X-ray exposure control device, X-ray image detection apparatus, and X-ray imaging system

An X-ray exposure control device comprises: an X-ray detection element including a plurality of pixels for dose detection each detecting a dose during X-ray radiation; a region setting unit configured to set a use pixel region including pixels for use in dose detection from the plurality of pixels for dose detection during the X-ray radiation; a signal generating unit configured to generate a stop signal for stopping the X-ray radiation from an X-ray source according to the dose detected by each of the pixels for use in the dose detection within the use pixel region set by the region setting unit; and a transmission unit configured to transmit to the X-ray source the stop signal to stop the X-ray radiation as generated by the signal generating unit.

BACKGROUND OF THE INVENTION

The present invention relates to an X-ray exposure control device having the function of controlling the exposure to X-rays, an X-ray image detection apparatus including the same, and an X-ray imaging system including the same.

In medical imaging using X-rays, there have conventionally been used, in general, an X-ray film method which utilizes a screen-film radiographic system in which a fluorescent screen is combined with an X-ray film and which involves directly recording an X-ray image on an X-ray film and developing the recorded X-ray image; a computed radiography (CR) method which involves recording an X-ray image on a storage phosphor sheet called imaging plate (IP) as a latent image and then reading photostimulated luminescence through laser scanning to acquire X-ray image data as digital data; and a digital radiography (DR) method which involves directly and instantaneously reading an X-ray image with an X-ray detection element such as a flat panel detector (FPD) having an X-ray sensitive layer disposed on a thin film transistor (TFT) substrate to directly acquire X-ray image data as digital data.

Any X-ray imaging system adopting any of the foregoing methods is provided with an automatic exposure control (AEC) mechanism to stop X-ray radiation in a case where X-ray radiation from an X-ray source is detected and a proper X-ray radiation dose is reached. Since the foregoing X-ray imaging system is provided with the AEC mechanism, an X-ray image of proper density can be acquired at all times in the same radiographic environment even in radiographing a variety of different sites.

Such conventional AEC is called phototimer using a so-called ionization chamber or an ion chamber having a photoelectric conversion element.

A conventional X-ray imaging system having such conventional AEC is shown inFIGS. 21A, 21B and 21C.

As shown in these drawings, an X-ray imaging system200includes an X-ray source202, a dedicated device for X-ray image detection (hereinafter referred to as “X-ray detection device”)206provided so as to be opposed to the X-ray source202and receiving an image of X-rays having passed through a subject204(radiographic site), a dosimeter210disposed between the position where the subject204is radiographed and the X-ray detection device206, and provided with X-ray sensors208at a plurality of positions constituting a lighting field, and an AEC section212which controls the stop of the X-ray source202according to the X-ray integrated dose (exposure dose) as detected by the X-ray sensors208in the lighting field of the dosimeter210. Here, the dosimeter210and the AEC section212constitute the AEC mechanism. A plurality of X-ray sensors, and in the illustrated case, three X-ray sensors208(suffixed by R, G and B symbols) are attached to the dosimeter210. Since the X-ray sensors are fixed with respect to the X-ray detection device206, the lighting field is fixed. For instance, the lighting field includes “blue (B)” and “green (G)” corresponding to the lung field in the front chest and “red (R)” in the abdomen.

In the X-ray imaging system200as described above, X-rays are radiated from the X-ray source202toward the subject204(radiographic site); the X-rays radiated to the lighting field of the subject204are detected by the X-ray sensors208of the dosimeter210; the detection signals are integrated in the AEC section212; when the X-ray dose detected by the X-ray sensors208and integrated in the AEC section212reaches an X-ray radiation dose (exposure dose) suitable to the subject204, a stop signal Sp for stopping the X-ray source202is generated in the AEC section212and transmitted from the AEC section212to the X-ray source202to stop the X-ray source202.

In recent years, in order to reduce the dose loss due to the dosimeter210shown inFIG. 21Band to reduce the cost (cut down the cost) involved in separately providing the dosimeter210, an attempt is also made to integrate the dosimeter210with the X-ray detection device206(see JP 7-2014901 A and JP 2011-174908 A (hereinafter referred to as Patent Literatures 1 and 2)).

In Patent Literatures 1 and 2, some pixels of the X-ray image detection device are used as pixels for detecting the X-ray dose.

SUMMARY OF THE INVENTION

In the meanwhile, also including the techniques disclosed in Patent Literatures 1 and 2, in the AEC of the conventional X-ray imaging system200, the lighting field (X-ray sensors208) is fixed as described above and hence is selected in advance according to the subject204(radiographic site). However, the AEC had a problem in that if the lighting field is deviated from the position of the subject, exposure cannot be properly controlled, whereby an X-ray image of proper density cannot be acquired.

Accordingly, these techniques had a problem in that an X-ray technologist needs to radiograph after the subject204(radiographic site) is positioned in advance, for example, at the positions of “blue (B)” and “green (G)” corresponding to the lung field in the front chest and at the position of “red (B)” in the abdomen so as to coincide with the lighting field (positions of the X-ray sensors208).

The present invention has been made to solve the above-described prior art problems and aims at providing an X-ray exposure control device, an X-ray image detection apparatus and an X-ray imaging system which are capable of recognizing and determining the lighting field of a radiographic subject during the radiography from an image in the course of X-ray photography, of stopping X-ray radiation at a proper exposure dose (exposure) suitable to the subject, that is, the radiographic site, in other words, of properly controlling the radiation dose during the X-ray photography according to the subject, and of acquiring an X-ray image of suitable density at all times in the same radiographic environment even in the radiography of a variety of different sites.

In order to achieve the above-described object, a first aspect of the present invention provides an X-ray exposure control device which controls a dose of X-ray radiation from an X-ray source to a radiography target, more specifically, an X-ray exposure control device which is used in an X-ray image detection apparatus to detect an X-ray image of a radiography target exposed to X-ray radiation from an X-ray source and which controls an accumulated dose of the X-ray radiation received by the radiography target, the X-ray exposure control device comprising: an X-ray detection element including a plurality of pixels for dose detection each detecting a dose during the X-ray radiation; a region setting unit configured to set a use pixel region including pixels for use in dose detection from the plurality of pixels for dose detection during the X-ray radiation; a signal generating unit configured to generate a stop signal for stopping the X-ray radiation from the X-ray source according to the dose detected by each of the pixels for use in the dose detection within the use pixel region set by the region setting unit; and a transmission unit configured to transmit to the X-ray source the stop signal to stop the X-ray radiation as generated by the signal generating unit.

The region setting unit preferably sets the use pixel region by analyzing dose information of the plurality of pixels for dose detection at a preset timing.

Preferably, the preset timing is a preset, fixed timing or a specified timing as specified from outside, and the specified timing is preferably based on at least one of a set value preset according to the radiography target, a tube current of the X-ray source and a tube voltage of the X-ray source.

Preferably, the region setting unit identifies subject pixels representing the radiography target constituting a subject or pixels within a radiation field exposed to the X-ray radiation by combining a plurality of pixel characteristics and neighboring pixel characteristics from dose information of the plurality of pixels for dose detection, and sets the use pixel region containing the subject pixels or the pixels within the radiation field as the pixels for use in the dose detection.

Preferably, the region setting unit identifies the subject pixels or the pixels within the radiation field, and sets a part of the subject pixels or a part of the pixels within the radiation field as the pixels for use in the dose detection.

The region setting unit preferably sets the use pixel region by combining a plurality of pixel characteristics and neighboring pixel characteristics from dose information of the plurality of pixels for dose detection.

The region setting unit preferably sets the use pixel region by using pixel characteristics from dose information of the plurality of pixels for dose detection.

The region setting unit preferably identifies the pixels for use in the dose detection based on pixel characteristics of a reduced image obtained by unifying the plurality of pixels for dose detection into one pixel.

The region setting unit preferably has a plurality of modes selectable according to the radiography target which was preset and sets the pixels for use in the dose detection according to a mode selected according to the radiography target.

The region setting unit preferably has a plurality of modes and sets the pixels for use in the dose detection according to a mode selected according to characteristics of an image.

The region setting unit preferably determines the selected mode based on characteristics in a subject region or a region within the radiation field.

The region setting unit preferably has a plurality of modes, and detects the pixels for use in the dose detection in the plurality of modes and determines the pixels for use in the dose detection to be set according to characteristics of an image.

Preferably, the plurality of modes include at least one mode of a first mode which sets pixels on a high dose side as the pixels for use in the dose detection in a cumulative dose histogram in a region set from the identified subject pixels, the identified pixels with the radiation field, or the plurality of pixel characteristics and the neighboring pixel characteristics; a second mode which sets pixels on a low dose side in the cumulative histogram as the pixels for use in the dose detection; and a mode which sets pixels in a vicinity of a median value in the cumulative histogram as the pixels for use in the dose detection. For example, the first and second modes may be used as a mode for radiographing the lung field and a mode for radiographing bones, respectively.

The plurality of modes preferably include a mode for specifying the use pixel region from outside or a mode for radiographing at a preset dose.

Preferably, the plurality of modes include at least one mode of a first mode which sets the pixels for use in the dose detection based on a dose of the identified subject pixels; a second mode which sets the pixels for use in the dose detection based on a dose of the identified pixels within the radiation field; a third mode which sets the use pixel region by combining the plurality of pixel characteristics and the neighboring pixel characteristics; and a fourth mode which sets the use pixel region using characteristics of an image.

The X-ray detection element preferably starts to detect, hold and accumulate the dose in each of the plurality of pixels for dose detection at a start timing at which the X-ray source starts the X-ray radiation toward the radiography target.

The X-ray detection element preferably detects the start timing with the plurality of pixels for dose detection.

Preferably, the X-ray exposure control device according to the first aspect further comprises an acquisition unit configured to acquire a start signal representing the start timing for starting the X-ray radiation from the X-ray source toward the radiography target, and the X-ray detection element starts to detect the dose in each of the plurality of pixels for dose detection according to the start signal acquired by the acquisition unit.

The acquisition unit preferably acquires the start signal from outside.

The signal generating unit preferably generates the stop signal for stopping the X-ray radiation at a point in time when the dose detected by each of the pixels for dose detection within the use pixel region has reached or exceed a preset threshold.

The threshold is preferably set based on the radiography target, radiographic conditions or a plurality of modes.

Preferably, the threshold is corrected so as to absorb differences in characteristics of the X-ray detection element or corrected so as to absorb differences in delay due to the transmission unit.

The X-ray exposure control device according to the first aspect preferably further comprises a second signal generating unit configured to generate a second stop signal for stopping the X-ray radiation from the X-ray source based on information different from the dose detected by each of the pixels for dose detection within the use pixel region.

Preferably, the information different from the dose detected is information on the radiography target, information on radiographic conditions or information on a plurality of modes.

Preferably, the X-ray exposure control device further comprises at least a notification unit configured to notify which type of signal is issued, the stop signal based on the dose detected by each of the pixels for dose detection within the use pixel region or the second stop signal based on the information different from the dose detected by each of the pixels for dose detection.

The region setting unit preferably reads out, from the X-ray detection element, dose information of the plurality of pixels for dose detection to be analyzed at a preset timing.

Preferably, after the region setting unit sets the use pixel region containing as the pixels for use in the dose detection, the signal generating unit reads out, from the X-ray detection element, the dose in each of the pixels for use in the dose detection within the use pixel region at each preset monitoring timing, compares the read-out dose with a threshold preset according to radiographic conditions and generates the stop signal at a point in time when the accumulated dose has reached or exceeded the threshold.

Preferably, the X-ray exposure control device according to the first aspect further comprises a storage unit which reads out the dose detected by each of the plurality of pixels for dose detection in the X-ray detection element at each preset sampling timing during the X-ray radiation and stores the read-out dose as dose information, and the region setting unit reads out, from the storage unit, the dose information of the plurality of pixels for dose detection to be analyzed at a preset timing.

The storage unit preferably accumulates the dose detected and read out at each preset sampling timing and stores the accumulated dose as the dose information.

Preferably, after the region setting unit sets the use pixel region containing as the pixels for use in the dose detection, the signal generating unit reads out, from the storage unit, the dose in each of the pixels for use in the dose detection within the use pixel region at each preset monitoring timing, compares the read-out dose with a threshold preset according to radiographic conditions and generates the stop signal at a point in time when the accumulated dose has reached or exceeded the threshold.

Preferably, the X-ray exposure control device according to the first aspect further comprises an accumulation unit configured to perform, for each of the pixels for dose detection, accumulation processing which includes reading out the dose detected by each of the plurality of pixels for dose detection in the X-ray detection element during the X-ray radiation at each preset sampling timing and accumulating the read-out dose, the storage unit comprises a first storage area to which the accumulation unit refers for the accumulation processing and which stores a dose accumulated for each of the pixels for dose detection and a second storage area to which the region setting unit refers for analytical processing and which stores a dose for analysis from each of the pixels for dose detection for use in the analytical processing for setting the use pixel region, and the dose accumulated for each of the pixels for dose detection in the first storage area is read out at a preset timing and stored in the second storage area as the dose for analysis for each of the pixels for dose detection.

Preferably, the accumulation unit updates the dose accumulated for each of the pixels for dose detection as stored in the first storage area by adding the dose detected by each of the pixels for dose detection in the X-ray detection element to the dose accumulated for each of the pixels for dose detection as read out from the first storage area of a storage section and the region setting unit reads out the dose for analysis for each of the pixels for dose detection as stored in the second storage area of the storage section at a preset timing, performs the analytical processing based on the read-out dose for analysis for each of the pixels for dose detection, determines the pixels for dose detection for use in generating the stop signal, and sets the use pixel region including the pixels for dose detection.

The accumulation processing in the accumulation unit and the analytical processing in the region setting unit are preferably controlled to be performed in parallel.

In order to achieve the above-described object, a second aspect of the present invention provides an X-ray image detection apparatus comprising: the X-ray exposure control device according to the first aspect; and an X-ray image detection unit configured to detect X-rays having passed through the radiography target between start of the X-ray radiation from the X-ray source and radiation stop, thereby detecting an X-ray image of the radiography target.

The X-ray image detection unit preferably comprises an X-ray image detection element including a plurality of X-ray image detection pixels for detecting the X-rays having passed through the radiography target between the start of the X-ray radiation from the X-ray source and the radiation stop.

Preferably, the X-ray image detection element is integrated with the X-ray detection element, and the plurality of pixels for dose detection have a configuration different from the plurality of X-ray image detection pixels and are incorporated between the plurality of X-ray image detection pixels or the X-ray image detection element is a non-destructive readable element and some of the plurality of X-ray image detection pixels double as the plurality of pixels for dose detection.

In order to achieve the above-described object, a third aspect of the present invention provides an X-ray imaging system comprising: an X-ray source for radiating X-rays; and the X-ray image detection apparatus according to the second aspect, wherein the X-ray source receives a start signal of the X-ray radiation from an external apparatus or the X-ray image detection apparatus to start the X-ray radiation, and receives the stop signal of the X-ray radiation from the X-ray image detection apparatus to stop the X-ray radiation.

As described above, according to the invention, it is possible to perform consistent X-ray exposure control regardless of the positioning of a radiographic subject by recognizing and determining the lighting field of the subject during the X-ray photography.

Accordingly, the present invention is capable of stopping X-ray radiation at a proper exposure dose (exposure) according to the subject (radiographic site), in other words, of properly controlling the radiation dose during the X-ray photography according to the subject, and of acquiring an X-ray image of suitable density at all times in the same radiographic environment even in the radiography of a variety of different sites.

In other words, the present invention is capable of consistent radiography at a proper dose regardless of the position of the subject or its radiographic site or of the position of the subject in the whole body.

DETAILED DESCRIPTION OF THE INVENTION

An X-ray exposure control device having the function of controlling the exposure to X-rays, an X-ray image detection apparatus including the same, and an X-ray imaging system including the same according to the present invention are described below in detail with reference to preferred embodiments shown in the accompanying drawings.

FIG. 1is a schematic diagram schematically showing a configuration of an exemplary X-ray imaging system which uses an X-ray image detection apparatus provided with an X-ray exposure control device according to a first embodiment of the invention.

As shown in this drawing, the X-ray imaging system10according to the first embodiment of the invention includes an X-ray source12and the X-ray image detection apparatus14. The X-ray image detection apparatus14includes an X-ray image detection device (hereinafter referred to simply as “image detection device”)18which is provided at a position opposed to the X-ray source12and which receives an image of X-rays having passed through a subject16(radiographic site) and a control unit20which controls the whole operation of the X-ray imaging system10including the operation control of the X-ray source12and the image detection device18and image processing of an X-ray image.

Although not shown, the X-ray imaging system10includes a radiographic table such as an upright radiographic table for radiographing the subject16at a standing posture or a decubitus radiographic table for radiographing the subject16at a lying posture, and a radiation source moving apparatus for setting the X-ray source12in a desired direction and at a desired position.

Although described later in detail, according to the example shown inFIG. 3, in the X-ray image detection apparatus14, pixels for exposure control76except a portion of normal pixels for image detection44in the image detection device18and each component of the control unit20except a portion where an X-ray image from the normal pixels44is processed mainly constitute the X-ray exposure control device according to the first embodiment of the invention.

The X-ray source12includes an X-ray tube for X-ray radiation and a radiation field limiter (collimator) for limiting the radiation field of the X-ray radiation from the X-ray tube. The X-ray tube has a cathode composed of a filament emitting thermoelectrons and an anode (target) which the thermoelectrons emitted from the cathode strike to cause X-ray radiation. The radiation field limiter has, for example, a plurality of lead sheets for blocking out X-rays which are disposed in a curb shape so that a radiation opening for passing X-rays therethrough is formed in the center, and the size of the radiation opening is changed through positional movements of the lead sheets to limit the radiation field.

As shown inFIG. 2, the control unit20includes an X-ray detection controller (hereinafter referred to simply as “detection controller”)22comprehensively controlling the whole operation of the apparatus, and a high voltage generator24, a radiation switch26, an input device28, a display30and a memory32connected to the detection controller22.

The detection controller22includes a device controller34, a radiation source controller36, a memory38and an X-ray exposure controller (hereinafter also referred to as “AEC section”)40.

The high voltage generator24increases the input voltage using a transformer to generate a high tube voltage and supplies the generated high tube voltage to the X-ray source12through a high voltage cable. The radiation source controller36controls the tube voltage which determines the energy spectrum of X-ray radiation from the X-ray source12, the tube current which determines the radiation dose per unit time, the start of radiation and the stop or termination of radiation from the X-ray source12, and the X-ray radiation time.

The radiation switch26is, for example, a two-stage push switch operated by an operator such as a radiation technologist. One-stage pressing generates a warm-up start signal for causing the X-ray source12to start warm-up and two-stage pressing generates a radiation start signal for causing the X-ray source12to start radiation. These signals are input to the radiation source controller36through a signal cable.

The radiation source controller36causes supply of electric power from the high voltage generator24to the X-ray source12to be started upon receipt of a radiation start signal from the radiation switch26and causes the supply of electric power from the high voltage generator24to the X-ray source12to be stopped to terminate X-ray radiation from the X-ray source12upon receipt of a radiation stop signal from the AEC section40.

The memory32stores in advance several types of radiographic conditions such as the tube voltage and the tube current. The radiographic conditions are manually set by the operator through the input device28. The radiation source controller36intends to radiate X-rays according to the product of the exposure time and the set radiographic conditions such as the tube voltage and the tube current. When it is detected that a necessary and sufficient dose has been reached, AEC in the AEC section40functions to stop the X-ray radiation even if the dose is equal to or smaller than the tube current−exposure time product (exposure time) according to which the radiation source controller36intended to radiate. In order to prevent the X-ray radiation from being finished before receiving a radiation stop signal from the AEC section40as a result of a target dose reached, thus leading to lack of dose, the maximum value of the tube current−exposure time product (exposure time is also possible) is set in the radiographic conditions of the X-ray source12. The set product of the tube current and the exposure time preferably takes a value suitable to the radiographic site.

The memory38and the AEC section40will be described later in detail.

The device controller34controls the operation of the image detection device18in response to an input operation from the operator through the input device28. More specifically, the device controller34performs various controls including power on/off of the image detection device18and mode switching to standby mode or radiographic mode.

In addition to this, the device controller34preferably has the function of performing various image processing steps such as offset correction, sensitivity correction and defect correction on X-ray image data in the memory38. These various image processing steps will be described later.

The X-ray image data from the image detection device18is stored in the memory38and then subjected to the above-described various image processing steps in the device controller34of the control unit20. The X-ray image having undergone such image processing steps is displayed on the display30or its data is stored again in the memory38or a storage device (not shown), or a data storage such as an image storage server connected to the control unit20through a network.

The control unit20has the function of a so-called console, and receives the input of a testing order including information on the gender and age of a patient, the radiographic site, purpose of radiography and the like and displays the received testing order on the display30. The testing order is input from external systems such as HIS (hospital information system) and RIS (radiography information system) which manage the patient information and the testing information on the radiographic testing, or is manually input by an operator. The testing order includes radiographic sites such as head, chest and abdomen, and radiographic directions such as front side, lateral side, oblique position, PA (X-ray radiation from the back side of a subject) and AP (X-ray radiation from the front side of a subject). The operator checks the testing order contents on the display30and inputs radiographic conditions suitable to the contents through the operation screen on the display30.

Next, the image detection device18is an X-ray detection element of the invention and includes a DR flat panel detector (hereinafter abbreviated as “FPD”)42(seeFIG. 3) and a casing containing the FPD42. The casing of the image detection device18has a substantially rectangular, flat shape, and is used to fix the FPD42to a radiographic table (not shown). Although described in detail later, the image detection device18may be an electronic cassette of a detachable and transportable cassette type. In the case of an electronic cassette, the casing containing the FPD42should be a transportable casing and its planar size should be the same as that of a film cassette or an IP cassette (also called a CR cassette) (the size according to International Standard ISO 4090:2001). If so, it is also possible to attach the device18to an existing radiographic table for a film cassette or an IP cassette. In the case of the image detection device18of an electronic cassette type, it is also possible to use the device18alone by placing it on a bed on which a subject lies or by making a subject carry it instead of setting it on a radiographic table.

The FPD42includes a TFT active matrix substrate. The imaging surface46in which a plurality of pixels44for accumulating charges according to the X-ray dose reached are arrayed is formed on top of the substrate. The plurality of pixels44are two-dimensionally arrayed at a predetermined pitch in a matrix of n rows (x direction) and m columns (y direction).

The FPD42is of an indirect conversion type which includes a scintillator (phosphor) capable of converting X-rays into visible light and which photoelectrically converts in the pixels44visible light obtained by conversion in the scintillator. The scintillator is composed of CsI:TI (thallium-activated cesium iodide), GOS (Gd2O2S:Tb; gadolinium oxysulfide) or the like, and is disposed so as to face the whole of the imaging surface46on which the pixels44are arrayed. The scintillator and the TFT active matrix substrate may be of a PSS (Penetration Side Sampling) type in which they are disposed in the order of the scintillator and the substrate when seen from the side on which X-rays enter, or be, conversely, of an ISS (Irradiation Side Sampling) type in which they are disposed in the order of the substrate and the scintillator. Use may be made of a direction conversion type FPD which does not use a scintillator but uses a conversion layer (e.g., amorphous selenium) that may directly convert X-rays into charges. Moreover, use may be made of a CMOS type instead of a TFT type.

Each pixel44includes a photodiode48which is a photoelectric conversion element that may generate charges (electron-hole pairs) in response to incidence of visible light, a capacitor (not shown) that may accumulate the charges generated by the photodiode48, and a thin film transistor (TFT)50as a switching element. It is also possible to accumulate charges in the photodiode48instead of separately providing a capacitor.

The photodiode48has a configuration including a semiconductor layer (e.g., PIN type) which may generate charges, and an upper electrode and a lower electrode which are provided above and below the semiconductor layer, respectively. In the photodiode48, the TFT50is connected to the lower electrode and a bias line52is connected to the upper electrode. Bias lines52whose number corresponds to the number of rows of the pixels44(n rows) on the imaging surface46are integrated to form a single connection. The connection53is connected to a bias power source54. A bias voltage is applied to the upper electrodes of the photodiodes48from the bias power source54through the connection53and the bias lines52. Application of a bias voltage causes an electric field in the semiconductor layer and charges (electron-hole pairs) generated in the semiconductor layer by photoelectric conversion transfer to the upper electrode and the lower electrode, one of them having a positive polarity and the other having a negative polarity. The charges having transferred are accumulated in the capacitor.

In the TFT50, a gate electrode, a source electrode, and a drain electrode are connected to a scanning line56, a signal line58and the photodiode48, respectively. The scanning lines56and the signal lines58are formed in a grid shape and the number of the scanning lines56provided corresponds to the number of rows of the pixels44(n rows) on the imaging surface46and the number of the signal lines58provided corresponds to the number of columns of the pixels44(m columns) on the imaging surface46. The scanning lines56are connected to a gate driver60and the signal lines58are connected to a signal processing circuit62.

The gate driver60drives each TFT50so that the TFT50performs the accumulating operation for accumulating signal charges in the pixel44according to the X-ray dose reached, the readout (main reading) operation for reading out the signal charges from the pixel44, and the reset (void reading) operation. A controller64controls the start timing of each of the foregoing operations executed by the gate driver60.

In the accumulating operation, the TFTs50are turned off and signal charges are accumulated in the pixels44during this period. In the readout operation, gate pulses G1to Gn which drive the TFTs50in the same rows all together are successively generated from the gate driver60to sequentially activate the scanning lines56on a row by row basis and the TFTs50connected to the scanning lines56are turned on on a row by row basis. When the TFTs50are turned on, the charges accumulated in the capacitors of the pixels44are read out to the signal lines58and are input to the signal processing circuit62.

The signal processing circuit62includes integrating amplifiers66, CDS circuits (CDS)68, a multiplexer (MUX)70, an A/D converter (A/D)72, and the like. The integrating amplifiers66are individually connected to the signal lines58. Each integrating amplifier66is composed of an operational amplifier66aand a capacitor66bconnected between the input and output terminals of the operational amplifier66a, and the signal line58is connected to one of the input terminals of the operational amplifier66a. The other of the input terminals of the operational amplifier66ais connected to ground (GND). A reset switch66cis connected in parallel to the capacitor66b. The integrating amplifiers66integrate the charges input from the signal lines58, convert them into analog voltage signals V1to Vm and output the analog voltage signals. The output terminal of the operational amplifier66ain each column is connected to the MUX70through an amplifier74and the CDS68. The output side of the MUX70is connected to the A/D72.

Each CDS68has sample-and-hold circuits and subjects an output voltage signal from the integrating amplifier66to correlated double sampling to remove noise while holding the output voltage signal from the integrating amplifier66in the sample-and-hold circuits for a preset period of time (sample holding). The MUX70uses an electronic switch to sequentially select one CDS68from the CDSs68in the respective columns connected in parallel based on an operation control signal from a shift resistor (not shown) and serially inputs the voltage signals V1to Vm output from the selected CDSs68to the A/D72. The A/D72converts the input voltage signals V1to Vm into digital voltage signals and output the digital voltage signals as image data representing an X-ray image to (the memory38and/or the AEC section40of the detection controller22of) the control unit20. An amplifier may be connected between the MUX70and the A/D72. It is also possible to provide an A/D for each signal line58, and in this case the A/Ds are followed by the MUX.

When the MUX70reads out the voltage signals V1to Vm in one row from the integrating amplifiers66, the controller64outputs a reset pulse RST to the integrating amplifiers66to turn on the reset switches66c. The signal charges in one row as accumulated in the capacitors66bare thereby discharged and the integrating amplifiers66are reset. After the integrating amplifiers66have been reset, the reset switches66care turned off again. After the lapse of a preset period of time, one of the sample-and-hold circuits of each of the CDSs68is held to sample the kTC noise component of the integrating amplifiers66. Thereafter, a gate pulse for the next row is output from the gate driver60to start readout of signal charges from the pixels44in the next row. In addition, after the lapse of a preset period of time from the output of the gate pulse, the signal charges from the pixels44in the next row are held by the other sample-and-hold circuit of each of the CDSs68. These operations are sequentially repeated to read out signal charges from the pixels44in all the rows. High-speed drive is possible by adopting pipeline processing which performs these processing steps at a time.

Outputting image data of an X-ray image in one row to the control unit20for each readout in the one row and recording the output image data in the memory38are repeatedly performed. Upon completion of readout in all the rows, image data of the X-ray image in one screen is recorded in the memory38. The X-ray image of the subject is thus detected. Another configuration is also possible in which a memory connected to the A/D72in the image detection device18is incorporated and digital image data output from the A/D72is once stored in the incorporated memory such that after image data representing an X-ray image in one screen has been stored, the image data in the one screen is immediately read out from the incorporated memory, output from the image detection device18to the control unit20and recorded in the memory38.

Dark charge occurs in the semiconductor layer of each photodiode48regardless of whether X-rays enter. The dark charge is accumulated in the capacitor of the pixel44because a bias voltage is applied. The dark charge occurring in the pixel44constitutes a noise component of image data. Thus, the reset operation is performed to remove the dark charge at preset time intervals. The reset operation is an operation for sweeping the dark charge occurring in the pixel44through the signal line58.

The reset operation is carried out by, for example, a sequential reset method in which the pixels44are reset on a row by row basis. In the sequential reset method, the gate pulses G1to Gn are sequentially issued from the gate driver60to the scanning lines56to turn on the TFTs50of the pixels44on a row by row basis, as in the readout operation of the signal charges. While the TFTs50are turned on, the dark charges flow from the pixels44through the signal lines58to the capacitors66bof the integrating amplifiers66. In the reset operation, the MUX70does not read out the charges accumulated in the capacitors66b, unlike the readout operation. A reset pulse RST is output from the controller64in synchronism with occurrence of each of the gate pulses G1to Gn to turn on the reset switches66c, whereby the charges accumulated in the capacitors66bare discharged to reset the integrating amplifiers66.

Instead of the sequential reset method, use may be made of a parallel reset method in which a plurality of rows of arrayed pixels are unified into one group, the pixels in the group are sequentially reset and the dark charges in the rows of the group are simultaneously swept and an all-pixel reset method in which gate pulses are applied to all the rows to simultaneously sweep the dark charges in all the pixels. The parallel reset method and the all-pixel reset method make it possible to accelerate the reset operation.

In addition to the normal pixels44to which the TFTs50driven by the gate driver60and the scanning lines56as described above are connected, the FPD42includes within the same imaging surface46the pixels for exposure control76which are the pixels for dose detection according to the invention and are short-circuited to the signal lines58without the TFTs50. The control pixels76are pixels used to detect the reached dose of X-rays incident on the imaging surface36after having passed through the subject16and functions as the AEC sensors for generating a radiation stop signal in the AEC section40of the detection controller22in the control unit20. The control pixels76account for about several percent of the pixels44in the imaging surface36.

The control pixels76are preferably provided, for example, along a wavy trajectory which is bilaterally symmetric with respect to the center of the imaging surface46so that these pixels are not disposed locally within the imaging surface46but are evenly scattered within the imaging surface46. One control pixel76is provided in each column of pixels44to which the same signal line58is connected, and it is preferable for one column having the control pixel76and columns (e.g., two or three columns) having no control pixel76to be alternately disposed. The positions of the control pixels76are already known at the time of manufacture of the FPD42, and the FPD42preferably stores the positions (coordinates) of all the control pixels76in advance, for example, in a non-volatile memory (not shown). Conversely, the control pixels76may be disposed in a locally concentrated manner. The arrangement of the control pixels76may be appropriately changed. In the mammography apparatus for use in imaging the breast, for instance, the control pixels76are preferably disposed so as to be concentrated on the chest wall side.

In the illustrated case, the pixels for exposure control76are disposed at the positions of the normal pixels for image detection of the FPD42so as to be substituted for the normal pixels horizontally and vertically at intervals of a few pixels. However, the present invention is not limited thereto and the pixels for exposure control76may be disposed in the space between the normal pixels. In this case, it is not necessary to use the positions of the normal pixels for the control pixels76and hence the pixel density can be correspondingly increased.

The TFT50is not provided between the control pixel76and the signal line58and the control pixel76is directly connected to the signal line58. Accordingly, the signal charge having occurred in the control pixel76is immediately read out to the signal line58. The same applies to the case where the TFTs50of the normal pixels44in the same column are turned off and the normal pixels44are in the course of accumulating operation for accumulating signal charges. Therefore, the charge having occurred in the control pixel76always flows into the integrating amplifier66on the signal line58to which the control pixel76is connected. During the accumulating operation, the charge from the control pixel76which was accumulated in the integrating amplifier66is output as a voltage value to the A/D72through the MUX70with a preset sampling period. The A/D72converts the input voltage value into a digital voltage value and output to (the AEC section40of the detection controller22of) the control unit20as pixel dose data for exposure control.

The image detection device18in the embodiment under consideration has the plurality of control pixels76and hence constitutes the X-ray detection element of the invention.

The image detection device18is basically configured as described above.

As described above, the device controller34of the control unit20is provided with circuits (not shown) having the function of performing various image processing steps such as offset correction, sensitivity correction and defect correction on X-ray image data in the memory38. The offset correction circuit removes fixed pattern noise due to individual differences and radiographic environment of the signal processing circuit62by subtracting the offset correction image acquired from the FPD42without X-ray radiation from the X-ray image on a pixel unit basis.

The sensitivity correction circuit is also called a gain correction circuit and corrects, for example, variations in the sensitivity of the photodiode48of each pixel44and variations in the output characteristics of the signal processing circuit62. The sensitivity correction is performed based on the sensitivity correction data generated based on an image obtained by subtracting the offset correction image from an image obtained by X-ray radiation at a predetermined dose in the absence of a subject. The sensitivity correction data has a coefficient for correcting the deviation from a reference value for each pixel so that each pixel output may be the same without any exception by multiplying the X-ray image after the offset correction by the sensitivity correction data upon X-ray radiation at a predetermined dose in the absence of a subject. For instance in a case where the output of Pixel A is a reference value of 1, whereas the output of Pixel B is 0.8, Pixel B has a coefficient of 1.25 (1/0.8=1.25).

The defect correction circuit linearly interpolates the pixel value of a defect pixel by the pixel values of its surrounding normal pixels based on the defect pixel information included with shipment. The pixel value of each control pixel76in the lighting field which was used to detect the dose in the AEC is also interpolated in the same manner.

The offset correction image and the sensitivity correction data are, for example, acquired at the time of shipment of the image detection device18, or acquired by a manufacturer's serviceman at the time of a periodic maintenance or by an operator during the working hours of a hospital, thereby being recorded in the internal memory of the device controller34and read out at the time of correction.

Various image processing steps may be performed by providing the above-described various image processing circuits within the detection controller22of the control unit20except the device controller34.

Next, the AEC section40of the detection controller22of the control unit20is a characteristic portion of the invention. The AEC section40recognizes and automatically determines the lighting field of the subject16based on dose data which includes digital voltage signals (dose detection signals) as detected by the control pixels76of the image detection device18, and generates a radiation stop signal Sp for stopping the X-ray radiation from the X-ray source12at a point in time when the amount of dose data as accumulated by the control pixels76in the lighting field has reached a threshold.

The AEC section40constitutes the main part of the X-ray exposure control device of the invention. As shown inFIG. 4, the AEC section includes a readout/accumulation portion78which reads out pixel dose data for exposure control (hereinafter referred to simply as “dose data”) from the image detection device18and accumulates the read-out dose data; a memory84provided with a storage area for accumulation80which stores the accumulated dose data and a storage area for analysis82which stores the dose data for analysis; a lighting field recognizing portion86which automatically recognizes the lighting field of the subject16based on the dose data for analysis stored in the storage area for analysis82of the memory84; a radiation stop determining portion88which determines based on the accumulated dose data in the lighting field as to whether radiation is stopped; a radiation stop signal generating portion90which generates the radiation stop signal Sp (first radiation stop signal Sp1) according to the determination for stopping radiation; a second radiation stop signal generating portion92which generates a second radiation stop signal (Sp2) to stop the X-ray radiation from the X-ray source12; and a transmitter94which transmits the (first and second) radiation stop signals Sp (Sp1, Sp2) to the X-ray source12through the radiation source controller36and the high voltage generator24to stop the X-ray radiation from the X-ray source12.

The readout/accumulation portion78reads out the dose data which was detected by the control pixels76of the image detection device18, acquired through the signal lines58to which the control pixels76are connected, and output after A/D conversion in the A/D72at a preset sampling timing, for example, with a preset sampling period; reads out accumulated dose data which was acquired by adding the read-out dose data in each sampling by, for example, the preset sampling timing from the start of the X-ray radiation and was stored in the storage area for accumulation80of the memory84; calculates accumulated dose data newly integrated and accumulated by adding the dose data read out at the sampling timing to the read-out accumulated dose data; and stores the resulting accumulated dose data in the storage area for accumulation80.

Here, the readout/accumulation portion78starts to measure the time that elapses before the preset timing, i.e., time period (sampling period) at a point in time when the readout/accumulation portion78receives from the radiation switch26a radiation start signal representing the start timing for starting X-ray radiation from the X-ray source12to the subject14as the radiography target. The radiation start signal from the radiation switch26is also transmitted to the image detection device18, which transfers from the reset operation to the dose detection operation upon receipt of the radiation start signal and starts detecting, holding and accumulating the dose in the control pixels76.

In the practice of the invention, the start timing for starting the X-ray radiation from the X-ray source12may be detected as a point in time when the control pixels76of the image detection device18detect the dose so that a signal of the start timing is transmitted to the readout/accumulation portion78.

The memory84stores, as accumulated dose data, the (cumulative) dose accumulated in the control pixels76of the image detection device18and includes the storage area for accumulation80which stores the cumulative dose accumulated at each sampling timing and the storage area for analysis82which stores the dose data for analysis for recognizing the lighting field.

The storage area for accumulation80of the memory84is a storage area for storing, for each control pixel76, the cumulative dose data accumulated in the control pixels76from the start of the X-ray radiation up until the sampling timing, and the stored cumulative dose data is updated at each sampling timing to new cumulative dose data in which the dose data sampled with the sampling period is accumulated. Since the normal pixels44of the FPD42in the image detection device18are of a TFT system using the TFTs50, in the case of the TFT system, once the dose data accumulated in the normal pixels44and the control pixels76is read out, the dose data accumulated in the normal pixels44and the control pixels76is reset and hence the cumulative dose data read out at each sampling timing needs to be accumulated in another memory each time the dose data is read out. If not, the whole cumulative dose data cannot be obtained. To do this, in the AEC section40to which the image detection device18is connected, the storage area for accumulation80of the memory84is necessary in order to sequentially accumulate the cumulative dose data read out from the control pixels76at each sampling timing each time the dose data is read out.

The storage area for analysis82is a storage area for storing the dose data for analysis for recognizing the lighting field which is used to automatically determine the lighting field in the lighting field recognizing portion86. The storage area for analysis82reads out the cumulative dose data stored in the storage area for accumulation80at a preset timing and stores the read-out cumulative dose data.

The cumulative dose data accumulated in the storage area for accumulation80and the dose data for analysis transferred into the storage area for analysis82preferably have different spatial resolutions and bit resolutions, and at least one of the spatial resolution and the bit resolution in the dose data for analysis is preferably lower than in the cumulative dose data. The data volume of the dose data for analysis can be thus compressed to accelerate the speed for analytical processing in the lighting field recognizing portion86based on the dose data for analysis. In other words, the dose data for analysis transferred into the storage area for analysis82is deemed to be image data from the control pixels76. Because the number of pixels is small, the image is rough but is sufficient for analytical processing for recognizing the lighting field.

The preset timing for reading out the cumulative dose data from the storage area for accumulation80as the dose data for analysis is preferably in such a state that the SN ratio (S/N) reaches a certain level so that the lighting field recognition functions consistently.

It is preferable for a timing to be applied in advance in a lot of cases of X-ray imaging of the subject14using the image detection device18to determine the preset timing as described above according to at least one of the radiographic site of the subject, and the tube current and the tube voltage of the X-ray source12, and the like, as the timing in which the S/N has reached a level suitable to make the lighting field recognition function consistently. In this case, the preset timing may be a preset fixed timing or a specified timing as specified from the exterior. In addition, the specified timing as described above may have a set value preset according to the radiographic site, or a value set based on at least one of the tube current and the tube voltage of the X-ray source12. The larger the tube current of the X-ray source12is, the more the radiation time necessary to obtain the same exposure dose can be shortened, and hence the more the timing can be shortened. On the other hand, the higher the tube voltage of the X-ray source12is, the more the dose of X-rays passing through the subject increases even at the same exposure dose, and hence the more the timing can be shortened.

The lighting field recognizing portion86performs analytical processing of an image produced by the dose data of the control pixels76in the course of X-ray radiation (each control pixel76has a value corresponding to the cumulative dose data) to determine one or more use pixels (lighting field) for use in determining the radiation stop in the radiation stop determining portion88. The lighting field recognizing portion86performs analytical processing based on the dose data for analysis as stored in the storage area for analysis82of the memory84to automatically recognize and automatically determine the lighting field of the radiographic site of the subject16to be radiographed. The lighting field recognizing portion86reads out the dose data for analysis stored in the storage area for analysis82by reference to the storage area for analysis82of the memory84in order to determine the lighting field (use pixels) through analytical processing based on the read-out dose data for analysis in each control pixel76, and the readout may be performed any time after a point in time when the dose data for analysis is stored in the storage area for analysis82, and is preferably performed just after the dose data for analysis is stored in the storage area for analysis82.

The method of recognizing the lighting field in the lighting field recognizing portion86is not particularly limited but any recognition method may be applied as long as the lighting field can be recognized as one or more use pixels or a use pixel region containing one or more use pixels and be determined automatically. For instance, methods of recognizing the lighting field as described below can be performed. The methods of recognizing the lighting field are described below.

First of all, as a first example, the control pixels76satisfying predetermined conditions can be determined as the lighting field from the statistics of the dose data values of all the control pixels76. The dose data value of the control pixel76is hereinafter simply referred to as a pixel value. The pixel as used herein refers to the control pixel76.

In other words, the pixel characteristics may be used from the pixel values (dose information) of the control pixels (pixels for dose detection)76to determine and set the lighting field.

For instance, the region excluding high density side pixels having the potential for the direct X-ray region and low density side pixels having the potential for the diaphragm region according to histogram analysis, in other words, the control pixels76having a median value of all the pixel values (40 to 60% in a cumulative histogram) are determined as the lighting field. Alternatively, the variance (σ2) of all the pixel values is determined and the control pixels76far from the average value by at least α×σ (α is a constant) are excluded to determine the remaining control pixels76as the lighting field.

The method of determining the lighting field only from the distribution of all the pixel values has a problem in that the method is likely to fail when the area ratio between the subject region, the direct X-ray region and the region outside the radiation field is excessively unbalanced and hence this method is preferably used in the case of having a standard area ratio between the subject region, the direct X-ray region and the region outside the radiation field.

Still alternatively, it is also possible to calculate a binarized threshold (e.g., center of a histogram) by histogram analysis and to determine the region having a specified size including the centers of gravity of the pixels equal to or larger than the threshold as the lighting field. Since the subject is surrounded by the direct X-ray region and the high density region such as a vicinity of the periphery of the skin, and the high density center of gravity is more likely to be within the subject, the region containing the center of gravity can be set as the lighting field. The predetermined size may be, for example, a circular region having a diameter of about 8 cm as adopted in conventional AEC or be determined based on the image size. For instance, a circular region whose diameter is a half of the image side may be applied. In addition, a high density region having the potential for the direct X-ray region and a low density region having the potential for the region outside the radiation field may be excluded from the circular region.

Next, in a second example, a plurality of characteristics, for example, a plurality of pixel characteristics or characteristics in the neighboring pixels are combined to identify and extract subject pixels, and a subject region containing some or all of the identified subject pixels or the whole of the subject region is determined as the lighting field.

For instance, the subject region can be extracted by excluding the direct X-ray region based on a histogram and excluding the region outside the radiation field based on a difference histogram as disclosed in JP 63-233658 A. Alternatively, it is also possible to use segmentation or machine learning as disclosed in JP 2004-078939 A. It is only necessary to learn three elements including subject, direct X-ray and outside the radiation field as targets to be identified, for instance, the likelihood of being a subject pixel, the likelihood of being a direct X-ray and the likelihood of being outside the radiation field.

In addition to this, it is possible to identify the subject region by learning the conditions and characteristics for discriminating the three elements (subject, direct X-ray, outside the radiation field) with the use of known machine learning methods (e.g., AdaBoost, Support Vector Machine).

It is also possible to detect a subject region and to set pixels having a median value (e.g., 30 to 70%) of pixel values in the subject region as the lighting field.

The region outside the radiation field may be excluded to determine the field within the radiation field composed of the subject region and the direct X-ray region as the lighting field instead of the subject region. In other words, for instance, the pixels within the radiation field exposed to X-ray radiation may be identified and extracted by combining the plurality of pixel characteristics or characteristics in the neighboring pixels to determine the region within the radiation field including some or all of the identified pixels within the radiation field or the whole of the region within the radiation field as the lighting field.

For instance, it is possible to detect edge candidate points along a radial linear direction set for a given point in an image by applying the radiation field recognition method disclosed in commonly assigned JP 3923131 B, to determine a preset number of reference candidate lines for these edge candidates using Hough conversion, and to determine the region surrounded by these reference candidate lines as the region within the radiation field.

In addition, a method of defining the radiation field that may be used in the apparatus for blackening a region outside the radiation field as disclosed in commonly assigned JP 3765920 B may be applied, the method including storing template information on a plurality of radiation field shapes, inputting information capable of identifying the radiation field shape of a radiation image, selecting a template corresponding to the information on the radiation field shape from the template information on the stored plurality of radiation field shapes based on the input information on the radiation field shape, performing positional and directional matching between the selected template and the radiation image to define the radiation field of the radiation image as the region within the radiation field.

As described above, it is also possible to detect a region within the radiation field and set pixels having a median value (e.g., 30 to 70%) of pixel values in the region within the radiation field as the lighting field.

As described above, it is also possible to detect a region within the radiation field and to set a region obtained by excluding high density pixels (e.g., 30% on the black side in the whole width of a histogram) having the potential for the direct X-ray region from the pixels in the region within the radiation field, or a median value (e.g., 30 to 70%) of the region as the lighting field.

In addition, the lighting field may be determined and set from pixel values of the control pixels76by combining a plurality of pixel characteristics or characteristics in the neighboring pixels.

For instance, it is possible to calculate the differential center of gravity based on the pixel values of adjoining pixels and set the region having a predetermined size including the differential center of gravity as the lighting field. The predetermined size can be considered in the same manner as in the above-described pixel center of gravity.

Furthermore, the lighting field may be identified based on the pixel characteristics of a reduced image obtained by treating pixel values of a plurality of control pixels76as one pixel.

In a third example, specific control pixels76are further extracted from the subject region and determined as the lighting field. In other words, the lighting field is determined by identifying the subject pixels and statistically analyzing the identified subject pixels.

For instance, pixels having a median value (40 to 60% in a cumulative histogram) in the subject region are determined as the lighting field. Influences in a case where the direct X-ray region or the region outside the radiation field is incorporated in the subject region can be reduced by excluding the high dose side and the lower dose side. Conversely, the high dose (high exposure) control pixels may be set as the lighting field or the low dose (low exposure) control pixels may be set as the lighting field.

Influences in a case where the direct X-ray region or the region outside the radiation field is incorporated in the subject region can also be reduced by placing more importance on the center of gravity of the subject region through a combination of barycentric positions and variance of the pixels belonging to the subject region.

It is also possible to calculate the identification result of the subject region for each pixel by multiple values and to weight it according to the degree of reliability.

Alternatively, it is also possible to prepare several selectable modes in the lighting field recognizing portion86according to the radiography target such as the radiographic site so as to switch from one to another for use.

For instance, it is preferable to prepare, for instance, Mode A (preferentially specifying high exposure pixels) in which high dose side pixels (e.g., 80 to 90% in a cumulative dose histogram) in the subject region, the region within the radiation field, or the region set from the pixel characteristics or the neighboring pixel characteristics are recognized as the lighting field; and Mode B (preferentially specifying low exposure pixels) in which low dose side pixels (e.g., 20% to 40% in a cumulative histogram) in the foregoing regions such as the subject region are recognized as the lighting field, such that the mode is switched between them by specifying from outside Mode A in the examination for observing the lung field and Mode B in the examination for observing bones. In addition to Modes A and B, Mode C (standard specification) in which the pixels having a medium dose (40 to 60% in a cumulative histogram) are recognized as the lighting field may be prepared to enable switching among the three modes.

For instance, the method of calculating the medium-dose pixels, i.e., the median value is also not limited to 40% to 60% in a cumulative histogram as described above, but a variety of calculation methods may be used. For instance, there is also a method in which the median value is set in a range of 30 to 70% in a cumulative histogram, and there is also another method in which a given rate in a cumulative histogram is taken after excluding a portion of a given rate on the high density side from the whole width of the histogram as the high density region having the potential for the direct X-ray region. The latter case has the advantage of being less likely to depend on the area of the direct X-ray region. The same applies to the low density side region and this method can be used in an application in which a protector used in, for example, radiographing the hip joint is excluded. In this case, a given rate in a cumulative histogram is taken as a median value after excluding a portion of a given rate on the low density side from the whole width of the histogram as the low density region having the potential for the protector. As a result, this method has the advantage of being less likely to be affected by the area of the protector.

The selectable modes may include, at least Mode D in which the lighting field is set by analyzing the pixel values (dose data for analysis) of the control pixels76at a preset timing during X-ray radiation, and Mode E in which the lighting field is specified from outside.

The selectable modes may include at least one mode of Mode F in which the lighting field (control pixels) is set based on the pixel values (dose data) of the identified subject pixels, Mode G in which the lighting field is set based on the pixel values of the identified pixels within the radiation field, Mode H in which the lighting field is set by combining the pixel characteristics and the neighboring pixel characteristics as described above, and Mode I in which the lighting field is set using the above-described pixel characteristics.

A plurality of modes may be prepared in the lighting field recognizing portion86to set the lighting field (control pixels) by the mode selected according to the image characteristics.

The plurality of modes prepared in the lighting field recognizing portion86may be the above-described various modes.

For instance, when a histogram shape in which a given range on the high density side having the potential for the direct X-ray region is excluded from a histogram of the whole image is seen, if the histogram width is narrow, the image can be judged as not having a narrowed focus to determine the vicinity of the median value in a cumulative histogram excluding the direct X-ray region as the lighting field without recognizing the radiation field.

Here, the mode selected from the plurality of modes can be determined based on the characteristics in the subject region or the region within the radiation field.

For instance, the mode can be switched from one to another according to the area of the detected region (the subject region, the region within the radiation field, the region except the direct X-ray region in the region within the radiation field (region except a predetermined range on the black side)). For instance, in a small site such as a finger (site for which the diaphragm is narrowed down), the subject region may not be precisely detected and hence the lighting field is determined based on the pixels in the region within the radiation field. On the other hand, in a case where the detected region has a large area, the lighting field is determined based on the pixels in the subject region. For instance, when the area is large, the median value is defined to lie between 30 to 70% in a histogram and when the area is small, the median value is defined to lie between 10 to 90% in a histogram.

If the subject region and the region within the radiation field cannot be detected, the method is changed to a mode switching method which is not based on the regions. For instance, the method can be changed to a method which involves setting the vicinity of the median value in a histogram of the whole image as the lighting field, or a method which involves setting to a fixed lighting field given from outside. In addition, the method may be changed to radiography under dose conditions set in advance for each site instead of setting the lighting field.

It is also possible to prepare a plurality of modes in the lighting field recognizing portion86, to detect the respectively used control pixels76by the plurality of modes, and to determine the control pixels76within the lighting field to be set according to the characteristics of the detected control pixels76.

The plurality of modes prepared in the lighting field recognizing portion86may be the above-described various modes.

For instance, in a case where the lighting field detected based on the subject region and the region within the radiation field is small, the degree of reliability is judged to be low, and the lighting field may be determined based on a histogram of the whole image, or a fixed lighting field be determined as the lighting field, or switching to radiographing under dose conditions preset for each site be made instead of determining the lighting field.

The median value of a cumulative histogram in the subject pixels or the pixels within the radiation field may be calculated to have plural definitions: 10 to 90%, 30 to 70%; and 40 to 60%, and the definition to be used be switched according to the area of the region of the subject pixels or the pixels within the radiation field.

In addition, if the subject region and the region within the radiation field are not detected, the method is changed to a mode switching method which is not based on the regions. For instance, the method can be changed to a method which involves setting the vicinity of the median value in a histogram of the whole image as the lighting field, or a method which involves setting to a fixed lighting field given from outside. In addition, the method may be changed to radiography under dose conditions set in advance for each site instead of setting the lighting field.

Furthermore, it is also possible to detect the lighting field in a plurality of modes and to determine the mode used to calculate the lighting field to be selected, according to the degree of reliability of the lighting field calculated from, for example, the image characteristics such as the area of the subject region.

In a case where the lighting field recognizing portion86failed in identifying the lighting field, it is preferable to update the dose data for analysis for use in lighting field recognition processing and re-execute the lighting field recognition processing, more specifically, to subject new dose data for analysis to the lighting field recognition processing after the accumulated dose data read out from the storage area for accumulation80of the memory84at a different timing is stored in the storage area for analysis82as the new dose data for analysis.

The radiation stop determining portion88determines the stop of radiation based on the accumulated dose data. The radiation stop determining portion88acquires a threshold of the reached dose within the lighting field as determined by the lighting field recognizing portion86and monitors the accumulated dose data in the storage area for accumulation80of the memory84with respect to the threshold. The radiation stop determining portion88reads out the accumulated dose data from the storage area for accumulation80of the memory84at a preset monitoring timing and compares the read-out accumulated dose data with the acquired threshold. As a result, if the read-out accumulated dose data reaches or exceeds the threshold, the radiation stop determining portion88outputs this result to the radiation stop signal generating portion90and if the read-out accumulated dose data does not reach or exceed the threshold, the radiation stop determining portion88waits for the next monitoring timing and determines at the next monitoring timing as to whether radiation is stopped. In other words, the radiation stop determining portion88successively performs determination as to whether radiation is stopped until the read-out accumulated dose data reaches or exceeds the threshold.

The threshold that may be used in the radiation stop signal generating portion90is a threshold of the lighting field as determined by the lighting field recognition processing in the radiation stop determining portion88and is preferably set according to or based on the radiography target such as the radiographic site corresponding to the lighting field to be determined in advance, the radiographic conditions or the above-described plurality of modes. Accordingly, the threshold is preferably switched based on the plurality of modes.

Even if the threshold is set in this way, the threshold is preferably corrected to absorb the differences in device characteristics of the image detection device18before use. Given that the time lag of the wired communication is different from that of the wireless communication as for the communication from the transmitter94, the threshold is preferably corrected to absorb the difference in delay of the communication from the transmitter94.

The radiation stop signal generating portion (first radiation stop signal generating portion)90generates a radiation stop signal Sp (first radiation stop signal Sp1) according to the radiation stop determination made by the radiation stop determining portion88. The radiation stop signal generating portion90receives from the radiation stop determining portion88a determination result indicating that accumulated dose data read out from the storage area for accumulation80of the memory84at a monitoring timing has reached or exceeded a previously acquired threshold, and generates a radiation stop signal Sp (first radiation stop signal Sp1) for stopping X-ray radiation from the X-ray source12.

In order to prevent excessive load or damage on the X-ray source12, the second radiation stop signal generating portion92generates a radiation stop signal Sp (second radiation stop signal Sp2) to stop X-ray radiation from the X-ray source12according to information different from the accumulated dose data of the control pixels76within the lighting field in the image detection device18, for example, X-ray radiation time, information based on the accumulated dose data of the control pixels76outside the lighting field in the image detection device18, information on the radiography target such as the radiographic site, and the like. For example, a backup timer for generating the radiation stop signal Sp to stop X-ray radiation when the X-ray tube load or the loading time of the X-ray source12has reached a specified set value (threshold) is used as the second radiation stop signal generating portion92.

It is also possible to measure the X-ray radiation time from the start of X-ray radiation and to generate the second radiation stop signal Sp2when the X-ray radiation time exceeds a preset threshold.

Dose data other than the lighting field pixel group may also be used to generate the second radiation stop signal Sp2. All the statistics in the group of the control pixels for dose detection76such as maximum value, minimum value and median value may also be used. The statistics in the subject pixel group and pixel groups excluding the subject pixel group, as exemplified by maximum value, minimum value and median value may also be used.

The threshold is preferably preset according to the dose to which the subject12as the radiography target or a radiographic site thereof may be exposed to radiation. More specifically, the threshold such as the set value of a backup timer is preferably switched between the chest and the lumbar spine.

In addition, the threshold is preferably preset based on at least one of information on the plurality of modes and information on the radiographic conditions. More specifically, the threshold such as the set value of a backup timer is preferably switched between a large subject and a small subject as the information on the plurality of modes. As for the information on the radiographic conditions, it is preferable to shorten the time set value of the backup timer at a large tube voltage, prolong the time set value of the backup timer at a small tube voltage, shorten the time set value of the backup timer at a large tube current and prolong the time set value of the backup timer at a small tube current.

The transmitter94transmits the radiation stop signal Sp (first or second radiation stop signal Sp1or Sp2) to the X-ray source12through the radiation source controller36and the high voltage generator24to stop X-ray radiation from the X-ray source12, and controls and executes communication between the first and second radiation stop signal generating portions90,92and the radiation source controller36, between the radiation source controller36and the high voltage generator24, and between the high voltage generator24and the X-ray source12.

The radiation stop signal Sp (Sp1or Sp2) for stopping X-ray radiation from the X-ray source12may be transmitted from the transmitter94by wire or wirelessly. In the illustrated example, the communication between the first and second radiation stop signal generating portions90and92, and the radiation source controller36, between the radiation source controller36and the high voltage generator24and between the high voltage generator24and the X-ray source12is performed by wire. However, the present invention is not limited to this but may be configured to perform a part or the whole of the communication wirelessly.

In the above-described example, the first and second radiation stop signal generating portions90and92generate the radiation stop signals Sp (Sp1, Sp2), respectively, and transmit the signals to the X-ray source12to stop X-ray radiation. However, this is not the sole case of the invention, and the first or second radiation stop signal generating portion90or92may continue to transmit at all times successive radiation signals (radiation enable signals) with a preset period such that stopping the X-ray radiation from the X-ray source12through the radiation source controller36and the high voltage generator24may be executed by stopping the transmission of the successive radiation signals instead of generating the radiation stop signals Sp (Sp1and Sp2).

Moreover, as shown inFIG. 4, it is preferable to display the signal type of the radiation stop signal Sp on a notification unit such as the display30to notify a user (e.g., radiation technologist) whether the radiation stop signal is, for example, the first radiation stop signal Sp1based on the accumulated dose (data) of the control pixels76in the lighting field or the second radiation stop signal Sp2based on the information different from the accumulated dose (data) of the control pixels76within the lighting field, for example, due to the X-ray radiation time exceeding the threshold. In the present invention, instead of displaying on the display30, the signal type of the radiation stop signal Sp may be indicated by a notification unit such as an indicator although not shown, or be notified by a voice generating unit such as an alarm or a speaker as voice, identifiable sound or melody.

The AEC section40that may be used in the invention is basically configured as described above.

Next, the operation of the X-ray exposure control device of the X-ray image detection apparatus in the X-ray imaging system according to the invention and the AEC procedure in the AEC section are described.

FIGS. 5 and 6are a flow chart and a schematic explanatory diagram, respectively, which illustrate an exemplary procedure of the AEC performed in the AEC section of the X-ray exposure control device of the X-ray image detection apparatus in the X-ray imaging system according to the invention.FIG. 7is a chart schematically showing an exemplary flow of X-ray imaging in the X-ray imaging system of the invention.

The preparation in the case of X-ray photography in the X-ray imaging system10is first described prior to the AEC using the X-ray exposure control device of the X-ray image detection apparatus14.

First, the subject16is made to stand at a predetermined position in front of a radiographic table and the height and the horizontal position of the image detection device18set on the upright radiographic table are adjusted to set the image detection device in position with respect to the radiographic site of the subject16. The height and the horizontal position of the X-ray source12and the size of the radiation field are adjusted according to the position of the image detection device18and the size of the radiographic site. Then, the radiographic conditions are set in the control unit20.

At this time, in the standby mode before X-ray photography, the controller64causes the FPD42to repeatedly perform the reset operation.

The preparation in the case of X-ray photography is thus finished.

In Step S10shown inFIG. 5, when a radiation start signal is output from the control unit20in response to two-stage pressing of the radiation switch26, the pixels44and the control pixels76outside the lighting field are transferred from the reset operation to the accumulating operation and the mode is switched to the radiographic mode, and X-ray radiation from the X-ray source12is started in Step S12, as shown inFIG. 7. In the case of the pixels44, the charges having concomitantly occurred are accumulated in the photodiodes48and in the case of the control pixels76, they are flown into the integrating amplifiers66through the signal lines58, are integrated and are converted into analog voltage values, which are held in the CDSs68for a preset period of time.

Next, in Step S14, the following operations are repeated until the radiation stop signal is generated in the subsequent Step S22: The analog voltage values held for a preset period of time in the CDSs68are output to the A/D72as dose detection signals with a preset sampling period, converted into digital dose data in the A/D72, output from the image detection device18with the preset sampling period, and as shown inFIG. 6, read out with the preset sampling period by the readout/accumulation portion78of the AEC section40of the detection controller22in the control unit20and added up in the storage area for accumulation80of the memory84to be stored as cumulative dose data.

In Step S16, as shown inFIG. 6, the cumulative dose data in the storage area for accumulation80of the memory84is retrieved into the storage area for analysis82at a preset timing and transferred as the dose data for analysis (image). In other words, an image derived from the dose data for analysis is acquired. The preset timing may be fixed or variable according to the radiographic site and the radiographic conditions.

In Step S18, as shown inFIG. 6, the lighting field recognizing portion86executes lighting field recognition processing by reference to the dose data for analysis in the storage area for analysis82to automatically determine the lighting field.

In Step S20, as shown inFIG. 6, the radiation stop determining portion88refers, with a preset monitoring period (monitoring timing), to the lighting field as determined by the lighting field recognizing portion86and the cumulative dose data (image) of the storage area for accumulation80at the monitoring timing to thereby acquire the data on the dose reached at the reached lighting field at the monitoring timing.

Next, in Step S22, as shown inFIG. 6, the radiation stop determining portion88performs radiation stop determination which includes determining whether the acquired data on the dose reached at the lighting field has reached a preset threshold through comparison between the dose data and the threshold. If the data on the dose reached at the lighting field reaches or exceeds the threshold, the radiation stop signal generating portion90generates a radiation stop signal in Step S24, as shown inFIG. 6.

On the other hand, if the acquired data on the dose reached at the lighting field does not reach the threshold in Step S22, the process returns to Step S20and acquiring the data on the dose reached at the lighting field at the next monitoring timing and determining as to whether the radiation is stopped in Step S22are repeated until the acquired data on the dose reached at the lighting field reaches the threshold to generate a radiation stop signal in Step S24.

In other words, as shown inFIG. 7, the acquisition of the data on the dose reached at the lighting field in Step S20and the radiation stop determination in Step S22are performed in the period when the dose at the lighting field is monitored.

The procedure for generating the AEC radiation stop signal in the AEC section40is thus finished.

Then, as shown inFIG. 7, the radiation stop signal generated in the radiation stop signal generating portion90is transmitted from the transmitter94to the X-ray source12through the radiation source controller36and the high voltage generator24and the X-ray radiation from the X-ray source12is stopped. In other words, in the control unit20, the radiation source controller36stops supply of electric power from the high voltage generator23to the X-ray source12, whereby the X-ray radiation is terminated.

Under the control of the controller64, the charges accumulated in the photodiodes48of the pixels44flow into the integrating amplifiers66through the signal lines58, output from the integrating amplifiers66to the A/D72as X-ray image detection signals with a preset sampling period, converted into digital X-ray image data, and output from the image detection device18to the memory38of the detection controller22of the control unit20. The X-ray image data output to the memory38by the readout operation is subjected to various image processing steps in various image processing circuits and a sheet of X-ray image is thus produced, as shown inFIG. 7. The X-ray image is displayed on the display30of the control unit20and used in, for example, diagnosis.

In the above-described example, as shown inFIG. 6, readout of dose data of the control pixels76made by the readout/accumulation portion78of the AEC section40is performed at a timing at which the whole of one radiographic image is read out at one time but the present invention is not limited thereto. As shown inFIG. 8, the whole of one radiographic image may be read out in a time-shared manner at a timing of a plurality of (in the illustrated case, four) readout operations. The volume of data read out by one readout operation can be thus reduced to enable high-speed readout.

In a case where the normal pixels44and the pixels for exposure control76are included together as in the image detection device18according to the embodiment under consideration, the number of the control pixels76is advantageously smaller in terms of the calculation amount when configuring the control pixels76but the smaller the number of the control pixels76is, the lower the S/N is. Therefore, the S/N can be improved by having such a configuration as to generate one piece of pixel information for exposure control by addition of a plurality of pixel groups, as shown inFIGS. 9A, 9B and 9C. In the image detection device18shown inFIG. 9A, the S/N can be improved by unifying 2×2 pixels (four pixels) in the pixels for exposure control76as shown inFIGS. 9B and 9Cso as to generate a piece of information.

Moreover, it is preferable to use the configuration of the pixels for exposure control76shown inFIGS. 9B and 9Cin combination in order for the lighting field recognizing portion86of the AEC section40to perform lighting field recognition processing with high accuracy in a short period of time. More specifically, the subject region (subject pixel group) is found as a rough region based on a piece of pixel information generated using the configuration of the pixels for exposure control76in which 2×2 (4) pixels are unified as shown inFIG. 9C, and when determining the use pixels (lighting field) within the found subject region, analysis is performed within the subject region after the pixel configuration is returned to the configuration of 2×2 (4) pixels of the pixels for exposure control76shown inFIG. 9B. The pixels which form the lighting field can be determined with high accuracy in a short period of time.

In a case where some of the normal pixels44are used as the pixels for exposure control76as in the image detection device18in the embodiment under consideration, information is accumulated in the normal pixels44during X-ray radiation, read out after the stop of the radiation and used for image production, whereas information is already read out from the pixels for exposure control76during the X-ray radiation and hence cannot be used without any processing for image formation after the stop of the radiation and the pixels for exposure control76become defect pixels unlike the normal pixels44.

Accordingly, it is also possible to complement the image data of the pixels for exposure control76corresponding to the defect pixels by applying the same method as known pixel defect correction to the positions corresponding to the pixels for exposure control76in image production after the stop of the radiation. Alternatively, it is also possible to use the pixels for exposure control76in image production similarly to the image data of the pixels44by reading out from the memory84information read out for exposure control during the X-ray radiation and using the read-out information.

As described above, according to the invention, it is possible to perform consistent X-ray exposure control regardless of the positioning of a subject by recognizing and determining the lighting field of the radiographic subject during the X-ray photography.

Accordingly, the present invention is capable of stopping X-ray radiation at a proper exposure dose (exposure) according to the subject (radiographic site), in other words, of properly controlling the radiation dose during the X-ray photography according to the subject, and of acquiring an X-ray image of suitable density at all times in the same radiographic environment even in the radiography of a variety of different sites.

In other words, the present invention is capable of consistent radiography at a proper dose regardless of the position of the subject or its radiographic site or of the position of the subject in the whole body.

In the above-described first embodiment of the invention, the X-ray imaging system10uses the X-ray image detection device18in which the normal pixels for image detection44to detect an X-ray image and the pixels for exposure control76are included together. However, the invention is not limited to this but the X-ray imaging system10may be, as shown inFIG. 10, an X-ray imaging system which uses a dedicated device for X-ray image detection only composed of normal pixels44for image detection to detect an X-ray image and an X-ray exposure control device only composed of pixels for exposure control76.

FIG. 10is a schematic explanatory diagram schematically showing another example of the X-ray imaging system to which an X-ray image detection apparatus provided with an X-ray exposure control device according to a second embodiment of the invention is applied.

FIGS. 11A and 11Bare explanatory diagrams illustrating an example of a dedicated device for X-ray image detection and an example of an X-ray exposure control device, respectively, that may be used in the X-ray imaging system shown inFIG. 10.

FIG. 12is an explanatory diagram illustrating an example of a control unit of the X-ray image detection apparatus that may be used in the X-ray imaging system shown inFIG. 10.FIG. 13is a chart schematically showing an exemplary flow of X-ray imaging in the X-ray imaging system shown inFIG. 10.

An X-ray imaging system100according to the second embodiment of the invention as shown in these drawings has the same configuration as the X-ray imaging system10according to the first embodiment of the invention as shown inFIGS. 1 to 7except that a dedicated device for X-ray image detection104and an X-ray exposure control device106are used as separate devices instead of the X-ray image detection device18according to the first embodiment. So, like components are denoted by the same reference numerals and their detailed description is omitted.

As shown in this drawing, the X-ray imaging system100includes an X-ray source12and an X-ray image detection apparatus102. The X-ray image detection apparatus102includes the dedicated device for X-ray image detection (hereinafter referred to as “image specific device”)104which is provided at a position opposed to the X-ray source12and which receives an image of X-rays having passed through a subject16(radiographic site), the X-ray exposure control device (hereinafter simply referred to as “control device”)106which is disposed between the position at which the subject16is radiographed and the image specific device104, and a control unit108which controls the whole operation of the X-ray imaging system100including the operation control of the X-ray source12, the image specific device104and the control device106and image processing of an X-ray image.

In the embodiment under consideration, the X-ray exposure control device106and the portion of the control unit108which controls the operation of the control device106constitute the X-ray exposure control device of the invention.

The image specific device104has quite the same configuration as the X-ray image detection device18shown inFIG. 3except that all the pixels are normal pixels44as shown inFIG. 11A. So,FIG. 11Aomits a detailed configuration and schematically shows only the array of the pixels44.

On the other hand, the control device106has quite the same configuration as the X-ray image detection device18shown inFIG. 3except that nine pixels disposed in a dispersed manner in the illustrated case are all pixels for exposure control76, as shown inFIG. 11B. So,FIG. 11Bomits a detailed configuration and schematically shows only the array of the pixels76. In the illustrated case, the control pixels used in the control device106have a larger pixel size than the normal pixels44used in the image specific device104in order to improve the S/N of the control pixels76.

As shown inFIG. 12, the control unit108includes an X-ray detection controller (hereinafter referred to simply as “detection controller”)110comprehensively controlling the whole operation of the apparatus, and a high voltage generator24, a radiation switch26, an input device28, a display30and a memory32connected to the detection controller110.

The detection controller110includes a device controller34, a radiation source controller36, a memory38and an X-ray exposure controller (hereinafter also referred to as “AEC section”)40.

Although the configuration of the detection controller110is the same as that of the detection controller22shown inFIG. 2, the control device106is directly connected to the AEC section40and dose data of the control device106is input to the AEC section40, whereas the image specific device104is directly connected to the memory38and image data of the image specific device104is input to the memory38and stored.

As shown inFIG. 13, the AEC uses dose data of the control pixels76of the control device106, but the flow of the X-ray imaging in the X-ray imaging system100according to the second embodiment can be performed in the same manner as the automatic exposure control (AEC) in the flow of the X-ray imaging in the X-ray imaging system10according to the first embodiment as shown inFIG. 7as for the lighting field recognition (see Step S18inFIG. 5) and the dose monitoring in the lighting field (see Steps S20to S22inFIG. 5) which are performed between the start of X-ray radiation (see Step S10inFIG. 5) and the stop of X-ray radiation (see Step S24inFIG. 5). On the other hand, X-ray image data of the image specific device104is used in X-ray image formation in the X-ray imaging system100according to the second embodiment, but the X-ray image formation itself can be performed in the same manner as the image formation in the flow of the X-ray imaging in the X-ray imaging system10according to the first embodiment as shown inFIG. 7.

From the above, the X-ray imaging system100according to the second embodiment can perform the AEC and X-ray image formation in quite the same manner as the X-ray imaging system10according to the first embodiment although the X-ray image detection device18according to the first embodiment is separated into the dedicated device for X-ray image detection104and the X-ray exposure control device106which are used herein. Accordingly, the X-ray imaging system100according to the second embodiment can achieve quite the same effects as the X-ray imaging system10according to the first embodiment.

Since the generation of an X-ray radiation stop signal Sp by the AEC can be performed independently of the X-ray image formation as in the X-ray imaging system100according to the second embodiment, the present invention is also applicable to an X-ray imaging system in which the X-ray image formation and the AEC are performed in discrete entities. For instance, the present invention is also applicable to the radiation detection device as disclosed in JP 9-73144 A in which the image detection and the AEC are performed in discrete entities, or a CR X-ray imaging system using a storage phosphor sheet (IP) instead of the X-ray exposure control device106provided with the FPD of a DR type for X-ray image formation in the X-ray imaging system100according to the second embodiment, and an X-ray imaging system of an X-ray film type.

The present invention is of course applicable to any X-ray imaging system, as long as it uses an integrated X-ray image detection device in which normal pixels for image detection and pixels for exposure control are included together as in the X-ray imaging system according to the first embodiment. For instance, the present invention is also applicable to the X-ray diagnostic apparatus described in Patent Literature 1 which uses an integrated device, the radiation imaging apparatus in Patent Literature 2, the radiation detection apparatus disclosed in JP 2004-170216 A, the radiation imaging apparatus using an electronic cassette having a built-in phototimer as disclosed in JP 2003-302716 A, and the like.

In addition, the X-ray image detection device, the dedicated device for X-ray image detection and the X-ray exposure control device that may be used in the X-ray imaging systems according to the first and second embodiments as described above use the TFTs but may use a CMOS disclosed in, for example, JP 2005-143802 A.

The X-ray image detection device18for use in the X-ray imaging system10according to the first embodiment of the invention as described above is used by being fixed to a radiographic table or is used by being connected to the X-ray detection controller22of the control unit20including the memory38and the AEC section40. However, the present invention is not limited thereto and use may be made of a so-called electronic cassette which is a transportable type X-ray image detection device including the memory38and the AEC section40.

An X-ray imaging system and a transportable X-ray image detection device according to a third embodiment of the invention are shown, inFIGS. 14 to 16.

FIG. 14is an explanatory diagram illustrating an example of the control unit of the X-ray image detection apparatus that may be used in the X-ray imaging system according to the embodiment of the invention;FIGS. 15 and 16are each an explanatory diagram illustrating an example of the X-ray image detection device that may be used in the X-ray imaging system shown inFIG. 14.

An X-ray imaging system120according to the third embodiment of the invention as shown inFIG. 14has the same configuration as the X-ray imaging system10according to the first embodiment of the invention as shown inFIG. 2except that the X-ray image detection device18according to the first embodiment is replaced by transportable X-ray image detection devices18aand18b. So, like components are denoted by the same reference numerals and their detailed description is omitted.

As shown in this drawing, the X-ray imaging system120according to the third embodiment of the invention includes an X-ray source12and an X-ray image detection apparatus122. The X-ray image detection apparatus122includes the X-ray image detection devices18aand18bwhich are provided at positions opposed to the X-ray source12and which receive an image of X-rays having passed through a subject16(radiographic site) and a control unit124which controls the whole operation of the X-ray imaging system120including the operation control of the X-ray source12and the image detection devices18aand18b, and image processing of an X-ray image.

According to the embodiment under consideration, in the X-ray image detection apparatus122, pixels for exposure control76except a portion of normal pixels44in the image detection devices18aand18b, and each component of the control unit124except a portion where an X-ray image from the normal pixels44is processed mainly constitute the X-ray exposure control device according to the third embodiment of the invention, as in the above-described first embodiment.

As shown inFIG. 14, the control unit124includes an X-ray detection controller126comprehensively controlling the whole operation of the apparatus, as well as a high voltage generator24, a radiation switch26, an input device28, a display30, a memory32, a wireless communication section128and a wired communication section130connected to the detection controller126.

The detection controller126include a device controller34, a radiation source controller36, and the wireless communication section128and the wired communication section130for connection with the image detection devices18aand18b.

The control unit124is provided with the wireless communication section128and the wired communication section130. The wireless communication section128is wirelessly connected to the image detection devices18aand18bin a case where the X-ray radiation stop timing is defined based on the output from control pixels76a(seeFIG. 15) and76(seeFIG. 16) of the image detection devices18aand18b, respectively. In this case, upon receipt of a warm-up start signal from the radiation switch26, the radiation source controller36transmits an inquiry signal to the image detection devices18aand18bthrough the wireless communication section128. Upon receipt of the inquiry signal, the image detection devices18aand18bcheck whether they are ready for radiographing and transmits a radiation enable signal if they are ready for radiographing. Upon receipt of the radiation enable signal at the wireless communication section128and further receipt of a radiation start signal from the radiation switch26, the radiation source controller36starts electric power supply from the high voltage generator24to the X-ray source12. Upon receipt of a radiation stop signal issued from the image detection devices18aand18bat the wireless communication section128, the radiation source controller36stops electric power supply from the high voltage generator24to the X-ray source12to terminate X-ray radiation.

The wireless communication section128wirelessly communicates with the image detection devices18aand18bnot only for the AEC signals but also other signals for radiographic conditions and X-ray image data. The wired communication section130is connected by wire to the image detection devices18aand18bin a case where wireless communication of radiographic conditions, image data and the like is impossible. The wired communication section130has the power supply function and supplies electric power for drive to the image detection devices18aand18bin a case where the wired communication section130is connected by wire to the image detection devices18aand18b.

The image detection device18aincludes an FPD42a(seeFIG. 15) and a transportable casing containing the FPD42a. The image detection device18bincludes an FPD42b(seeFIG. 16) and a transportable casing containing the FPD42b. The casing in each of the image detection devices18aand18bhas a substantially rectangular, flat shape.

A plurality of image detection devices18a,18b, for example, two image detection devices18a,18bare provided to be used in an upright radiographic table and a decubitus radiographic table which are not shown in a radiographic room having the X-ray imaging system120disposed therein. Each of the image detection devices18aand18bis detachably set in a holder (not shown) of an upright photographic table or a decubitus radiographic table so that an imaging surface46(seeFIG. 3) of the FPD42aor the FPD42bis held in such a position as to be opposed to the X-ray source12. It is also possible to use the image detection devices18aand18balone not by setting them on an upright radiographic table or a decubitus radiographic table but by putting them on a bed (not shown) on which a subject is lying supine or by making the subject carry them.

InFIGS. 15 and 16, a wireless communication section132and a wired communication section134for communicating with the control unit124by a wireless system or a wired system, and a battery138are incorporated into each of the image detection devices18aand18b. The wireless communication section132and the wired communication section134mediate the transmission and reception of various information and signals, including image data of a controller64abetween the control unit124and the image detection devices18a,18b. In particular, the wireless communication section132communicates with the wireless communication section128of the control unit124for the AEC signals. In wireless communication, the battery138supplies electric power for operating the respective portions of the image detection device18aor18b. The battery138used is of a comparatively small size so as to be placed in the thin image detection device18aor18b. The battery138can also be taken out of the image detection device18aor18b, set on a dedicated cradle and charged. The battery138may be configured to be capable of wireless power supply.

The wired communication section134is connected by wire to the wired communication section130of the control unit124in a case where wireless communication between the image detection devices18a,18band the control unit124is made impossible for lack of power of the battery138. In a case where a cable from the control unit124is connected to the wired communication section134, the function of the wireless communication section132is stopped and the wired communication section134functions instead, thus enabling wired communication with the control unit124. At this time, power supply from the control unit124to the image detection devices18a,18bis made possible and power supply from the battery138is stopped. The battery138may be charged by the electric power from the control unit124. A conventionally known technique including measurement of the contact current between a connector and a cable socket may be used as the method of detecting cable connection.

Each of the FPDs42aand42bincludes a TFT active matrix substrate, and the imaging surface46in which the plurality of pixels44for accumulating charges according to the reached X-ray dose are arrayed is formed on top of the substrate.

Each of the FPDs42aand42bis of an indirect conversion type which includes a scintillator (phosphor) capable of converting X-rays into visible light and which photoelectrically converts in the pixels44visible light obtained by conversion in the scintillator.

The FPD42ais different from the FPD42shown inFIG. 3in the configuration of the control pixels76aand their drive system, and in the presence of a memory38a, an AEC section40, a communication section136having the wireless communication section132and the wired communication section134as well as the battery138but the FPD42ahas the same configuration as the FPD42except these points. So, a detailed description is omitted.

As in the FPD42a, the FPD42bis different from the FPD42shown inFIG. 3in that the former includes a memory38a, an AEC section40, a communication section136having the wireless communication section132and the wired communication section134as well as the battery138but the array of the normal pixels44and the control pixels76on the imaging surface46, and the configuration of a signal processing circuit62are quite the same. So, their description is omitted and the FPD42ais described below as a typical example.

In a TFT50, a gate electrode, a source electrode, and a drain electrode are connected to a scanning line56a, a signal line58and a photodiode48, respectively. The scanning lines56aand the signal lines58are formed in a grid shape and the number of the scanning lines56aprovided corresponds to the number of rows of the pixels44(n rows) on the imaging surface46aand the number of the signal lines58provided corresponds to the number of columns of the pixels44(m columns) on the imaging surface46. The scanning lines56aare connected to a gate driver60aand the signal lines58are connected to the signal processing circuit62.

The gate driver60adrives each TFT50so that the TFT50performs the accumulating operation for accumulating signal charges in the pixel44according to the X-ray dose reached, the readout (main reading) operation for reading out signal charges from the pixel44, and the reset (void reading) operation. The controller64acontrols the start timing of each of the foregoing operations executed by the gate driver60a.

In the accumulating operation, the TFTs50are turned off and signal charges are accumulated in the pixels44during this period. In the readout operation, gate pulses G1to Gn which drive the TFTs50in the same rows all together are successively generated from the gate driver60ato sequentially activate the scanning lines56aon a row by row basis and the TFTs50connected to the scanning lines56aare turned on on a row by row basis. When the TFTs50are turned on, the charges accumulated in the capacitors of the pixels44are read out to the signal lines58and are input to the signal processing circuit62.

The signal processing circuit62includes integrating amplifiers66, CDS circuits (CDS)68, a multiplexer (MUX)70, an A/D converter (A/D)72, and the like. The integrating amplifiers66integrates the charges input from the signal lines58, converts them into analog voltage signals V1to Vm and outputs the analog voltage signals. The output terminal of an operational amplifier66ain each column is connected to the MUX70through an amplifier74and the CDS68. The output side of the MUX70is connected to the A/D72. The A/D72converts the input voltage signals V1to Vm into digital voltage signals and outputs the digital voltage signals to the memory38aor the AEC section40incorporated in the image detection device18a. An amplifier may be connected between the MUX70and the A/D72. It is also possible to provide an A/D for each signal line58, and in this case the A/Ds are followed by the MUX.

When the MUX70reads out the voltage signals V1to Vm in one row from the integrating amplifiers66, the controller64aoutputs a reset pulse RST to the integrating amplifiers66to turn on reset switches66c. The signal charges in one row as accumulated in capacitors66bare thereby discharged and the integrating amplifiers66are reset. After the integrating amplifiers66have been reset, the reset switches66care turned off again. After the lapse of a preset period of time, one of sample-and-hold circuits of each of the CDSs68is held to sample the kTC noise component of the integrating amplifiers66. Thereafter, a gate pulse for the next row is output from the gate driver60ato start readout of signal charges from the pixels44in the next row. In addition, after the lapse of a preset period of time from the output of the gate pulse, the signal charges from the pixels44in the next row are held by the other sample-and-hold circuit of each of the CDSs68. These operations are sequentially repeated to read out signal charges from the pixels44in all the rows. High-speed drive is possible by adopting pipeline processing which performs these processing steps at a time.

Upon completion of readout in all the rows, image data representing an X-ray image corresponding to a screen is recorded in the memory38a. This image data is immediately read out from the memory38aand output to the control unit124through the wireless communication section132or the wired communication section134. The X-ray image of the subject is thus detected.

The memory38ahas such a capacity that X-ray image data in one screen can be radiographed a plurality of times, for example, 100 times and stored. In a case where X-ray image data cannot be transmitted from the wireless communication section132or the wired communication section134because of a communication failure, the memory38atemporarily accumulates the X-ray image data output from the FPD42aduring that time. The X-ray image data temporarily accumulated in the memory38ais transmitted at a time or in several batches at the time of recovery from the communication failure. A storage unit for temporarily accumulating X-ray image data at the time of a communication failure may be provided separately from the memory38a. A removable medium which is detachable from the image detection device18amay be used as the storage unit so that the removable medium can be detached from the image detection device18aat the time of a communication failure and directly set to the control unit124to take X-image data therefrom.

The reset operation is carried out by, for example, a sequential reset method in which the pixels44are reset on a row by row basis. In the sequential reset method, the gate pulses G1to Gn are sequentially issued from the gate driver60ato the scanning lines56ato turn on the TFTs50of the pixels44on a row by row basis, as in the readout operation of the signal charges. While the TFTs50are turned on, the dark charges flow from the pixels44through the signal lines58to the capacitors66bof the integrating amplifiers66. In the reset operation, the MUX70does not read out the charges accumulated in the capacitors66b, unlike the readout operation. A reset pulse RST is output from the controller64ain synchronism with occurrence of each of the gate pulses G1to Gn to turn on the reset switches66c, whereby the charges accumulated in the capacitors66bare discharged to reset the integrating amplifiers66.

In addition to the normal pixels44to which the TFTs50driven by the gate driver60aand the scanning lines56aas described above are connected, the FPD42aincludes within the same imaging surface46the control pixels76ato which TFTs50adriven by a driver60bdifferent from that for the normal pixels44and scanning lines56bare connected. The TFTs50aare turned on by gate pulses g1to gn from the gate driver60b. The basic configuration of each control pixel76asuch as the photodiode48is the same except only the drive source and the accumulated charges can be read out from the signal lines58independently of the pixels44. As for the reset operation and readout operation, after the operation of the normal pixels44has been completely finished, gate pulses g1to gn are issued from the gate driver60bin the same manner to perform the reset operation or the readout operation of the control pixels76a. Alternatively, the reset operation or the readout operation of the pixels44and the control pixels76ain the same row is simultaneously performed in synchronism with the operation of the gate driver60a. The control pixels76aare pixels used to detect the X-ray dose reached at the imaging surface46and functions as the AEC sensors. The control pixels76aaccount for about several ppm to several percent of the pixels44in the imaging surface46.

Similarly to the control pixels76shown inFIG. 3, the control pixels76aare not disposed locally within the imaging surface46but are evenly scattered within the imaging surface46.

When a gate pulse is generated from the gate driver60bto turn on the TFT50a, a signal charge generated in the control pixel76ais read out to the signal line58. Since the drive source of the control pixel76ais different from that of the pixels44, the signal charge of the control pixel76acan be read out even when the TFTs50of the pixels44in the same column are turned off and the pixels44are in the course of accumulating operation for accumulating the signal charges. At this time, the charge generated in the control pixel76aflows into the capacitor66bof the integrating amplifier66on the signal line58to which the control pixel76ais connected. During the accumulating operation of the pixels44, the TFT50ais turned on and the charge from the control pixel76awhich is accumulated in the integrating amplifier66is output to the A/D72with a preset sampling period.

As in the controller64shown inFIG. 3, the controller64ais provided with circuits (not shown) which perform various image processing steps such as offset correction, sensitivity correction and defect correction on X-ray image data in the memory38a.

The drive of the AEC section40is controlled by the controller64a. The AEC section40acquires from the A/D72digital voltage signals (hereinafter referred to as “dose detection signals”) from the signal lines58to which the control pixels76aare connected, and performs the AEC based on the acquired dose detection signals.

The AEC section40has the configuration shown inFIG. 4, so its description is omitted. In order to stop X-ray radiation more rapidly, the AEC section40may be provided upstream of the A/D72so that a radiation stop signal is generated based on an analog signal and output. Alternatively, an analog signal may be transmitted to the control unit124as a dose detection signal so that a radiation stop signal is generated in the radiation source controller36of the control unit124.

The wireless communication section132performs transmission and reception of AEC signals, to be more specific, reception of an inquiry signal, transmission of a radiation enable signal in response to the inquiry signal, reception of a radiation start signal, and transmission of a radiation stop signal.

Ad hoc communication is used as the wireless communication system between the wireless communication section128of the control unit124and the wireless communication section132of the image detection device18a. The ad hoc communication is used in direct wireless communication between wireless communication devices. Therefore, as compared with infrastructure communication in which communication of medical devices other than the X-ray imaging system120and communication of various data such as electronic medical records, medical reports and accounting data are performed through a wireless access point, a hospital LAN or a switching device such as a hub, data communication delays (lags) are less likely to occur and the average delay time in data communication is small. Accordingly, it can be said that the communication speed of the ad hoc communication is higher than that of the infrastructure communication.

The control unit124is often disposed in a radiographic room. Accordingly, by adopting the ad hoc communication in the communication of AEC signals including the radiation stop signal which is communicated between the control unit124and the image detection device18a, it is possible to perform consistent communication and also realize high speed communication without causing data communication delay because the distance between the control unit124and the image detection device18ais small and radio waves are also easily received. Since a relay is not used therebetween, it is possible to immediately restore from a communication failure only by operation check of the wireless communication section132or part replacement.

It is preferable to adopt, for example, an optical beacon typified by IrDA or other infrared communication or a radio beacon as the wireless communication system between the wireless communication section128and the wireless communication section132. The optical beacon and the radio beacon are suitable to the communication of the AEC signal which is used to immediately stop X-ray radiation as soon as the target dose has been reached, because the number of bits of communication signals is comparatively small and the communication systems are also simple and are less likely to cause delay.

In case of a communication failure, in the wired communication, the cable disconnection or the contact failure of a connector must be checked to explore the cause of the communication failure, or if a relay such as a hub is connected between the control unit124and the image detection device18a, its operation must be also checked. However, as described above, according to the present invention, the AEC signals including the radiation stop signal are always wirelessly communicated even in a case where a wired connection is established between the control unit124and the image detection device18a. Therefore, the cause of the communication failure can be simply identified in the wireless communication only by checking the operation of the wireless communication section128and the wireless communication section132, which also enables rapid recovery from the communication failure. Accordingly, the present invention is not likely to encounter a situation in which radiographing cannot be performed for a prolonged period of time due to a communication failure to make a patient wait unnecessarily long and this configuration is resistant to accidents.

In a case where image data cannot be transmitted, X-ray image data is temporarily accumulated in the memory38aand hence radiographing can be continued if wireless communication of the AEC signals is alive. If the memory38ahas a capacity that may resist a menu such as tomosynthesis imaging in which imaging is performed several times in succession, continuous imaging can be continued to the end without being stopped even in a situation where image data cannot be transmitted.

The AEC signal is after all an ON/OFF signal and hence has an extremely smaller capacity than the image data and the like. Therefore, the power (radio field intensity) required for wireless communication is small and the AEC signal can also be used without any problem in a subject having a pacemaker. The power consumption involved in wireless communication is also small. The capacity of image data is large and hence wireless communication requires a large amount of power but the power consumption can be suppressed by switching so as to perform the transmission and reception of image data by wire when a cable is connected.

By applying ad hoc communication to the AEC signal between the control unit124and the image detection device18a, an investigation to explore the cause of a communication failure and recovery from the communication failure can be made more speedily because an unnecessary device such as a hub does not intervene between the control unit124and the image detection device18a. The same applies to the case where a beacon which has a simple configuration and facilitates failure analysis is adopted.

The AEC signal and the other signal such as image data may have a common or different resource for the wireless communication function. In a case where the resource is common, the number of parts can be reduced and in a case where the resource is different, even if the transmission/reception timing of the AEC signal is the same as that of the other signal, this problem can be solved.

In the embodiment under consideration, the inquiry signal, the radiation enable signal in response to the inquiry signal, the radiation start signal and the radiation stop signal were described as the AEC signals but only the radiation stop signal is more preferably used as the AEC signal. In this case, when the image detection device18ais connected by wire to the control unit124, wireless communication is performed only for the radiation stop signal and the transmission and reception of the inquiry signal, the radiation enable signal in response to the inquiry signal and the radiation start signal as well as image data are performed by wired communication. The power consumption of the battery can be thus minimized. In case of a wireless communication failure, the control is switched so that X-ray radiation is detected by the control pixels76a, the reset operation is performed and the control pixels76aturn into an accumulation state. It is thus possible to start radiographing without synchronization in the transmission and reception of signals in the wired communication although there is more or less X-ray loss. In other words, radiographing can be started by detecting X-rays but communication is essential to stop radiographing and hence radiographing can be continued also in case of a wired communication failure by wirelessly communicating only the AEC stop signal. What is more, as in the above embodiment, in case of a wireless failure, the failure is immediately analyzed and recovery from the failure is also speedy.

In the above embodiment, the radiation stop signal is output at a point in time when the integrated value of the dose detection signal has reached a radiation stop threshold, but the estimated time at which the cumulative X-ray dose would reach a target value may be calculated in the AEC section40based on the integrated value of the dose detection signal so that the radiation stop signal is output when the calculated estimated time is reached.

Alternatively, it is also possible to continuously transmit successive radiation signals from the wireless communication section132of the image detection device18a(18b) toward the wireless communication section128of the radiography control unit124from the start of X-ray radiation until a determination that the integrated value of the reached X-ray dose has reached a target value is made in the AEC section40, and to stop the X-ray radiation when the wireless communication section128cannot receive the successive radiation signals. In the embodiment under consideration, in a case where a situation occurs in which transmission and reception of the radiation stop signal cannot be performed between an electronic cassette and the control unit, X-ray radiation is continuously performed even after the time at which the X-ray radiation should be stopped, which may cause a patient to be excessively exposed to radiation. However, the patient may at least not be exposed to excessive radiation although the dose may be insufficient because the X-ray radiation is anyway stopped when the reception of the successive radiation signals is stopped.

In the above-described embodiments, since the FPD of a TFT type, that is, the FPD composed of the normal pixels44each including the TFT50or50aand the control pixels are used as the FPD42,42aor42bof the image detection device18,18aor18b, destructive readout in which all the charges accumulated in the normal pixels44are read out in each readout is performed. However, the present invention is not limited thereto and an element which is capable of non-destructive readout, for example, a CMOS sensor capable of non-destructive readout as disclosed in JP 2005-143802 A may be used as the dose detection element.

A description is given of a case where a non-destructive readable element, for example, a non-destructive readable CMOS sensor as disclosed in JP 2005-143802 A is used as the image detection device.

FIG. 17is an explanatory diagram illustrating an example of an X-ray image detection device using a non-destructive readable CMOS circuit that may be used in the X-ray imaging system shown inFIG. 14.

An X-ray image detection device18cshown inFIG. 17has the same configuration as the X-ray image detection device18ashown inFIG. 15except that each normal pixel44and each control pixel76aconstituting the imaging surface46aof the FPD42aare provided with the TFT50and the TFT50a, respectively, whereas each of normal pixels45and normal pixels also serving as control pixels (hereinafter also referred to as “dual-purpose pixels”)45aconstituting an imaging surface46bof an FPD42cis provided with a CMOS circuit51and that the AEC section40is replaced by an AEC section40a. So, like components are denoted by the same reference numerals and their detailed description is omitted.

The image detection device18cshown inFIG. 17includes the FPD42cand a casing containing the FPD42c.

