Patent Publication Number: US-2018028141-A1

Title: Radiography system, radiography method, and radiography program

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     The present application claims priority under 35 U.S.C § 119 to Japanese Patent Application No. 2016-150590, filed on Jul. 29, 2016, which is hereby expressly incorporated by reference, in its entirety, into the present application. 
     BACKGROUND 
     Technical Field 
     The present disclosure relates to a radiography system, a radiography method, and a radiography program. 
     Related Art 
     For example, as disclosed in WO2013/047193A, a radiography apparatus has been known that includes two radiation detectors each of which includes a plural pixels that accumulate a larger amount of charge as they are irradiated with a larger amount of radiation and which are provided so as to be stacked. 
     In addition, a technique has been known which detects the time related to the emission of radiation, such as the time when the emission of radiation starts and the time when the emission of radiation ends, on the basis of an electric signal of which the level generally increases as the amount of charge output from each pixel of a radiation detector of a radiography apparatus increases and controls an operation related to the accumulation of charge in each pixel. 
     However, in a case in which radiographic images are captured by two radiation detectors disclosed in, for example, WO2013/047193A, radiation that has been transmitted through the radiation detector provided on the incident side of the radiation reaches the radiation detector provided on the emission side of the radiation. Therefore, the amount of radiation that reaches the radiation detector provided on the emission side of the radiation is less than the amount of radiation that reaches the radiation detector provided on the incident side and the amount of radiation used to generate a radiographic image is reduced. 
     Therefore, in the radiation detector provided on the incident side of the radiation and the radiation detector provided on the emission side of the radiation, the detection results of the time related to the emission of radiation are different from each other. As a result, in some cases, the accumulation of charge in each pixel of each radiation detector is asynchronous. 
     SUMMARY 
     The present disclosure has been made in view of the above-mentioned problems and an object of the present disclosure is to provide a technique that can synchronize the accumulation of charge even when the amount of radiation emitted to a second radiation detector is less than the amount of radiation emitted to a first radiation detector. 
     In order to achieve the object, according to an aspect of the invention, there is provided a radiography system including: a radiography apparatus including a first radiation detector in which a first plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are two-dimensionally arranged, and a second radiation detector provided on a side of the first radiation detector from which the radiation is transmitted and emitted and in which a second plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are two-dimensionally arranged; and a controller that executes a process, the process including: obtaining an electric signal, which is converted from charge generated in the first plural pixels and of which the level increases as the amount of charge increases; detecting a time related to the emission of the radiation from the obtained electric signal; and controlling a first charge accumulation operation in the first plural pixels and a second charge accumulation operation in the second plural pixels on the basis of the detected time. 
     In order to achieve the object, according to another aspect of the present disclosure, there is provided a radiography method that is performed by a radiography apparatus including a first radiation detector in which a first plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are two-dimensionally arranged, and a second radiation detector which is provided on a side of the first radiation detector from which the radiation is transmitted and emitted and in which a first plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are two-dimensionally arranged, the method including: obtaining an electric signal, which is converted from charge generated in the first plural pixels and of which the level increases as the amount of charge increases; detecting a time related to the emission of the radiation from the obtained electric signal; and controlling a first charge accumulation operation in the first plural pixels and a second charge accumulation operation in the first plural pixels on the basis of the detected time. 
     In order to achieve the object, according to still another aspect of the present disclosure, there is provided A non-transitory computer readable storage medium storing a radiography program that causes a computer to execute a process of controlling a radiography apparatus, the radiography apparatus including: a first radiation detector in which a first plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are two-dimensionally arranged, and a second radiation detector which is provided on a side of the first radiation detector from which the radiation is transmitted and emitted and in which a first plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are two-dimensionally arranged, and the process including: obtaining an electric signal, which is converted from charge generated in the first plural pixels and of which the level increases as the amount of charge increases; detecting a time related to the emission of the radiation from the obtained electric signal; and controlling a first charge accumulation operation in the first plural pixels and a second charge accumulation operation in the first plural pixels on the basis of the detected time. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a block diagram illustrating an example of the structure of a radiography system according to an embodiment. 
         FIG. 2  is a side cross-sectional view illustrating an example of the structure of a radiography apparatus according to this embodiment. 
         FIG. 3  is a block diagram illustrating an example of the structure of a main portion of an electric system of the radiography apparatus according to this embodiment. 
         FIG. 4  is a circuit diagram illustrating an example of the structure of a signal processing unit according to this embodiment. 
         FIG. 5  is a block diagram illustrating an example of the structure of a main portion of an electric system of a console according to this embodiment. 
         FIG. 6  is a graph illustrating the amount of radiation that reaches each of a first radiation detector and a second radiation detector according to this embodiment. 
         FIG. 7  is a flowchart illustrating an example of the flow of an overall imaging process according to this embodiment. 
         FIG. 8  is a flowchart illustrating an example of the flow of an image generation process in the overall imaging process according to this embodiment. 
         FIG. 9  is a front view schematically illustrating a bone tissue region and a soft tissue region according to this embodiment. 
         FIG. 10  is a timing chart illustrating an example of the flow of the capture of a radiographic image by the radiography apparatus  16  according to this embodiment. 
         FIG. 11  is a diagram schematically illustrating a change in the amount of radiation emitted from a radiation source over an irradiation time. 
         FIG. 12  is a flowchart illustrating an example of the flow of an accumulation synchronization process according to this embodiment. 
         FIG. 13  is a flowchart illustrating an example of the flow of a first imaging process according to this embodiment. 
         FIG. 14  is a flowchart illustrating an example of the flow of a second imaging process according to this embodiment. 
         FIG. 15  is a timing chart illustrating another example of oversampling in the second radiation detector according to this embodiment. 
         FIG. 16  is a timing chart illustrating still another example of the oversampling in the second radiation detector according to this embodiment. 
         FIG. 17  is a timing chart illustrating an example of a reading method for collectively reading charge from pixels connected to a plurality of adjacent gate lines. 
         FIG. 18  is a timing chart illustrating an example of a reading method for collectively reading charge from pixels connected to a plurality of adjacent data lines. 
     
    
    
     DETAILED DESCRIPTION 
     Hereinafter, an embodiment of the invention will be described in detail with reference to the drawings. 
     First, the structure of a radiography system  10  according to this embodiment will be described with reference to  FIG. 1 . As illustrated in  FIG. 1 , the radiography system  10  includes a radiation emitting apparatus  12 , a radiography apparatus  16 , and a console  18 . The console  18  according to this embodiment is an example of an image processing apparatus according to the invention. 
     The radiation emitting apparatus  12  according to this embodiment includes a radiation source  14  that irradiates a subject W, which is an example of an imaging target, with radiation R such as X-rays. An example of the radiation emitting apparatus  12  is a treatment cart. A method for instructing the radiation emitting apparatus  12  to emit the radiation R is not particularly limited. For example, in a case in which the radiation emitting apparatus  12  includes an irradiation button, a user, such as a doctor or a radiology technician, may press the irradiation button to instruct the emission of the radiation R such that the radiation R is emitted from the radiation emitting apparatus  12 . In addition, for example, the user may operate the console  18  to instruct the emission of the radiation R such that the radiation R is emitted from the radiation emitting apparatus  12 . 
     When receiving a command to start the emission of the radiation R, the radiation emitting apparatus  12  emits the radiation R from the radiation source  14  according to exposure conditions, such as a tube voltage, a tube current, and an irradiation period. 
     The radiography apparatus  16  according to this embodiment includes a first radiation detector  20 A and a second radiation detector  20 B that detect the radiation R which has been emitted from the radiation emitting apparatus  12  and then transmitted through the subject W. The radiography apparatus  16  captures radiographic images of the subject W using the first radiation detector  20 A and the second radiation detector  20 B. Hereinafter, in a case in which the first radiation detector  20 A and the second radiation detector  20 B do not need to be distinguished from each other, they are generically referred to as “radiation detectors  20 ”. 
     Next, the structure of the radiography apparatus  16  according to this embodiment will be described with reference to  FIG. 2 . As illustrated in  FIG. 2 , the radiography apparatus  16  includes a plate-shaped housing  21  that transmits the radiation R and has a waterproof, antibacterial, and airtight structure. The housing  21  includes the first radiation detector  20 A, the second radiation detector  20 B, a radiation limitation member  24 , a control board  25 , a control board  26 A, a control board  26 B, and a case  28 . 
     The first radiation detector  20 A is provided on the incident side of the radiation R and the second radiation detector  20 B is provided so as to be stacked on the side of the first radiation detector  20 A from which the radiation R is transmitted and emitted in the radiography apparatus  16 . The first radiation detector  20 A includes a thin film transistor (TFT) substrate  30 A and a scintillator  22 A which is an example of a light emitting layer that is irradiated with the radiation R and emits light corresponding to the amount of radiation R emitted. The TFT substrate  30 A and the scintillator  22 A are stacked in the order of the TFT substrate  30 A and the scintillator  22 A from the incident side of the radiation R. The term “stacked” means a state in which the first radiation detector  20 A and the second radiation detector  20 B overlap each other in a case in which the first radiation detector  20 A and the second radiation detector  20 B are seen from the incident side or the emission side of the radiation R in the radiography apparatus  16  and it does not matter how they overlap each other. For example, the first radiation detector  20 A and the second radiation detector  20 B, or the first radiation detector  20 A, the radiation limitation member  24 , and the second radiation detector  20 B may overlap while coming into contact with each other or may overlap with a gap therebetween in the stacking direction. 
     The second radiation detector  20 B includes a TFT substrate  30 B and a scintillator  22 B which is an example of the light emitting layer. The TFT substrate  30 B and the scintillator  22 B are stacked in the order of the TFT substrate  30 B and the scintillator  22 B from the incident side of the radiation R. 
     That is, the first radiation detector  20 A and the second radiation detector  20 B are so-called irradiation side sampling (ISS) radiation detectors that are irradiated with the radiation R from the side of the TFT substrates  30 A and  30 B. 
     In the radiography apparatus  16  according to this embodiment, the scintillator  22 A of the first radiation detector  20 A and the scintillator  22 B of the second radiation detector  20 B have different compositions. