a voltage corresponding to the tube current set in the irradiation condition is applied to the grid electrode.

  When the communication unit 16c receives an irradiation stop signal from the imaging control device 13, the control unit 16b operates the voltage generation unit 16a to stop the X-ray irradiation by the X-ray source 15. Specifically, the voltage applied to the grid electrode is switched to the voltage applied during warm-up, the voltage application to the target is subsequently stopped, and finally the voltage application to the filament is stopped.

  The controller 16b has a timer (not shown) that starts timing when X-ray irradiation is started. The control unit 16b is set in the radiation source control device 16 when the time measured by the timer reaches the irradiation time set in the irradiation condition without receiving the irradiation stop signal by the communication unit 16c, or for safety regulations. When the maximum irradiation time is reached, X-ray irradiation is stopped in the same manner as when the irradiation stop signal is received by the communication unit 16c.

  The imaging control device 13 includes a control unit 13a and a communication unit 13b. The communication unit 13 b relays transmission and reception of information between the radiation source control device 16 and the electronic cassette 12 and between the electronic cassette 12 and the console 14. The information exchanged between the radiation source control device 16 and the electronic cassette 12 includes the above-described synchronization signal, and the information exchanged between the electronic cassette 12 and the console 14 includes X-ray images and imaging condition information.

  When receiving an irradiation start request signal from the radiation source control device 16 via the imaging control device 13, the electronic cassette 12 starts an accumulation operation for performing image detection, and directs the irradiation permission signal to the imaging control device 13. To send. At the same time, sampling of the X-ray dose for the AEC function is started, and the remaining time until the accumulated X-ray dose reaches the target dose is calculated for each sampling of the dose. The electronic cassette 12 transmits the predicted remaining time to the communication unit 13b as a countdown signal.

  The control unit 13a of the imaging control device 13 predicts in advance the time (stop time) when the remaining time becomes zero based on the countdown signal sequentially received by the communication unit 13b. The communication unit 13b transmits an irradiation stop signal to the radiation source control device 16 when the stop time comes. The controller 13a recalculates the stop time based on the latest countdown signal and updates it. Since the stop time is predicted by the control unit 13a every time a countdown signal is received, an irradiation stop signal can be generated and transmitted based on the received countdown signal even if reception of the countdown signal stops.

  When it is determined by the AEC function that the accumulated dose has reached the target dose, the electronic cassette 12 ends the accumulation operation and starts the image reading operation. The electronic cassette 12 transmits the X-ray image generated by the image reading operation to the imaging control device 13. The imaging control device 13 transfers the X-ray image from the electronic cassette 12 to the console 14.

  In FIG. 2, an imaging condition table 25 is stored in the storage device 14 c of the console 14. The imaging conditions include information on the subject H such as the imaging part, the sex of the subject H, the age, the body thickness of the subject H (only the imaging part is shown in FIG. 2), and the irradiation conditions of the X-rays irradiated by the X-ray source 15. , Information on the lighting field and information on the target dose when the electronic cassette 12 performs AEC are included. The irradiation conditions include the tube voltage, the tube current, and the X-ray irradiation time as described above, and are determined in consideration of information regarding the imaging region and the subject H.

  The daylighting field indicates an area where the accumulated dose of X-rays is calculated when performing AEC. For example, a region of interest that is most noticeable at the time of diagnosis is set. When the imaging region is, for example, the chest, the region of interest is the left and right lung fields as indicated by Aa and Bb in FIG. The target dose is a threshold for making a determination to stop the irradiation of X-rays by comparing with the accumulated dose in the daylighting field in AEC.

  In the imaging condition table 25, the correspondence between the imaging region such as the chest and the abdomen and the irradiation condition and the like corresponding to the imaging region is recorded, and when the operator selects the imaging region via the input device 14a, the corresponding irradiation is performed. Conditions and others are read out and displayed on the operation screen of the display 14b. Each value of the irradiation condition read from the imaging condition table 25 can be finely adjusted according to the sex, age, and body thickness of the subject H. The operator confirms the contents of the inspection order on the display 14b, and inputs imaging conditions corresponding to the contents on the input device 14a. Information on the set shooting conditions is transmitted from the console 14 to the shooting control device 13 and transferred from the shooting control device 13 to the electronic cassette 12. In the imaging condition table 25 of this example, the tube current and the irradiation time are individually recorded. However, since the total amount of X-ray irradiation dose is determined by the product of the tube current and the irradiation time, the tube that is the product of the two products is used. The value of the current irradiation time product (mAs value) may be recorded.

  In the radiation source control device 16, the same irradiation conditions as those input to the console 14 are set by the operator through the operation panel 19. However, when the AEC function is used, the irradiation time set in the radiation source control device 16 is set to a time sufficiently longer than the irradiation time set in the console 14. The reason for this setting is that when the AEC function is used, the X-ray irradiation stop timing is determined not by the radiation source control device 16 but by the imaging control device 13 side. This is because it is necessary to prevent radiation exposure from being stopped and causing a shortage of dose. When the AEC function is used, the irradiation time set in the radiation source control device 16 and the safety regulations preset in the radiation source control device 16 for each imaging region in order to avoid excessive exposure to the subject H The maximum irradiation time functions as a backup time when the AEC function does not work normally. In addition, when using the AEC function, the irradiation time set on the console 14 does not mean that the X-ray irradiation is stopped during the irradiation time, and depends on the imaging region and the body thickness of the subject H. It is just a recommended value. Hereinafter, the irradiation time set in the console 14 when using the AEC function is referred to as a recommended irradiation time.

  In FIG. 3, the electronic cassette 12 includes a detection panel 30, a control board 31, a battery 32, a communication unit 33, and a portable box-shaped casing 34 that accommodates these. The housing 34 is made of, for example, a conductive resin. A rectangular opening is formed in the front surface 34a of the housing 34 on which X-rays are incident, and a transmission plate 35 is attached to the opening as a top plate. The transmission plate 35 is made of a carbon material that is lightweight, has high rigidity, and has high X-ray permeability.

  The housing 34 has a size compliant with, for example, an international standard ISO 4090: 2001 for a film cassette or an IP cassette. For this reason, it can be attached to an existing photographing stand for a film cassette or an IP cassette. In addition to being set on the photographing stand, the electronic cassette 12 may be used alone by being placed on a bed on which the subject H lies, or on the subject H itself.

  The battery 32 supplies power to each part of the electronic cassette 12 through a power supply circuit (not shown). The battery 32 can be taken out from the inside of the housing 34, and can be charged by being set in a dedicated charger (not shown). The communication unit 33 is wired or wirelessly connected to the imaging control device 13 and transmits / receives various information such as the above-described synchronization signal, countdown signal, imaging conditions, and X-ray image to / from the imaging control device 13.

The detection panel 30 includes a scintillator (phosphor) 36 and a light detection substrate 37. The scintillator 36 and the light detection substrate 37 are laminated in the order of the scintillator 36 and the light detection substrate 37 when viewed from the X-ray incident side. The scintillator 36 has a phosphor such as CsI: Tl (thallium-activated cesium iodide) or GOS (Gd ₂ O ₂ S: Tb, terbium-activated gadolinium oxysulfide), and is incident through the transmission plate 35. Is converted into visible light and emitted.

  The light detection substrate 37 detects visible light emitted from the scintillator 36 and converts it into an electrical signal. The control board 31 controls driving of the light detection board 37 and generates an X-ray image based on the electrical signal output from the light detection board 37.

  In FIG. 4, the light detection substrate 37 includes a pixel 40 arranged in a two-dimensional matrix of n rows × m columns, n scanning lines 41, and m signal lines on a glass substrate (not shown). 42 is provided. The scanning lines 41 extend in the X direction along the row direction of the pixels 40 and are arranged at a predetermined pitch in the Y direction along the column direction of the pixels 40. The signal lines 42 extend in the Y direction and are arranged at a predetermined pitch in the X direction. The scanning lines 41 and the signal lines 42 are orthogonal to each other, and the pixels 40 are provided corresponding to the intersections of the scanning lines 41 and the signal lines 42. An area of the glass substrate on which the pixels 40 are arranged forms an imaging area 43. n and m are integers of 2 or more, for example, n and m≈2000. The position of the pixel 40 is represented by, for example, XY coordinates with the coordinates of the upper left pixel 40 at the origin (0, 0). The arrangement of the pixels 40 may not be a square arrangement as in this example, but may be a honeycomb arrangement.

  As is well known, each pixel 40 includes a photoelectric conversion unit 44 that generates charges (electron-hole pairs) by the incidence of visible light and accumulates them, and a TFT (Thin-Film Transistor) 45 that is a switching element. Prepare. The photoelectric conversion unit 44 has a structure in which a semiconductor layer that generates electric charges and an upper electrode and a lower electrode are arranged above and below the semiconductor layer. The semiconductor layer is, for example, a PIN (p-intrinsic-n) type, and an N-type layer is formed on the upper electrode side and a P-type layer is formed on the lower electrode side. The TFT 45 has a gate electrode connected to the scanning line 41, a source electrode connected to the signal line 42, and a drain electrode connected to the lower electrode of the photoelectric conversion unit 44.

  A bias line (not shown) is connected to the upper electrode of the photoelectric conversion unit 44. Bias lines are provided for the number of rows of pixels 40 (n rows) and connected to one bus line. The bus is connected to a bias power source. A positive bias voltage is applied from the bias power source to the upper electrode of the photoelectric conversion unit 44 through the bus line and its bias line. Application of a positive bias voltage generates an electric field in the semiconductor layer. The photoelectric conversion unit 44 is used in a reverse bias state. Electrons among the electron-hole pairs generated in the semiconductor layer by photoelectric conversion move to the upper electrode and are absorbed by the bias line, and the holes move to the lower electrode and are collected as signal charges.

  The pixel 40 is divided into a normal pixel 40a and a detection pixel 40b. The normal pixel 40a is used for generating an X-ray image. On the other hand, the detection pixel 40 b functions as an X-ray detection unit for detecting a dose (X-ray intensity) per unit time of X-rays reaching the imaging region 43. The TFT 45 of the detection pixel 40 b is short-circuited by connecting the source electrode and the drain electrode by a short-circuit line 46. On the other hand, the TFT 45 of the normal pixel 40a is not short-circuited. In the figure, the detection pixel 40b is hatched to be distinguished from the normal pixel 40a.

  The TFT 45 of the detection pixel 40b is always on regardless of the voltage applied to the gate electrode because the source and drain electrodes are short-circuited. Therefore, the signal charge collected on the lower electrode of the photoelectric conversion unit 44 flows out to the signal line 42 via the short-circuit line 46. On the other hand, in the normal pixel 40a, the signal charge is read to the signal line 42 in accordance with the voltage applied to the gate electrode of the TFT 45.

  The detection pixel 40b is different from the detection pixel 40b only in that the source and drain electrodes of the TFT 45 are short-circuited, and other basic configurations such as the photoelectric conversion unit 44 are exactly the same as those of the normal pixel 40a. Therefore, both can be formed by substantially the same manufacturing process. The detection pixel 40b can be configured by directly connecting the photoelectric conversion unit 44 to the signal line 42 without providing the TFT 45. However, in this embodiment, the detection pixel 40b has a structure including the normal pixel 40a and the detection pixel 40b. In order not to cause a difference in characteristics, the TFT 45 is provided in the same manner as the normal pixel 40a, and the source and drain electrodes are short-circuited by the short-circuit line 46 to form the detection pixel 40b.

  The number of detection pixels 40b is, for example, about 1 / million to several hundredths of the number of all pixels 40. The signal line 42 is divided into a first signal line 42a in which one detection pixel 40b and a plurality of normal pixels 40a are connected, and a second signal line 42b in which only the normal pixels 40a are connected. The first signal lines 42a are periodically provided with about 1 to 3 (in this embodiment, two) second signal lines 42b.

  The positions of the detection pixels 40 b in the Y direction are changed for each first signal line 42 a and are arranged so as to be evenly distributed in the imaging region 43. For example, as illustrated in FIG. 5, the detection pixels 40 b are arranged so as to draw a zigzag locus 47 that is symmetrical with respect to the center of the imaging region 43. In this way, a plurality of detection pixels 40 b are arranged in a distributed manner throughout the imaging region 43. For this reason, even when the X-ray irradiation range is limited to a part of the imaging region 43, X-rays can be detected by any of the detection pixels 40b. The position of the detection pixel 40b is known when the detection panel 30 is manufactured, and the electronic cassette 12 stores in advance the XY coordinates of all the detection pixels 40b in the internal memory 72 (see FIG. 11) of the control unit 54.

  The control board 31 is provided with a gate driver 50, a signal processing circuit 51, a memory 52, an AEC unit 53, and a control unit 54 for controlling them. The gate driver 50 is connected to the end of each scanning line 41. As shown in FIG. 6, the controller 54 drives the TFT 45 through the gate driver 50, thereby reading out and resetting (discarding) dark charges from the normal pixels 40 a and the X-ray arrival dose. The detection panel 30 is caused to perform an accumulation operation for accumulating the signal charge in the normal pixel 40a and an image reading operation for reading the signal charge from the normal pixel 40a. In the image reading operation and the pixel reset operation, the gate driver 50 sequentially applies a gate pulse to each scanning line 41 from the first row to the n-th row, and sets the TFT 45 of the normal pixel 40a connected to each scanning line 41 to one row. Turn on sequentially one by one. The time during which the TFT 45 is turned on is defined by the pulse width of the gate pulse, and the TFT 45 returns to the off state after the time defined by the pulse width has elapsed. On the other hand, in the accumulation operation, the gate driver 50 does not give a gate pulse to any scanning line 41. Therefore, during the accumulation operation, the TFT 45 of the normal pixel 40a is turned off.

  The signal processing circuit 51 is connected to the end of each signal line 42. The signal processing circuit 51 includes an integrating amplifier 55, a gain amplifier 56, a CDS (Correlated Double Sampling) circuit 57, a multiplexer (MUX) 58, and an A / D converter 59.

  An integrating amplifier 55 is provided for each signal line 42. The integrating amplifier 55 includes an operational amplifier 55a, a capacitor 55b, and an amplifier reset switch 55c. The operational amplifier 55a has two input terminals and one output terminal. The signal line 42 is connected to one of the two input terminals, and the ground line is connected to the other. The capacitor 55b and the amplifier reset switch 55c are connected in parallel between the input terminal to which the signal line 42 is connected and the output terminal.

  The integrating amplifier 55 integrates the signal charges input from the signal line 42 by accumulating them in the capacitor 55b, and outputs an analog voltage value (signal voltage) corresponding to the integrated value. The unpreset switch 55 c is driven and controlled by the control unit 54. By turning on the unpreset switch 55c, the signal charge accumulated in the capacitor 55b is reset (discarded).

  The gain amplifier 56 is connected to the output terminal of the operational amplifier 55a, and amplifies the signal voltage output from the integrating amplifier 55 with a predetermined gain value. The gain value is set by the control unit 54 based on the shooting conditions from the console 14.

  The CDS circuit 57 is connected to the output terminal of the gain amplifier 56, and performs a well-known correlated double sampling process on the signal voltage amplified by the gain amplifier 56, and the reset noise component of the integrating amplifier 55 is obtained from the signal voltage. Remove. Specifically, the CDS circuit 57 includes two sample and hold circuits (not shown) and one difference circuit (not shown). One sample and hold circuit samples and holds the signal voltage output from the gain amplifier 56, and the integration amplifier 55 reset noise output from the gain amplifier 56 when the integration amplifier 55 is reset by the other sample and hold circuit. Sample and retain ingredients. By taking the difference between the two in the difference circuit, a signal voltage from which noise has been removed is obtained.

  The MUX 58 is connected to the output terminal of each CDS circuit 57, selects the CDS circuits 57 in the 1st to m-th columns one by one in order, and is output from each CDS circuit 57 to the A / D converter 59. Input signal voltage serially. The A / D converter 59 performs A / D conversion processing on the input signal voltage and outputs a digital signal voltage.

  In the image reading operation, the control unit 54 drives the gate driver 50 and each part of the signal processing circuit 51 (integration amplifier 55, CDS circuit 57, MUX 58, A / D converter 59) at a predetermined cycle, respectively. 41 is sequentially selected from the first row to the n-th row, and as a result, the signal voltage for one row output from the A / D converter 59 is stored in the memory 52 sequentially. When the image reading operation from the first row to the n-th row is completed, the memory 52 stores a signal voltage representing one X-ray image in association with the XY coordinates of each pixel 40. The signal voltage data is read from the memory 52 to the control unit 54, subjected to various correction processes by the control unit 54, and then transmitted to the console 14 through the communication unit 33. In this way, an X-ray image of the subject H is detected. Hereinafter, a signal voltage read by the image reading operation is referred to as an image signal (see FIG. 6).

  In the pixel reset operation, each part of the gate driver 50 and the signal processing circuit 51 is driven at the same predetermined cycle as the image readout operation. In the pixel reset operation, dark charges flow from the normal pixel 40 a to the capacitor 55 b of the integration amplifier 55 through the signal line 42 while the TFT 45 is in the on state. Unlike the image readout operation, the signal voltage is not read out in the pixel reset operation, and in synchronization with the fall of the gate pulse, the amplifier reset pulse is output from the control unit 54, the amplifier reset switch 55c is turned on, and the capacitor 55b is turned on. The accumulated dark charge is discharged and the integrating amplifier 55 is reset.

  As described above, the signal charge collected on the lower electrode of the photoelectric conversion unit 44 of the detection pixel 40 b always flows out to the first signal line 42 a via the short-circuit line 46. This signal charge flows into the capacitor 55b of the integrating amplifier 55 connected to the first signal line 42a and is accumulated in the capacitor 55b. That is, even during the accumulation operation in which the normal pixel 40a is irradiated with the X-ray and the TFT 45 is turned off, the signal charge of the detection pixel 40b is converted into a signal voltage by driving each part of the signal processing circuit 51. Can be read out.

  As shown in FIG. 6, each part of the signal processing circuit 51 performs a dose sampling operation for reading out a signal voltage corresponding to the signal charge of the detection pixel 40b at the sampling period set by the control unit 54 simultaneously with the start of the accumulation operation. Start. Since the signal charge of the detection pixel 40b changes according to the amount of X-rays incident on the imaging region 43, the signal voltage obtained by one sampling is the unit time (sampling period) of the X-rays that reach the imaging region 43. ) Represents the dose per unit. Hereinafter, the signal voltage read by this dose sampling operation is referred to as a dose signal.

  In the dose sampling operation, an amplifier reset pulse is input from the control unit 54 to the amplifier reset switch 55c of the integrating amplifier 55 every time sampling is performed, and the accumulated charge is reset. In one sampling of the dose sampling operation, a plurality of CDS circuits 57 are sequentially selected by the MUX 58 and the dose signal for one row is sampled as in the case of reading the image signal for one row in the image reading operation. . Since the outputs of the plurality of detection pixels 40b dispersedly arranged in the imaging region 43 always flow out to the first signal line 42a, all the detection pixels are included in the dose signal for one row read by the dose sampling operation. A dose signal corresponding to 40b is included. In the memory 52, the dose signal for one row is recorded in association with the XY coordinates of the detection pixel 40b.

  As described above, the signal processing circuit 51 functions as an image signal reading unit that reads an image signal from the detection panel 30 and periodically outputs a dose signal corresponding to a dose per unit time of X-rays emitted from the X-ray source 15. It also functions as a dose sampling unit for sampling.

  The AEC unit 53 reads a dose signal from the memory 52 every time a dose signal for one row is recorded in the memory 52. Then, the accumulated dose is obtained based on the read dose signal, and it is determined whether or not the accumulated dose has reached the target dose.

  In FIG. 7, the AEC unit 53 includes a dose signal selection unit 65, a dose signal integration unit 66, and a countdown signal generation unit 67. The dose signal selection unit 65 selects a dose signal to be used for AEC from the dose signals for one row based on the XY coordinates of the detection pixel 40b and the information on the daylighting field among the imaging conditions from the console 14. The information of the lighting field is represented by XY coordinates, and when the lighting field is rectangular, for example, it is represented by two XY coordinates connected by diagonal lines. The dose signal selection unit 65 identifies the detection pixel 40b in the daylighting field based on the detection pixel 40b and the XY coordinates of the daylighting field, and outputs the dose signal output from the identified detection pixel 40b to the dose signal for one line. The selected dose signal is selected from the above, and the selected dose signal is delivered to the dose signal integrating unit 66 as a dose signal used for AEC.

  Each time the dose signal integrating unit 66 receives a dose signal from the detection pixel 40b in the daylighting field from the dose signal selection unit 65, the representative value of the dose signal from the detection pixel 40b in the daylighting field, for example, an average value or an intermediate value Calculate maximum value, mode value, total value, etc. The dose signal integration unit 66 integrates the representative values to obtain the accumulated dose in the lighting field. The dose signal integration unit 66 is provided with a memory (not shown) for stocking information on the accumulated dose in the daylighting field. The dose signal integrating unit 66 adds the representative value calculated in the current sampling to the accumulated dose of the sampling field stocked in the memory at the previous sampling, and overwrites this in the memory to calculate the cumulative dose of the sampling field. Information is updated every sampling. The dose signal integration unit 66 passes information on the accumulated dose in the lighting field to the countdown signal generation unit 67.

  The countdown signal generation unit 67 performs an extrapolation operation based on the time change of the accumulated dose of the daylighting field received from the dose signal integration unit 66 at each sampling, thereby calculating the time when the accumulated dose of the daylighting field reaches the target dose. Predict. Specifically, as shown in FIG. 8, first, an interpolation line L1 is calculated by linearly interpolating the accumulated dose (indicated by a circle) of the daylighting field obtained up to the current time T1, and this interpolation is performed. A straight line obtained by extending the line L1 after the current time T1 is defined as an extrapolation line L2. Based on this extrapolated line L2, a predicted arrival time T2 at which the accumulated dose in the lighting field reaches the target dose TH is obtained. Then, a remaining time TR that is the difference between the predicted arrival time T2 and the current time T1 is obtained, and a countdown signal representing the remaining time TR is generated.

  The countdown signal generation unit 67 performs the extrapolation calculation and generates the countdown signal every time it receives the accumulated dose of the daylighting field from the dose signal integration unit 66. The countdown signal generation unit 67 outputs the generated countdown signal to the control unit 54. Whenever the control unit 54 receives a countdown signal from the countdown signal generation unit 67, the control unit 54 outputs the countdown signal to the communication unit 33, and the communication unit 33 transmits this to the imaging control device 13. In addition, when the remaining time TR represented by the countdown signal becomes zero, the control unit 54 ends the accumulation operation of the detection panel 30 and starts the image reading operation.

  The control unit 54 is provided with a circuit (not shown) that performs various correction processes such as offset correction, sensitivity correction, and defect correction on the X-ray image stored in the memory 52 by the image reading operation. The offset correction circuit subtracts the offset correction image acquired by performing the image reading operation without irradiating the X-ray in units of pixels from the X-ray image, thereby fixing the fixed pattern caused by the individual difference of the signal processing circuit 51 and the imaging environment. Noise is removed from the X-ray image. The sensitivity correction circuit is also called a gain correction circuit, and corrects variations in sensitivity of the photoelectric conversion unit 44 of each pixel 40, variations in output characteristics of the signal processing circuit 51, and the like. The defect correction circuit linearly interpolates the pixel value of the defective pixel with the pixel values of the surrounding normal pixels based on the defective pixel information generated at the time of shipment or regular inspection. The defect correction circuit also treats the detection pixel 40b as a defective pixel. Since the pixel value of the normal pixel 40a in the column of the first signal line 42a in which the detection pixel 40b is arranged is also affected by the output of the detection pixel 40b that always flows out, the defective pixel correction circuit determines the pixel value of the detection pixel 40b. Alternatively, defect correction is also performed by linear interpolation on the pixel value of the normal pixel 40a in the column of the first signal line 42a in which the detection pixel 40b is arranged. Note that the various correction processing circuits described above may be provided in the imaging control device 13 and the console 14, and various correction processes may be performed by the imaging control device 13 and the console 14.

  As shown in FIGS. 9 and 10, the sampling period SP of the dose signal in this embodiment is a charge accumulation time in which the integrating amplifier 55 (abbreviated as CA: Charge Amp in FIGS. 9 and 10) accumulates and accumulates charges. (CA accumulation time) and the readout time (reading: abbreviated as TCA) in which the dose signal is read out from the integrating amplifier 55 and output to the memory 52. Since the readout time TCA is constant regardless of the amount of charge accumulation (CA accumulation amount) of the integration amplifier 55, changing the sampling period SP is synonymous with changing the charge accumulation time of the integration amplifier 55. is there. The sampling period SP is defined by the generation period of the CA readout pulse for reading out the dose signal from the integrating amplifier 55.

  As shown in FIGS. 9 and 10, even if the signal charge flowing out from the detection pixel 40b to the first signal line 42a is a constant amount regardless of time, the CA accumulation amount of the integrating amplifier 55 (indicated by hatching). Therefore, the signal component of the dose signal corresponding to the CA accumulation amount increases as the sampling period SP increases. Therefore, the signal component of the dose signal obtained by one sampling becomes larger when the sampling period SP is longer than the sampling period SPa shown in FIG. 9 as in the sampling period SPb shown in FIG. The dose signal includes a noise component due to stationary noise that is steadily generated by the signal processing circuit 51. Since the magnitude of the stationary noise is almost constant regardless of the CA accumulation amount, the signal component of the dose signal is The larger the value, the higher the S / N ratio of the dose signal.

  If the S / N ratio of the dose signal is too low, the remaining time TR obtained by the countdown signal generation unit 67 based on the dose signal also decreases in reliability, and when the accumulated dose of X-rays reaches the target dose, the X There is a possibility that the irradiation of the line cannot be stopped. Therefore, in order to stop the X-ray irradiation accurately when the cumulative dose of X-rays reaches the target dose, it is necessary to increase the S / N ratio of the dose signal to some extent.

  When the sampling period SP is the same, the signal component of the dose signal obtained by one sampling becomes smaller as the dose of X-rays reaching the imaging region 43 is lower. Therefore, when the X-ray dose reaching the imaging region 43 is relatively low, it is necessary to lengthen the sampling period SP in order to increase the S / N ratio of the dose signal. Conversely, when the dose of X-rays reaching the imaging region 43 is relatively high, a dose signal having a sufficiently high S / N ratio can be obtained even if the sampling period SP is somewhat short. As the sampling period SP is shorter, the number of times the remaining time TR is predicted and increased, and the accuracy of the prediction calculation is increased accordingly. Therefore, if a high S / N ratio is ensured, the sampling period SP should be short. As described above, the sampling period SP has an optimum value corresponding to the X-ray dose reaching the imaging region 43. Therefore, in the present invention, during the dose sampling operation, it is determined whether or not the sampling cycle SP is appropriate based on the dose signal. If it is determined that the sampling cycle SP is inappropriate, the sampling cycle SP is changed to an optimum value. Get close to.

  In FIG. 11, the control unit 54 is provided with a cycle suitability determination unit 70 and a cycle setting unit 71. The cycle suitability determination unit 70 reads the dose signal for one row from the memory 52 during the dose sampling operation, analyzes the read dose signal for one row, and determines whether the sampling cycle SP is appropriate for the dose. The determination result is output to the period setting unit 71. The period setting unit 71 sets the sampling period SP. Specifically, since the sampling period SP is defined by the CA read pulse generation period, the period setting unit 71 changes the sampling period SP by changing the CA read pulse generation period.

  The cycle suitability determination unit 70 first removes the dummy signal from the dose signal for one row. As described above, the column of the first signal lines 42a provided with the detection pixels 40b is provided by sandwiching a plurality of columns of the second signal lines 42b where the detection pixels 40b are not provided. The dose signal includes a dummy signal corresponding to the column of the second signal line 42b between the significant dose signals of the column of the first signal line 42a. The dummy signal is based on the leak charge leaking from the normal pixel 40a and is almost zero. For this reason, the cycle suitability determining unit 70 determines suitability of the sampling cycle SP after removing dummy signals that are meaningless as data. Instead of removing the dummy signal after reading the dose signal for one row from the memory 52, only the dose signal corresponding to the detection pixel 40b may be selected when reading from the memory 52.

  In FIG. 12, the cycle suitability determination unit 70 creates a histogram with the horizontal axis representing the signal value of the dose signal for one row from which the dummy signal is removed, and the vertical axis representing the number of signal values. The periodic propriety determination unit 70 analyzes the created histogram, and identifies a portion VF where the peak shape including the signal value VSmax having the maximum number falls to the left side of the horizontal axis of the histogram having a small signal value. Then, on the right side of the horizontal axis of the histogram with a large signal value, with the specified location VF as a boundary, the blank area where the subject H does not exist and the X-ray directly enters, the left side of the horizontal axis of the histogram with a small signal value, It is recognized as a subject area irradiated with X-rays transmitted through the subject H.

  The cycle suitability determination unit 70 reads the threshold value THV1 (corresponding to the first threshold value) stored in the internal memory 72 of the control unit 54. Then, for example, on the left side of the VF recognized as the subject area, the magnitude relationship between the signal value VHmax having the maximum number and the threshold value THV1 is compared. When the signal value VHmax exceeds the threshold value THV1 (when the signal value VHmax is on the right side of the threshold value THV1) as shown in FIG. 12, the cycle suitability determination unit 70 determines that the sampling cycle SP is appropriate, and vice versa. When the signal value VHmax falls below the threshold value THV1 (when the signal value VHmax is on the left side of the threshold value THV1), it is determined that the sampling period SP is inappropriate.

  The threshold value THV1 is set to a relatively small value close to zero. For this reason, when the signal value VHmax falls below the threshold value THV1, the sampling period SP is too short, and the signal values of the dose signals from the detection pixels 40b existing in the subject region and thus in the lighting field are almost all zero. This means that the S / N ratio of the dose signal is too low for accurate AEC.

  Note that there is a slight time lag from when the detection panel 30 starts the accumulation operation and the dose sampling operation in response to the irradiation start request signal to when the X-ray irradiation actually starts. Further, the dose per unit time of X-rays is small immediately after the start of irradiation, and gradually increases toward a set dose determined according to the tube current. Therefore, the signal value of the dose signal for one line obtained during the time lag and immediately after the start of irradiation is almost zero or close to zero regardless of the subject region and the missing region. If the appropriateness determination of the sampling period SP is started at this timing, it will be determined as inappropriate in all cases. For this reason, the cycle suitability determination unit 70 does not determine suitability of the sampling cycle SP while the maximum signal value VSmax in the unclear region is zero or a value close thereto, and the signal value VSmax becomes a relatively large value. Start adequacy determination. Specifically, after the accumulation operation is started by the detection panel 30, the suitability determination is started after being delayed by a preset time. Further, the cycle suitability determination unit 70 repeats the suitability determination until it determines that the sampling cycle SP is appropriate, and ends the suitability determination after determining that it is appropriate.

  The cycle setting unit 71 does not change the set sampling cycle SP when the cycle suitability determination unit 70 determines that the sampling cycle SP is appropriate. On the other hand, when the cycle suitability determining unit 70 determines that the sampling cycle SP is inappropriate, the cycle setting unit 71 changes the sampling cycle SP in accordance with a predetermined setting condition. The cycle setting unit 71 continues changing the sampling cycle SP until the cycle suitability determining unit 70 determines that the sampling cycle SP is appropriate.

  The setting conditions for the sampling period SP include an initial value SP1, a change step SPstp, and an upper limit value SPmax. These pieces of information are stored in the internal memory 72 of the control unit 54. The initial value SP1 is a sampling period initially set by the period setting unit 71. As shown in FIG. 13, 1 ms is set in this embodiment. In order to secure the CA accumulation time of the integrating amplifier 55, the initial value SP1 is set to a value longer than the reading time TCA from which the dose signal is read from the integrating amplifier 55 and output to the memory 52 (SP1> TCA). ). The change step SPstp defines the interval of the sampling period SP that is changed once, and is the difference or ratio between the sampling period SP before and after the change when the sampling period SP is changed stepwise. In the present embodiment, the change step SPstp is set to a ratio of 2 (× 2). For this reason, the sampling period SP set by the period setting unit 71 becomes longer in steps of 1 ms (SP1), 2 ms (SP2), 4 ms (SP3), 8 ms (SP4),.

  The upper limit value SPmax is literally the maximum value of the sampling period SP that can be set by the period setting unit 71, and is shorter than 1/2 of the X-ray recommended irradiation time TX in the imaging conditions from the console 14 in this embodiment. A value is set (SPmax <TX / 2). The control unit 54 functions as a setting condition changing unit, and changes the upper limit value SPmax according to the recommended irradiation time TX.

  The reason why the upper limit SPmax is set to a value shorter than ½ of the recommended irradiation time TX is as follows. In order to calculate the interpolation line L1 at the time of generating the countdown signal, data on the accumulated dose of at least two lighting fields is necessary, and for this purpose, it is necessary to sample the dose signal at least twice. However, if the sampling period SP is 1/2 or more of the recommended irradiation time TX, there is a possibility that sampling can be performed only once during the recommended irradiation time TX.

  As a result of stepwise changing the sampling period SP, the period setting unit 71 determines that the period appropriateness determining unit 70 determines that the sampling period SP is inappropriate when the sampling period SP reaches the upper limit value SPmax. In this case, the upper limit value SPmax remains unchanged without changing the sampling period SP.

  Next, the effect | action by the said structure is demonstrated with reference to the flowchart of FIG. 14 and FIG. When performing X-ray imaging with the X-ray imaging system 2, first, the subject H is set at one of the imaging positions 17 and 18 in the standing position and the standing position, and the height or horizontal position of the electronic cassette 12 is set. To adjust the position of the subject H to the imaging region. Then, the height, horizontal position, and irradiation field size of the X-ray source 15 are adjusted according to the position of the electronic cassette 12 and the size of the imaging region. Next, shooting conditions are set on the console 14. In addition, the same X-ray irradiation conditions as those set for the console 14 are set in the radiation source control device 16. The imaging conditions set by the console 14 are transmitted to the imaging control device 13 and transferred from the imaging control device 13 to the electronic cassette 12.

  The electronic cassette 12 waits for imaging conditions from the console 14 in a sleep state in which only the communication unit 33 is driven, for example. In FIG. 14, when the imaging conditions are received from the console 14 (YES in step S100), the electronic cassette 12 starts a pixel reset operation (step S110). In the pixel reset operation, a gate pulse is emitted from the gate driver 50, and unnecessary charges of the normal pixels 40a for one row are read to the signal line 42 and discarded by the integrating amplifier 55 in a predetermined cycle. Is called.

  When the setting of the imaging condition on the console 14 and the setting of the irradiation condition on the radiation source control device 16 are completed, the operator half-presses the irradiation switch 20. When the irradiation switch 20 is half-pressed, a warm-up instruction signal is input to the radiation source control device 16 and the warm-up of the X-ray source 15 is started. Further, an irradiation start request signal is transmitted from the radiation source control device 16 to the imaging control device 13. The imaging control device 13 transfers an irradiation start request signal to the electronic cassette 12.

  When the irradiation start request signal is received (YES in step S120), the electronic cassette 12 finishes the pixel reset operation up to the n-th row and then starts an accumulation operation (step S130), and simultaneously sends an irradiation permission signal to the imaging control device. 13 to send. In the accumulation operation, a gate pulse is not applied from the gate driver 50, and therefore signal charges corresponding to the dose of irradiated X-rays are accumulated in the normal pixel 40a.

  Simultaneously with the start of the accumulation operation, the cycle setting unit 71 sets the sampling cycle SP to the initial value SP1 (step S140), and the dose sampling operation is started (step S150). The signal charge of the detection pixel 40 b flows out to the signal line 42 and is accumulated in the integrating amplifier 55 because the source electrode and the drain electrode of the TFT 45 are short-circuited. In the dose sampling operation, the dose signal corresponding to the signal charge of the detection pixel 40 b accumulated in the integration amplifier 55 is sampled at the sampling period SP set by the period setting unit 71. The signal charge of the detection pixel 40b accumulated in the integrating amplifier 55 is reset every sampling. In one sampling, a plurality of CDS circuits 57 are sequentially selected by the MUX 58 and a dose signal for one row is recorded in the memory 52.

  The operator presses the irradiation switch 12 halfway, and then presses the irradiation switch 12 fully in response to the time required for warm-up. When the irradiation switch 20 is fully pressed, an irradiation start instruction signal is input to the radiation source control device 16 and X-ray irradiation is started from the X-ray source 15. When the X-ray irradiation is started, the signal value of the dose signal increases.

  The dose signal for one row recorded in the memory 52 is read out by the periodic suitability determination unit 70 every sampling. Then, as shown in step S160, the cycle suitability determining unit 70 determines whether or not the currently set sampling cycle SP is appropriate for the dose. Specifically, first, the dummy signal corresponding to the column of the second signal line 42b is removed from the dose signal for one row, and then the signal value of the dose signal for one row is plotted on the horizontal axis and the number of signal values. A histogram with the vertical axis as the vertical axis is created, and the missing region and the subject region are recognized in the histogram. Then, the suitability of the sampling period SP is determined by comparing the magnitude of the signal value VHmax having the maximum number in the portion of the histogram recognized as the subject region with the threshold value THV1. The suitability determination of the sampling cycle SP by the cycle suitability determination unit 70 is started after a predetermined time after the accumulation operation is started by the detection panel 30.

  In FIG. 15, when the signal value VHmax exceeds the threshold value THV1, the cycle suitability determining unit 70 determines that the sampling cycle SP is appropriate (YES in step S161). In this case, the currently set sampling period SP is not changed.

  Conversely, if the signal value VHmax falls below the threshold value THV1, the cycle suitability determining unit 70 determines that the sampling cycle SP is inappropriate (NO in step S161). In this case, the currently set sampling period SP is changed by the period setting unit 71 in accordance with the change step SPstp (step S163), and the dose signal for one line obtained in the changed sampling period SP is added. Based on this, the suitability determination of the sampling cycle SP is performed again by the cycle suitability determination unit 70. In the present embodiment, the sampling period SP is changed from the initial value of 1 ms to 2 ms, 4 ms, 8 ms,. However, if the currently set sampling period SP is the upper limit value SPmax (YES in step S162), the sampling period SP is not changed.

  In the present invention, the sampling period SP is changed when it is determined that the sampling period SP is inappropriate based on the dose signal. Therefore, when a relatively high dose is set in the irradiation condition or the body thickness of the subject is relatively thin. Even when the dose of X-rays reaching the imaging region 43 is relatively high as in the case of the case, on the contrary, when a relatively low dose is set under the irradiation conditions, or when the body thickness of the subject is relatively thick Thus, even when the dose of X-rays reaching the imaging region 43 is relatively low, an appropriate sampling period SP that can ensure a high S / N ratio can be obtained. Therefore, accurate AEC can be performed.

  In this embodiment, since the appropriateness determination is repeated by changing the sampling period SP until the appropriateness determination unit 70 determines that the appropriate sampling period SP is appropriate, the X-ray doses reaching the imaging region 43 are compared. Whether the target is high or low, the sampling period SP can be brought close to the optimum value. In addition, since the initial value SP1 is set to be relatively short 1 ms, and then the sampling period SP is increased stepwise, the sampling period SP is appropriately set at an early stage when the dose of X-rays reaching the imaging region 43 is relatively high. Can be set to any value.

  Even if the X-ray dose is extremely low, if the sampling period SP is changed to a long value in accordance with this, the amount of accumulated CA in one sampling is larger than when the sampling period SP is short. Therefore, the signal component of the dose signal is increased accordingly, and the S / N ratio of the dose signal is increased. Therefore, accurate AEC can be performed even in low-dose imaging.

  In the dose sampling operation, the dose signal for one row recorded in the memory 52 is also read out to the AEC unit 53. As shown in step S <b> 170 of FIG. 14, in the AEC unit 53, the dose signal output from the detection pixel 40 b in the lighting field is selected from the dose signal for one row by the dose signal selection unit 65. The dose signal integration unit 66 calculates the representative value of the dose signal from the detection pixel 40b in the daylighting field and the cumulative dose of the daylighting field as the integration value (step S180). The information on the accumulated dose in the daylighting field is stored in the memory and updated to the latest value every sampling.

  The countdown signal generator 67 predicts a time T2 when the accumulated dose in the lighting field reaches the target dose TH by extrapolation, and generates a countdown signal representing the remaining time TR of the difference between the time T2 and the current time T1. The generated countdown signal is transmitted to the imaging control device 13 (step S190). A series of operations from step S150 to step S190 is continued until the remaining time TR becomes zero (YES in step S200). However, the suitability determination in step S160 is not performed after the cycle suitability determination unit 70 determines that the sampling period SP is appropriate.

  The imaging control device 13 periodically receives a countdown signal from the electronic cassette 12. In the imaging control device 13, an irradiation stop signal is generated at a time when the remaining time TR becomes zero based on the latest countdown signal. The generated irradiation stop signal is transmitted to the radiation source control device 16. When the radiation source control device 16 receives the irradiation stop signal, the X-ray irradiation from the X-ray source 15 is stopped. If the radiation source control device 16 does not receive the irradiation stop signal, the X-ray irradiation is stopped after the backup time has elapsed.

  The irradiation stop signal is generated based on the remaining time TR obtained based on the dose signal having a high S / N ratio, and the X-ray irradiation is stopped by this irradiation stop signal, so that the accumulated dose of X-rays has reached the target dose. Sometimes X-ray irradiation can be stopped accurately, and the accuracy of AEC can be improved.

  When the remaining time TR becomes zero (YES in step S200), the electronic cassette 12 ends the accumulation operation and the dose sampling operation and starts an image reading operation (step S210). In the image reading operation, a gate pulse is generated from the gate driver 50, the signal charges of the normal pixels 40a for one row are read to the signal line 42, and the image signals for one row corresponding to the signal charges are integrated amplifier 55. The operation that is taken out of the data and recorded in the memory 52 after A / D conversion is repeated at a predetermined cycle. As a result, an image signal representing one X-ray image is recorded in the memory 52. The X-ray image is subjected to various correction processes by the control unit 54 and then transmitted to the imaging control device 13 by the communication unit 33. The X-ray image is further transferred from the imaging control device 13 to the console 14 and displayed on the display 14a for diagnosis. This completes one X-ray imaging.

  As described above, according to the present invention, when the X-ray dose is high, a short value is set as the sampling period SP, so that the number of times of predictive calculation of the remaining time TR is increased and the prediction calculation of the remaining time TR is performed. Can improve the accuracy. On the contrary, when the X-ray dose is low, a long value is set as the sampling period SP, and a dose signal with a high S / N ratio can be obtained.

When the X-ray dose is low, the effect of increasing the sampling period SP is as follows quantitatively. For example, in consideration of obtaining a dose signal having a signal component equivalent to the dose signal obtained by one sampling when the sampling cycle SP is 1 ms and the sampling cycle SP is 8 ms, sampling with a sampling cycle of 1 ms is performed. It is necessary to add 8 dose signals each time. In this case, the stationary noise generated in the signal processing circuit 51 for each sampling is (8) ^1/2 ≈2.83 times that of one sampling with a sampling period SP of 8 ms, and the noise of the dose signal The component increases and the S / N ratio decreases. Therefore, it is possible to obtain a dose signal with a higher S / N ratio by performing sampling once with a long sampling period SP than when performing sampling several times with a short sampling period SP and adding the dose signals for each time. . As calculated in the above example, the S / N ratio is 8 × when the sampling is performed once at the sampling period SP = 8 ms, rather than the sampling is performed 8 times at the sampling period SP = 1 ms and the dose signals are added each time. 2.83≈23 times improvement.

  Thus, when the dose of X-rays is low, a dose signal with a high S / N ratio can be obtained by increasing the sampling period SP once even if the total time is the same. In recent years, in response to the trend of stricter management of the exposure dose to the subject H, low-dose imaging in which the X-ray dose is suppressed as much as possible is increasing. When accurately performing AEC in low-dose imaging, it is important to ensure the S / N ratio of the dose signal, so that the sampling period SP is changed to an appropriate value based on the dose signal as in the present invention. Technology is effective.

  The method for recognizing the subject area is not limited to the histogram analysis exemplified above. Although the detection pixels 40b are distributed in the imaging region 43 although the number is smaller than that of the normal pixels 40a, the dose signal obtained by each detection pixel 40b is recorded in the memory 52 when viewed as the image signal of the detection pixel 40b. The dose signal for one line can be regarded as an X-ray image with low resolution. Accordingly, the subject area may be recognized by performing known image recognition such as pattern recognition on the X-ray image represented by the dose signal for one row in the cycle suitability determination unit 70.

  In this case, for example, when there is a signal value lower than a preset threshold among the dose signals of the detection pixels 40b existing in the recognized subject area, the sampling period SP is determined to be inappropriate. In this case, the threshold value may be zero, and when the threshold value is zero, the sampling period SP is set when there is even one signal value of zero among the dose signals of the detection pixels 40b existing in the recognized subject area. Judged inappropriate. The sampling period SP may be determined to be inappropriate when there are a predetermined number of signal values lower than the threshold. Further, when the signal value of the dose signal is too low to recognize the image, it may be determined that the sampling period SP is inappropriate.

  Further, the suitability of the sampling period SP may be determined by simply comparing the magnitude relationship between the signal value of the dose signal and the threshold without recognizing the subject area. For example, the cycle suitability determination unit 70 compares the signal value of the dose signal for one row from which the dummy signal is removed with a preset threshold value, and all the signal values of the dose signal for one row are less than the threshold value. May be determined to be inappropriate. Alternatively, if there is even one zero in the signal value of the dose signal for one row after removal of the dummy signal, the sampling period SP may be determined to be inappropriate.

  Thus, if it is found that the condition for determining the suitability of the sampling period SP is that the S / N ratio of the dose signal is too low to accurately perform AEC, it is necessary to change the sampling period SP to be longer than the current condition. Well, whether or not the subject area is recognized by performing histogram analysis or image recognition, the magnitude relationship between the signal value of the dose signal and a preset threshold value (first threshold value) is compared, and the dose signal determines the threshold value. The basic condition for determining that the sampling period is inappropriate when the value falls below is not changed. Note that these exemplified determination conditions may be used individually or in combination. For example, at the start of the dose sampling operation, a determination condition for comparing the signal value of the dose signal for one row from which the dummy signal is removed with a preset threshold is used, and when a certain time has elapsed from the start of the dose sampling operation. The determination conditions based on the histogram analysis exemplified in the first embodiment are used.

  In addition, when it is determined that the sampling period SP is inappropriate in one sampling, the sampling period SP is not changed immediately, but is continuously sampled at all times of a preset predetermined number of times. When it is determined that the sampling period SP is inappropriate, the sampling period SP may be changed. Rather than determining the suitability of the sampling cycle SP at a single time, it is more reliable to determine the suitability of the sampling cycle SP at a plurality of times.

[Second Embodiment]
In the first embodiment, the sampling period SP is changed so as to increase stepwise, but conversely, it may be changed so as to decrease stepwise. For example, as shown in FIG. 16, the initial value SP1 is 8 ms, the change step SPstp is ½ times (× ½), and the sampling period SP set by the period setting unit 71 is 8 ms (SP1), 4 ms (SP2). ) 2 ms (SP3), 1 ms (SP4),...

  In the present embodiment, the initial value SP1 corresponds to the upper limit value SPmax of the first embodiment. For this reason, the initial value SP1 is set to a value shorter than ½ of the recommended irradiation time TX in the imaging conditions from the console 14 as with the upper limit value SPmax of the first embodiment (SP1 <TX / 2).

  In the present embodiment, the lower limit value SPmin is set as the setting condition for the sampling period SP. The lower limit value SPmin corresponds to the initial value SP1 of the first embodiment. For this reason, the lower limit value SPmin is set to a value that is longer than the readout time TCA in which the dose signal is read from the integrating amplifier 55 and output to the memory 52, similarly to the initial value SP1 of the first embodiment (SPmin). > TCA).

  As in the first embodiment, the periodic suitability determination unit 70 removes the dummy signal, creates a histogram as shown in FIG. 17, and further identifies the VF to recognize the missing region and the subject region. . Then, for example, on the left side of the VF recognized as the subject area, the magnitude relationship between the signal value VHmin whose number is the smallest and the signal value is not zero and the preset threshold value THV2 (corresponding to the second threshold value) is compared. The cycle suitability determination unit 70 determines that the sampling cycle SP is appropriate when the signal value VHmin is equal to or less than the threshold value THV2, and conversely, when the signal value VHmin exceeds the threshold value THV2 as shown in FIG. It is determined that the period SP is inappropriate. When the signal value VHmin exceeds the threshold value THV2, it indicates that the currently set sampling period SP is too long and the sampling period SP may be set shorter.

  In the present embodiment, since the sampling period SP is changed stepwise, the condition for determining the suitability of the sampling period SP is contrary to the first embodiment, and the S / N ratio of the dose signal is more than necessary. However, it should be understood that it is necessary to change the sampling period SP to be shorter than the current state. As a determination condition, in addition to the above example, the magnitude relationship between the signal value of the dose signal for one row and the threshold value set to the signal value immediately before the CA accumulation amount reaches the saturation limit amount is compared. When the signal value of the dose signal for the row exceeds the threshold value, the sampling period SP may be determined to be inappropriate. As the threshold value, for example, 900 is set when the signal value is expressed by 1024 gradations of 0 to 1023. At this time, the subject region is recognized by histogram analysis or image recognition, and the signal value of the dose signal of the detection pixel 40b existing in the recognized subject region and the threshold value set to the signal value immediately before the CA accumulation amount reaches the saturation limit amount You may compare the magnitude relationship with. In any case, the basic condition for comparing the magnitude value between the signal value of the dose signal and a preset threshold value and determining that the sampling period is inappropriate when the dose signal exceeds the threshold value remains the same.

  As a result of changing the sampling cycle SP stepwise, the cycle setting unit 71 determines that the cycle suitability determination unit 70 determines that the sampling cycle SP is inappropriate when the sampling cycle SP reaches the lower limit SPmin. Also, the lower limit SPmin is set without changing the sampling period SP below that.

  In the present embodiment, the sampling period SP is shortened in a stepwise manner, contrary to the first embodiment, so that when the X-ray dose reaching the imaging region 43 is relatively low, the sampling period SP is appropriately set at an early stage. Can be set to any value.

[Third embodiment]
It is possible to select the first change mode for increasing the sampling period SP stepwise shown in the first embodiment and the second change mode for decreasing the sampling period SP stepwise shown in the second embodiment. May be provided. In this case, each change mode is switched by the control part 54 according to the manual operation from the console 14, for example. The control unit 54 functions as a mode setting unit. In this way, each change mode can be used properly according to the shooting conditions. The imaging conditions include an imaging condition, a body thickness of the subject H, and the like in addition to an irradiation condition that determines an irradiation amount of the X-ray source 15. Each change mode can be appropriately switched depending on the imaging region and the body thickness of the subject H.

  For example, when the tube current or tube voltage set in the radiation source control device 16 is low and the dose is relatively low, the body thickness of the subject H is thick, or the imaging region where the body thickness of the subject H is relatively thick is selected. When the imaging condition is such that the X-ray arrival dose to the imaging region 43 is extremely low and the signal value of the dose signal is lower than the reference, the mode is switched to the second change mode. If the shooting condition is higher than the reference, the first change mode is selected. In this way, when the optimum value of the sampling period SP is expected to be relatively long, the second change mode is set, and in the opposite case, the first change mode is set, and the X-ray doses reaching the imaging region 43 are compared. The sampling period SP can be set to an appropriate value at an early stage both when the target is low and when it is high.

  In addition, each change mode may be automatically switched according to the shooting conditions set in the console 14. Specifically, a table that records the correspondence between the shooting conditions and each change mode is stored in the internal memory 72 of the control unit 54 in advance. The control unit 54 functions as a mode setting unit, and sets a change mode suitable for the photographing condition received from the console 14 among the change modes with reference to the table. The control unit 54 sets the second change mode in the case of an imaging condition in which the X-ray arrival dose to the imaging region 43 is relatively low, and conversely, the imaging condition in which the X-ray arrival dose to the imaging region 43 is relatively high. In this case, the first change mode is set.

  In the first and second embodiments, the change step SPstp is doubled or halved. However, in order to make the sampling period SP closer to the optimum value, it is preferable to further set the change step SPstp. . For example, assuming that the initial value SP1 is 1 ms and the change step SPstp is 0.5 ms (+0.5), the sampling period SP is changed to 1 ms, 1.5 ms, 2 ms, 2.5 ms, 3 ms,. In the first embodiment, for example, when the sampling period SP is 4 ms, it is determined to be inappropriate, and when it is determined to be appropriate when the sampling period SP is 8 ms, the optimum value of the sampling period SP is 4 ms and 8 ms. However, since the sampling period SP that is actually set is 8 ms, it is closer to the optimum value than when the sampling period SP is 1 ms or 2 ms, but there is still a difference from the optimum value. is there. If the setting of the change step SPstp is made finer, the sampling period SP can be brought closer to the optimum value.

[Fourth embodiment]
In the third embodiment, one of the first change mode in which the sampling period SP is increased stepwise and the second change mode in which the sampling period SP is reduced stepwise is selected and selected during the dose sampling operation. However, in this embodiment, the sampling period SP is brought closer to the optimum value by switching each change mode during the dose sampling operation.

  For example, at the start of the dose sampling operation, the first change mode in which the sampling period SP is increased stepwise is used. At this time, the setting of the change step SPstp is made relatively coarse, and the interval of the sampling period SP to be changed stepwise is increased. Then, whether or not the sampling period SP is appropriate is determined and changed in the same manner as in the first embodiment, and when it is determined that the sampling period SP is appropriate in the first change mode, the appropriateness determination in the second change mode is further performed. . Even in the suitability determination in the second change mode, if it is determined that the sampling period SP is appropriate, the set sampling period SP is not changed.

  On the other hand, when it is determined that the sampling period SP is inappropriate in the suitability determination in the second change mode, the sampling period SP is changed to be shortened by switching to the second change mode. At this time, the change step SPstp is set to be finer than that in the first change mode, and the interval of the sampling period SP to be changed stepwise is shortened. This mode switching is performed at least once, and the dose sampling operation is performed at the sampling period SP finally determined to be appropriate.

  More specifically, as shown in FIG. 18A, for example, the initial value SP1 is set to 5 ms and the change step SPstp is set to 5 ms (+5), and the search for the sampling period SP is started in the first change mode. It is assumed that the sampling period SP is changed so as to be increased in steps of 5 ms, 10 ms,..., And it is determined that the sampling period SP is appropriate in the suitability determination in the first change mode at 10 ms, for example. In this case, the optimum value of the sampling period SP exists between 5 ms and 10 ms. Here, whether or not the second change mode is appropriate is determined. If 10 ms is determined to be inappropriate, the change step SPstp is set to 1 ms (−1) as shown in FIG. 18B, and the second change mode is set. Execute. The sampling period SP is changed from 10 ms to 9 ms, 8 ms, etc. so as to be shortened stepwise. For example, when it is determined as appropriate in the determination of suitability in the second change mode at 7 ms, the sampling period SP is changed. The subsequent dose sampling operation is performed after setting to 7 ms. As described above, if each mode is executed alternately to gradually narrow the range of the sampling period SP where the optimum value exists, the sampling period SP can be made closer to the optimum value.

  In addition, when the dose sampling operation is started, the first change mode is set, and when it is determined to be appropriate at the stage of the initial value SP1, the optimum value of the sampling period SP exists below the initial value SP1. Therefore, the sampling period SP is searched by switching to the second change mode immediately and changing the sampling period SP stepwise to a value shorter than the initial value SP1. Conversely, when the dose sampling operation is started, the second change mode is set, and when it is determined that the initial value SP1 is appropriate, the optimum value of the sampling period SP is greater than or equal to the initial value SP1, Switching to the 1 change mode, the sampling period SP is changed stepwise to a value longer than the initial value SP1, and the sampling period SP is searched. In this way, the sampling period SP can be made closer to the optimum value regardless of the setting of the initial value SP1.

[Fifth Embodiment]
In the first and second embodiments, the initial value SP1 and the change step SPstp, which are the setting conditions of the sampling period SP, are fixed values, but these may be changed. In this case, the initial value SP1 and the change step SPstp are changed by the control unit 54 by manual operation from the console 14, for example. The control unit 54 functions as a setting condition changing unit. In this way, the initial value SP1 and the change step SPstp can be set to appropriate values according to the shooting conditions.

  The change step SPstp may be changed according to the result of the suitability determination of the sampling period SP by the cycle suitability determination unit 70. For example, when the dose signal and the threshold value are slightly different from each other by comparing the dose signal and the threshold value, and it can be determined that the sampling cycle SP is appropriate in a short time, the current sampling cycle SP is close to the optimum value. Therefore, the change step SPstp is set to a fine value. Conversely, if the dose signal and the threshold value are far from each other, the current sampling cycle SP is far from the optimum value, so the change step SPstp is changed to a coarse value. In this way, the sampling period SP can be brought closer to the optimum value, and the sampling period SP can be set to an appropriate value at an early stage.

  When changing the initial value SP1, for example, when the tube current or tube voltage set in the radiation source control device 16 is low and the dose is relatively low, the body thickness of the subject H is thick, or the body thickness of the subject H In the case of an imaging condition in which the X-ray arrival dose to the imaging region 43 is extremely low and the signal value of the dose signal is lower than the reference, such as when a relatively thick imaging region is selected, the initial value SP1 is compared. Set to a long value. For example, when the default initial value SP1 is 1 ms, the initial value SP1 is changed to 5 ms, for example. Conversely, in the case of an imaging condition in which the signal value of the dose signal is higher than the reference, the initial value SP1 is set to a shorter value than in the case of the imaging condition in which the signal value of the dose signal is lower than the reference. In this way, when the optimum value of the sampling period SP is expected to be relatively long, the initial value SP1 is changed to be longer, and in the opposite case, if the initial value SP1 is changed to be shorter, the initial value SP1 is close to the optimum value. Therefore, the sampling period SP can be set to an appropriate value at an early stage both when the X-ray dose reaching the imaging region 43 is relatively low and high. Note that the change step SPstp in this case is preferably set to a fine value because the initial value SP1 is close to the optimum value.

  The initial value SP1 may be automatically changed according to the shooting conditions. Specifically, a table in which the correspondence relationship between the shooting conditions and the initial value SP1 is stored in the internal memory 72 of the control unit 54 in advance. The control unit 54 functions as a setting condition changing unit, and sets an initial value SP1 corresponding to the imaging condition received from the console 14 with reference to the table. The control unit 54 sets the initial value SP1 to a relatively long value when the X-ray arrival dose to the imaging region 43 is relatively low, and conversely the X-ray arrival dose to the imaging region 43 is In the case of relatively high shooting conditions, the initial value SP1 is set to a relatively short value.

  You may change the backup time which is the irradiation time set to the radiation source control apparatus 16 when using an AEC function according to the imaging | photography site | part or the body thickness of the to-be-photographed object H. FIG. When the body thickness of the subject H is thick, or when an imaging region where the body thickness of the subject H is relatively thick is selected, the backup time is longer than when the body H is not so. In this case, referring to the backup time, it is possible to roughly determine whether the signal value of the dose signal is higher or lower than the standard. Therefore, the mode switching of the third embodiment and the initial value SP1 of the fifth embodiment are set. The backup time may be used as a determination criterion for change. For example, when the backup time is long, the signal value of the dose signal is expected to be lower than the reference, so the mode is switched to the second change mode, and the initial value SP1 is set to a long value.

  Note that the information on the body thickness of the subject H may be included in the imaging conditions transmitted from the console 14 to the electronic cassette 12, and the body thickness of the subject H may be set to, for example, “obesity type”, “normal”, “lean type”. These operation buttons may be provided on the electronic cassette 12 or the imaging control device 13 and may be input by the operator before imaging as imaging conditions different from the imaging conditions input on the console 14.

  In the first and second embodiments, the upper limit SPmax of the sampling period SP is set to 1 / of the recommended irradiation time TX among the imaging conditions from the console 14 because the dose signal needs to be sampled at least twice. Although a value shorter than 2 is set, the present invention is not limited to this. In order to further increase the accuracy of the extrapolation calculation, it is preferable to sample the dose signal two or more times. For example, in order to sample the dose signal at least three times, the upper limit value SPmax is set to 1 of the recommended irradiation time TX. A value shorter than / 3 may be set (SPmax <TX / 3).

  In addition, when the irradiation start request signal is received from the radiation source control device 16 or when the irradiation permission signal is transmitted to the radiation source control device 16, the zero point is used as a sampling point for extrapolation calculation. For extrapolation, in addition to this zero point, at least one more line signal should be sampled and one more point sampling point should be added, so the upper limit SPmax is shorter than the recommended irradiation time TX. There should be (SPmax <TX). However, since there is a slight time lag from when the detection panel 30 starts the accumulation operation and the dose sampling operation in response to the irradiation start request signal as described above, the X-ray irradiation actually starts. When extrapolation is performed only at the zero point that has received the irradiation start request signal from the source control device 16 or the zero point that has transmitted the irradiation permission signal to the radiation source control device 16 and one more sampling point, the remaining time TR It is also conceivable that the reliability of the system is lowered. Therefore, when the time point when the irradiation start request signal is received from the radiation source control device 16 or the time point when the irradiation permission signal is transmitted to the radiation source control device 16 is used as a sampling point for extrapolation calculation, preferably at least twice. It is better to set the upper limit SPmax so that the dose signal can be sampled twice or more.

[Sixth Embodiment]
When the remaining time TR until the cumulative dose of X-rays reaches the target dose is predicted and calculated as in the first embodiment, it is better to predict and calculate the remaining time TR until immediately before the cumulative dose actually reaches the target dose. This is preferable because the reliability of the remaining time TR is increased. However, depending on the setting of the sampling period, sampling may be performed just after the remaining time TR becomes zero, and this sampling may not be utilized in the prediction calculation of the remaining time TR. Therefore, in this embodiment, the sampling period SP is changed so that sampling is performed just before the accumulated dose actually reaches the target dose.

  Specifically, as shown in FIG. 19A, the sampling period SP is set to a value longer than the remaining time TR, and the sampling point indicated by the dotted circle is not used for the prediction calculation of the remaining time TR as it is. In this case, the cycle setting unit 71 sets the sampling cycle SP to a value equal to or less than the remaining time TR as shown in FIG. At this time, in consideration of the delay time TD required for the prediction calculation of the remaining time TR, the generation and transmission of the countdown signal, and the generation and transmission of the irradiation stop signal, the time TD from the time T2 when the remaining time TR becomes zero The sampling period SP is set so that sampling is performed at least once before. Therefore, the sampling period SP is a value obtained by subtracting the delay time TD from the remaining time TR.

  As a result, the sampling period SP becomes a shorter value than the one determined to be appropriate by the period appropriateness determination unit 70 and the S / N ratio of the dose signal is lowered, but immediately before the accumulated dose actually reaches the target dose. It is preferable to perform sampling to predict the remaining time TR because it contributes to the accuracy of AEC.

  In the first embodiment, the countdown signal is transmitted from the communication unit 33 to the imaging control device 13 every time the countdown signal generation unit 67 generates a countdown signal. Therefore, the sampling period SP of the dose signal and the communication interval of the countdown signal However, the present invention is not limited to this. The communication interval of the countdown signal may be determined regardless of the sampling period SP of the dose signal. In this case, the upper limit value SPmax is preferably set to a value shorter than the communication interval of the countdown signal. When the sampling period SP is set to the upper limit value SPmax, if the upper limit value SPmax is equal to or greater than the communication interval of the countdown signal, the countdown signal having the same content must be repeatedly transmitted. If the value is shorter than the communication interval, the latest countdown signal can always be transmitted.

  In consideration of the fact that the X-ray irradiation time is never longer than the backup time set in the radiation source control device 16, the upper limit SPmax is set to a value shorter than the minimum backup time. Good. Further, since the CA accumulation amount has a saturation limit amount, the upper limit value SPmax needs to be set to a value at which the CA accumulation amount does not reach the saturation limit amount.

[Seventh Embodiment]
You may provide the change mode which changes sampling period SP like the said each embodiment, and the fixed mode which does not change sampling period SP. Depending on the hospital or operator, the accuracy of AEC may not be required so much. In such a case, the fixed mode is selected and X-ray imaging is performed. Can be used according to user's preference.

  In each of the above embodiments, two or more sampling periods SP to be changed are prepared in advance. However, an aspect in which the sampling period SP is changed only by the initial value SP1 and the other two stages is also included in the present invention.

  In the first embodiment, the countdown signal generator 52 generates the interpolation line L1 and the extrapolation line L2 by linear interpolation and predicts the remaining time TR. However, the interpolation is performed by nonlinear interpolation using a polynomial or the like. The line L1 and the extrapolation line L2 may be generated.

  Instead of predicting the remaining time TR, an irradiation stop signal may be issued from the electronic cassette 12 to stop X-ray irradiation when the accumulated dose exceeds the target dose TH. In this case, if the sampling period SP is fixed at a relatively long setting, and the sampling timing does not match the time when the accumulated dose actually reaches the target dose, the accumulated dose will be the target unless the next sampling is waited. The X-ray irradiation stop timing is delayed because the dose TH is not exceeded. However, in the present invention, if a short sampling period is determined to be appropriate, a short sampling period is set, so that the sampling period SP is set to be relatively long. There is little possibility that the X-ray irradiation stop timing will be delayed as compared with the case of being fixed.

  In the first embodiment, the case where the daylighting field is set in advance according to the shooting conditions is illustrated. However, as a method for determining the daylighting field, an arbitrary part of the imaging region 43 is designated as the daylighting field by an operator setting. Alternatively, the representative value of the dose signal may be integrated for each block obtained by dividing the imaging region 43 into a plurality of blocks, and the block having the lowest integrated value may be set as the daylighting field. Alternatively, the dose signal may be regarded as an X-ray image with a low resolution, and a region of interest may be identified by applying pattern recognition to the X-ray image represented by the dose signal, and this may be used as a lighting field.

  In each of the embodiments described above, histogram analysis and image recognition technology have been exemplified as the subject region recognition method. However, the subject region may be recognized based on information on the imaging region in the imaging conditions. For example, when the imaging region is the abdomen, the above-described suitability determination is performed by regarding almost the entire region except the end of the imaging region as the subject region. Further, the suitability of the sampling period SP may be determined based on the signal value of the dose signal of the detection pixel 40b existing in the lighting field, not in the subject area.

  In the first embodiment, the countdown signal indicating the remaining time until the accumulated dose reaches the target dose is transmitted from the electronic cassette 12 to the imaging control device 13. However, the dose signal and the accumulated dose itself are imaged from the electronic cassette 12. You may transmit to the control apparatus 13. In this case, the imaging control device 13 calculates the remaining time based on the received dose signal and information on the accumulated dose, and generates an irradiation stop signal.

  In the first embodiment, the imaging control device 13 generates the irradiation stop signal based on the countdown signal transmitted from the electronic cassette 12, but the countdown signal transmitted from the electronic cassette 12 is transmitted from the imaging control device 13 to the radiation source. It may be transferred to the control device 16 and the radiation source control device 16 may generate an irradiation stop signal. Further, the electronic cassette 12 may perform the generation of the irradiation stop signal.

  The imaging control device 13 and the console 14 may be separate as in the first embodiment or may be integrated.

  In the first embodiment, the detection panel in which the scintillator is arranged on the X-ray incident side of the light detection board is illustrated, but the detection panel in which the scintillator is arranged on the opposite side to the X-ray incidence side of the light detection board is used. May be. In this case, the scintillator absorbs X-rays transmitted through the light detection substrate to generate visible light, and the light detection substrate photoelectrically converts the visible light to generate a signal charge.

  In the first embodiment, an indirect conversion type detection panel that converts X-rays into visible light and converts the visible light into signal charges is illustrated, but X-rays are directly converted by a photoconductive film such as amorphous selenium. A direct conversion type detection panel that converts to signal charge may be used.

  In the first embodiment, the sequential pixel reset operation for resetting the normal pixels 40a row by row is performed. However, the pixel reset operation may be a parallel reset method or an all pixel reset method.

  In the first embodiment, the detection pixel 40b in which the source electrode and the drain electrode of the TFT 45 are short-circuited by the short-circuit line 46 is illustrated. However, as the detection pixel 40b, for example, for image signal readout and dose signal readout You may use what connected two TFT. In this case, one of the TFTs is connected to the scanning line 41 and functions as an image readout TFT, like the TFT 45 of the first embodiment. The other TFT is connected to a scanning line for reading out a dose signal different from the scanning line 41. A dose signal readout gate driver is provided so that the two TFTs can be controlled independently, and the other TFT is connected to this via a dose signal readout scanning line. The scanning line for reading out the dose signal is wired in the row where the detection pixel 40b is located.