The FPD42cincludes the imaging surface46bon which the normal pixels45and the dual-purpose pixels45aeach provided with the CMOS circuit51are arrayed; a gate driver60awhich drives all the CMOS circuits51of the normal pixels45and the dual-purpose pixels45a; a gate driver60bwhich drives the CMOS circuits51of the dual-purpose pixels45a; a signal processing circuit62to which signal lines58connected to the CMOS circuits51of the normal pixels45and the dual-purpose pixels45aare connected; a memory38a, the AEC section40a, a controller64b, a communication section136including a wireless communication section132and a wired communication section134; and a battery138.

The normal pixels45and the dual-purpose pixels45aare pixels of the same configuration and are each composed of a photodiode48and the CMOS circuit51. The only difference between the dual-purpose pixels45aand the normal pixels45is the drive system. The normal pixels45are driven by the gate driver60aand scanning lines56a, whereas the dual-purpose pixels45aare driven not only by the gate driver60aand the scanning lines56aas in the normal pixels45but also by the other gate driver60band scanning lines56b, and can read out accumulated charges converted into voltage values from the signal lines58independently of the normal pixels45. The dual-purpose pixels45aare pixels used to detect the X-ray dose reached at the imaging surface46band functions as the AEC sensors.

Since the reset operation and readout operation of the dual-purpose pixels45aare similar to those of the control pixels76ashown inFIG. 15, their detailed description is omitted.

The CMOS circuit51is provided in each of all the normal pixels45and all the dual-purpose pixels45a, is composed of a plurality of MOS transistors, and includes three terminals, one being connected to the scanning line56a, another being connected to the signal line58and the other being connected to the photodiode48. For instance, the CMOS circuit51is composed of three MOS transistors including a scan transistor, an output transistor and a reset transistor which are mutually connected; the gate electrode of the scan transistor is connected to the scanning line56a, the source electrode of the scan transistor to the signal line58, the drain electrode of the scan transistor to the source electrode of the output transistor, the gate electrode of the output transistor to the photodiode48, the drain electrode of the output transistor to the power supply voltage, the gate electrode of the reset transistor to a reset line (not shown), whereby the signal charge generated in the photodiode48and accumulated in the photodiode48or a capacitor (not shown) is converted in the output transistor into a voltage signal, the voltage signal output from the output transistor is selectively output to the signal line58through the source electrode of the scan transistor which is driven by the scanning line56a.

The CMOS circuit51is not limited to the above-described circuit but any element may be used without particular limitation if the accumulated signal charge converted into a voltage value can be read out with the signal charge accumulated in the capacitor or the like maintained, and further if a non-destructive readable element is used.

On the other hand, in the CMOS circuit51of the normal pixel also serving as the control pixel45a, the terminal connected to the scanning line56ais also connected to the scanning line56b, and when the CMOS circuit51is driven by the scanning line56b, for example, the accumulated signal charge is converted into a voltage signal in the output transistor, and the voltage signal output from the output transistor is selectively output to the signal line58through the source electrode of the scan transistor driven by the scanning line56b.

As described above, in each of the normal pixels45and the dual-purpose pixels45a(hereinafter also referred to simply as “pixels45and45a”), the signal charge accumulated in the photodiode48or the capacitor is not directly read out but is read out from the signal line58after conversion into a voltage signal in the output transistor and hence the accumulated signal charge is maintained without any processing and thereafter accumulated, and non-destructive readout is thus possible. In other words, even during the accumulating operation in which the normal pixels45and the dual-purpose pixels45aaccumulate signal charges, or even at any timing, accumulated signal charges of the normal pixels45and the dual-purpose pixels45aconverted into voltage signals can be read out.

In the imaging surface46bof the FPD42c, the scanning lines56a,56band the signal lines58are formed in a grid shape and the number of the scanning lines56a,56bprovided corresponds to the number of rows of the pixels45,45a(n rows) on the imaging surface46band the number of the signal lines58provided corresponds to the number of columns of the pixels45,45a(m columns) on the imaging surface46b. The scanning lines56aare connected to the gate driver60a, the scanning lines56bto the gate driver60band the signal lines58to the signal processing circuit62.

The normal pixels also serving as the control pixels45aare provided so as to be evenly scattered to account for about several ppm to several percent of the normal pixels45. In the illustrated example, up to one dual-purpose pixel45ais provided for each row of the normal pixels45and hence the CMOS circuit51of one dual-purpose pixel45ais connected to the scanning line56bin each row.

The gate drivers60aand60bdrive the CMOS circuits51so that the CMOS circuits50perform the accumulating operation for accumulating signal charges according to the reached X-ray dose in (the capacitors) of the pixels45and45a, the readout (main reading) operation for reading out signal charges converted into voltage values from the pixels45and45a, and the reset (void reading) operation. The controller64bcontrol the start timing of each of the foregoing operations executed by the gate driver60a.

In the accumulating operation, the CMOS circuits51are turned off and signal charges are accumulated in the pixels45and45aduring this period.

In the readout operation from the pixels45and45athrough the scanning lines56a, gate pulses G1to Gn which drive the CMOS circuits51in the same rows all together are successively generated from the gate driver60ato sequentially activate the scanning lines56aon a row by row basis and the CMOS circuits51connected to the scanning lines56aare turned on on a row by row basis.

On the other hand, in the readout operation from the dual-purpose pixels45athrough the scanning lines56b, gate pulses g1to gn which drive the CMOS circuits51in specified rows are successively generated from the gate driver60bto sequentially activate the scanning lines56bon a row by row basis and the CMOS circuits51connected to the scanning lines56bare turned on a row by row basis.

When the CMOS circuits51are thus turned on, the signal charges converted into voltage signals which are accumulated in the capacitors of the normal pixels45and the dual-purpose pixels45aare read out to the signal lines58to be input to the signal processing circuit62.

When gate pulses are generated from the gate driver60bto turn on the CMOS circuits51, the signal charges generated in the dual-purpose pixels45aand converted into voltage values are read out to the signal lines58. At this time, since the dual-purpose pixels45aare driven by a drive source different from that for the normal pixels45, the CMOS circuits51of the normal pixels45in the same column are turned off and the accumulated signal charge in the dual-purpose pixel45a, which is converted into a voltage signal, can only be read out.

In this way, the FPD42cis capable of reading out the accumulated signal charges converted into voltage signals from the dual-purpose pixels45aas accumulated dose data at a required timing in order to detect the X-ray dose reached at the imaging surface46bat a preset timing.

FIG. 18is a block diagram of another example of the AEC section that may be used in the control unit of the X-ray image detection apparatus shown inFIG. 2.

An AEC section40ashown inFIG. 18has the same configuration as the AEC section40shown inFIG. 4except that the readout/accumulation portion78is replaced by a readout portion79and that the memory84including both the storage area for accumulation80and the storage area for analysis82is replaced by a memory84aonly including a storage area for analysis82. So, like components are denoted by the same reference numerals and their detailed description is omitted.

As shown in the drawing, the AEC section40aincludes the readout portion78a, the memory84ahaving the storage area for analysis82, a lighting field recognizing portion86, a radiation stop determining portion88, a radiation stop signal generating portion90, a second radiation stop signal generating portion92and a transmitter94.

The readout portion79directly reads out accumulated dose data of the dual-purpose pixels45afrom the image detection device18cat a preset timing so that the lighting field recognizing portion86determines the lighting field, and directly reads out the accumulated dose data of the dual-purpose pixels45afrom the image detection device18cat each preset monitoring timing so that the radiation stop determining portion88determines the stop of radiation.

The memory84ahas the storage area for analysis82for storing the accumulated dose data of the dual-purpose pixels45adirectly read out from the image detection device18cat the preset timing by the readout portion79as the dose data for analysis.

Since the X-ray image detection device18cin the embodiment shown inFIG. 17uses the CMOS circuit51in each of the normal pixels45and the dual-purpose pixels45aof the FPD42c, in the case of a non-destructive readable device such the CMOS circuit, the dose data accumulated in the normal pixels45and the dual-purpose pixels45ais not reset even after the dose data accumulated in the CMOS circuits from the normal pixels45and the dual-purpose pixels45ahas been read out, and hence it is not necessary to accumulate in a different memory the dose data having been accumulated and read out. Therefore, the memory84aof the AEC section40ain the embodiment shown inFIG. 18does not need the storage area for accumulation80(seeFIG. 4) which is necessary to sequentially accumulate the dose data having been accumulated and read out in the memory84of the AEC section40in the case of using the TFT X-ray image detection devices18,18a,18b(FIG. 3,FIG. 15,FIG. 16).

The lighting field recognizing portion86automatically recognizes the lighting field of the subject16based on the dose data for analysis as stored in the storage area for analysis82of the memory84a, and the radiation stop determining portion88directly reads out the accumulated dose data of the dual-purpose pixels45awithin the lighting field from the image detection device18cat each preset monitoring timing as the accumulated dose data within the lighting field as determined by the lighting field recognizing portion86, and determines as to whether radiation is stopped, based on the accumulated dose data within the lighting field having been read out.

Next, the procedure of the AEC in the AEC section42ashown inFIG. 18is described.

FIGS. 19 and 20are a flow chart and a schematic explanatory diagram, respectively, which illustrate a procedure of the AEC performed in the AEC section42a.

Steps S30, S36, S40and S42in the flow chart and the explanatory diagram of the AEC procedure as shown inFIGS. 19 and 20are the same steps as Steps S10, S18, S22and S24in the flow chart and the explanatory diagram of the AEC procedure as shown inFIGS. 5 and 6. So, their detailed description is omitted.

After the end of the preparation for the X-ray photography, in Step S30shown inFIG. 19, in response to the radiation start signal output from the control unit20, the normal pixels45and the normal pixels also serving as the control pixels45aof the image detection device18care transferred from the reset operation to the accumulating operation and the mode is switched to the radiographic mode, and X-ray radiation from the X-ray source12is started in Step S32. The photodiodes48in the normal pixels45and the dual-purpose pixels45astart to accumulate the charges having concomitantly occurred as shown inFIG. 20.

Next, in Step S34ofFIG. 19, the conversion voltage signals of the charges accumulated in the dual-purpose pixels45aof the FPD42cof the image detection device18cuntil the preset timing are directly read out by the readout portion79of the AEC section40aat the preset timing and stored as the dose data for analysis (image) in the storage area for analysis82of the memory84a. In other words, an image derived from the dose data for analysis is acquired. The preset timing may be fixed or variable according to the radiographic site and the radiographic conditions.

In Step S36ofFIG. 19, the lighting field recognizing portion86executes lighting field recognition processing by reference to the dose data for analysis in the storage area for analysis82to automatically determine the lighting field, as shown inFIG. 20.

In Step S38ofFIG. 19, as shown inFIG. 20, the readout portion79of the AEC section40adirectly reads out at a preset timing the conversion voltage signals of the charges accumulated in the dual-purpose pixels45awithin the lighting field as determined by the lighting field recognizing portion86with the predetermining monitoring period (monitoring timing), and the radiation stop determining portion88acquires the data on the dose reached at the lighting field at the monitoring timing.

Next, in Step S40ofFIG. 19, as shown inFIG. 20, the radiation stop determining portion88performs radiation stop determination which includes determining whether the acquired data on the dose reached at the lighting field has reached a preset threshold through comparison between the dose data and the threshold. If the data on the dose reached at the lighting field reaches or exceeds the threshold, the radiation stop signal generating portion90generates a radiation stop signal in Step S42ofFIG. 19, as shown inFIG. 20.

On the other hand, if the acquired data on the dose reached at the lighting field does not reach the threshold in Step S40, the process returns to Step S38and acquiring the data on the dose reached at the lighting field at the next monitoring timing and determining as to whether the radiation is stopped in Step S40are repeated until the acquired data on the dose reached at the lighting field reaches the threshold to generate a radiation stop signal in Step S42.

The procedure for generating the AEC radiation stop signal in the AEC section40ais thus finished.

After the subsequent stop of the X-ray radiation from the X-ray source12based on the radiation stop signal, under the control of the controller64b, the charges accumulated in the photodiodes48of the normal pixels45and the dual-purpose pixels45aare converted into voltage signals in the CMOS circuits51, flow into integrating amplifiers66through the signal lines58, output to an A/D72from the integrating amplifiers66with a preset sampling period as X-ray image detection signals, converted into digital X-ray image data, and temporarily stored in the memory38a. In this case, the dual-purpose pixels45afunctions as normal pixels. Accordingly, in the image detection device18c, the dual-purpose pixels45ado not cause pixel defects as in the control pixels76and76aof the image detection devices18,18aand18b, and hence high-quality X-ray images can be obtained as compared with the image detection devices18,18aand18b.

The digital X-ray image data temporarily stored in the memory38ain this way is output from the wired communication section134of the image detection device18c, transmitted to the detection controller126through the wired communication section130of the control unit124. The X-ray image data is subjected to various image processing steps in the various image processing circuits to generate a sheet of X-ray image. The X-ray image is displayed on the display30of the control unit124and is used in, for example, diagnosis.

The same effects can also be achieved in the configuration of the control unit20by configuring the image detection device18shown inFIG. 3using the normal pixels45and the dual-purpose pixels45ain place of the normal pixels44and the control pixels76.

INDUSTRIAL APPLICABILITY

The X-ray exposure control device having the function of controlling the exposure to X-rays, the X-ray image detection apparatus including the same, and the X-ray imaging system including the same according to the present invention can be utilized as an X-ray imaging system for use in industrial imaging for medical imaging and non-destructive inspection using X-rays.

While the X-ray exposure control device having the function of controlling the exposure to X-rays, the X-ray image detection apparatus including the same, and the X-ray imaging system including the same according to the present invention have been described above with reference to various embodiments and examples, it should be understood that the present invention is by no means limited to those embodiments and examples, and various improvements and design changes may of course be made without departing from the spirit and scope of the invention.