  In this case, a gate pulse is always given to the other TFT from the gate driver for reading out the dose signal so that the other TFT is always turned on, and the same state as that of the detection pixel 40b of the first embodiment is set. Alternatively, a gate pulse is sequentially applied to the dose signal readout scanning line from the dose signal readout gate driver, and the other TFTs are sequentially turned on one by one. Each time, the signal charge stored in the integrating amplifier 55 is output to the memory 52 as a dose signal. In the latter sampling period SP, the other TFT is turned off and charge accumulation in the detection pixel 40b is started, and then the other TFT is given a gate pulse to output the accumulated charge in the detection pixel 40b to the signal line 42. In other words, this is the charge accumulation period of the detection pixel 40b. In this case, dose signals from all the detection pixels 40 b are recorded in the memory 52 when a gate pulse is given to the scanning line for reading the dose signal in each row. The AEC unit 53 reads the dose signal from all the detection pixels 40b from the memory 52 every time it is recorded in the memory 52, performs dose signal selection, dose signal integration, and generation of a countdown signal. Similarly, every time the dose signal from all the detection pixels 40 b is recorded in the memory 52, the cycle suitability determination unit 70 reads this from the memory 52 and determines the suitability of the sampling cycle SP.

  The arrangement of the detection pixels shown in FIG. 5 is an example, and other arrangements may be used. For example, the detection pixels may be arranged at a constant pitch in the X and Y directions so as to be evenly distributed over the entire surface of the imaging region 43, and the detection pixels may be arranged in a grid pattern. In addition, although an example in which one normal pixel is used as a detection pixel has been shown, the present invention is not limited to this. For example, a part of the photoelectric conversion unit in one pixel may be separated as a sub-pixel and used as a detection pixel. However, a dedicated detection pixel may be arranged in a gap between the pixels.

  Further, the X-ray detection unit may not be a detection pixel. For example, the detection panel may be configured with only normal pixels, and in the dose sampling operation, the dose signal may be sampled based on the leakage charge leaking from the normal pixels. In this case, all the pixels function as an X-ray detection unit. Even when the pixel has the TFT turned off, a small amount of leakage charge leaks out to the signal line 42. Since the leak charge increases in accordance with the charge accumulation amount of the pixel 40, if the leak charge is accumulated in the integrating amplifier 55 and taken out as a dose signal, it can be used for AEC.

  Further, the dose may be detected based on the current flowing in the bias line connected to a specific pixel by using the current based on the charge generated in the pixel flowing in the bias line that supplies the bias voltage to each pixel. In this case, the current detector that detects the current of the bias line is the X-ray detector. The dose sampling unit obtains a dose signal by integrating the current detected by the current detection unit.

  An X-ray detection unit may be provided around the imaging region. Alternatively, an X-ray detection unit that is completely separated and independent from the detection panel may be provided in the casing of the electronic cassette, or may be attached to the outer peripheral surface of the casing.

  In the first embodiment, since the signal processing circuit functions as a dose sampling unit and an image signal reading unit, the cost can be reduced. However, the configuration in which the signal processing circuit serves as the dose sampling unit and the image signal reading unit is not essential, and the dose sampling unit and the image signal reading unit may be performed by separate signal processing circuits. In this case, a dedicated wiring for reading a dose signal is provided separately from the signal line 42, and the dedicated sampling line and the dose sampling unit provided separately from the signal processing circuit are connected to the detection pixel.

  In the first embodiment, a TFT type detection panel is illustrated, but a CMOS (Complementary Metal Oxide Semiconductor) type detection panel may be used. In the case of the CMOS type, so-called non-destructive readout is possible, in which signal charges accumulated in the pixels are read out as signal voltages through amplifiers provided in the pixels while the signal charges are held in the pixels without flowing out to the signal lines. is there. Therefore, even during the accumulation operation, it is possible to select an arbitrary pixel in the imaging region and read the dose signal from the pixel. In the case of the CMOS type, all the pixels can function as an X-ray detection unit.

  The present invention can be modified as appropriate without departing from the spirit of the present invention, such as a combination of the above embodiments. The present invention is not limited to an electronic cassette that is a portable X-ray image detection apparatus, and may be applied to an X-ray image detection apparatus that is installed on an imaging table. Furthermore, the present invention can be applied not only to X-rays but also to other radiation such as γ rays as an imaging target.

DESCRIPTION OF SYMBOLS 2 X-ray imaging system 10 X-ray imaging apparatus 11 X-ray generator 12 Electronic cassette 30 Detection panel 40 Pixel 40a Normal pixel 40b Detection pixel 43 Imaging area 51 Signal processing circuit 53 AEC part 54 Control part 70 Period appropriateness determination part 71 Period setting Part

Claims (18)

A detection panel having an imaging region in which pixels that receive radiation emitted from a radiation generator and store signal charges representing an image signal are two-dimensionally arranged;
A radiation detector for detecting radiation;
Based on the output of the radiation detection unit, a dose sampling unit that periodically samples a dose signal according to the dose per unit time of radiation, and
An AEC unit that determines whether the cumulative dose of radiation has reached a target dose based on the dose signal;
Based on the dose signal, a cycle suitability determination unit that determines whether a sampling period of the dose signal by the dose sampling unit is appropriate for the dose, and
A radiographic image detection apparatus comprising: a cycle setting unit that changes a sampling cycle when the cycle suitability determination unit determines that the sampling cycle is inappropriate.   The radiological image detection apparatus according to claim 1, wherein the cycle suitability determination unit repeatedly performs the suitability determination of the sampling cycle until the sampling cycle is determined to be appropriate.   The radiological image detection apparatus according to claim 2, wherein the cycle setting unit changes the sampling cycle stepwise from an initial value.   The radiological image detection apparatus according to claim 3, wherein the cycle setting unit shortens the sampling cycle stepwise or stepwise.   The radiological image detection apparatus according to claim 4, wherein the cycle suitability determination unit compares the magnitude of the signal value of the dose signal with a preset threshold value, and performs a sampling cycle suitability determination. . When the sampling period is increased stepwise, the first threshold is set as the threshold,
The radiological image detection apparatus according to claim 5, wherein the cycle suitability determination unit determines that the sampling cycle is inappropriate when a signal value of the dose signal is lower than the first threshold value. When the sampling period is shortened stepwise, a second threshold is set as the threshold,
The radiological image detection apparatus according to claim 5, wherein the cycle suitability determination unit determines that the sampling cycle is inappropriate when a signal value of the dose signal exceeds the second threshold value.   The periodic suitability determination unit compares the magnitude relationship between the signal value of the dose signal of the radiation detection unit in the subject region irradiated with the radiation that has passed through the subject in the imaging region and the threshold value. The radiographic image detection apparatus of any one of Claim 5 thru | or 7. A plurality of the radiation detection units are arranged so as to be evenly distributed over the entire surface of the imaging region,
The periodic suitability determination unit recognizes a subject region irradiated with radiation that has passed through a subject out of the imaging region based on dose signals of a plurality of radiation detection units, and a radiation detection unit located within the subject region The radiological image detection apparatus according to claim 8, wherein the magnitude relationship between the signal value of the dose signal and the threshold value is compared.   10. The mode setting unit according to claim 4, further comprising: a first setting mode for increasing the sampling period in steps, and a second setting mode for setting the sampling period in steps. The radiographic image detection apparatus described.   The mode setting unit includes at least one of tube current, tube voltage, subject body thickness information, and imaging region set in the radiation generating device, and is set for each imaging and radiation to the imaging region The radiographic image detection apparatus according to claim 10, wherein the first and second change modes are set based on an imaging condition for determining a final dose. The mode setting unit sets the first change mode in the case of imaging conditions in which the arrival dose is relatively high,
On the contrary, the radiographic image detection apparatus according to claim 11, wherein the second change mode is set in the case of an imaging condition in which the arrival dose is relatively low.   A setting condition changing unit that changes a setting condition when the period setting unit sets a sampling period, an initial value of the sampling period, a changing step that defines an interval of the sampling period to be changed once, and an upper limit of the sampling period The radiological image detection apparatus according to claim 3, further comprising a setting condition changing unit that changes a setting condition including at least one of the value and the lower limit value.   The setting condition changing unit includes at least one of tube current, tube voltage, subject body thickness information, and imaging region set in the radiation generator, and is set for each imaging and is applied to the imaging region. The radiological image detection apparatus according to claim 13, wherein an initial value of a sampling period is set based on an imaging condition for determining an arrival dose of radiation.   The radiographic image detection apparatus according to claim 13, wherein the setting condition changing unit changes the changing step according to a determination result of the periodic suitability determining unit. The AEC unit predicts the time when the cumulative dose reaches the target dose based on the dose signal obtained by sampling multiple times,
The radiographic image detection according to claim 13, wherein the setting condition changing unit sets the upper limit value of the sampling period to a value shorter than ½ of the recommended irradiation time of radiation. apparatus.   The radiographic image detection apparatus according to claim 1, comprising a change mode for changing a sampling period and a fixed mode for not changing a sampling period. A detection panel having an imaging region in which pixels that receive radiation emitted from a radiation generator and store signal charges representing an image signal are two-dimensionally arranged;
A radiation detector for detecting radiation;
Based on the output of the radiation detection unit, a dose sampling unit that periodically samples a dose signal according to the dose per unit time of radiation, and
An AEC unit that determines whether the cumulative dose of radiation has reached a target dose based on the dose signal;
Based on the dose signal, a cycle suitability determination unit that determines whether a sampling period of the dose signal by the dose sampling unit is appropriate for the dose, and
In the operating method of the radiological image detection apparatus comprising a cycle setting unit for setting a sampling cycle,
An operation method of a radiological image detection apparatus, wherein the sampling period is changed by the cycle setting unit when the sampling cycle is determined to be inappropriate by the cycle suitability determination unit.

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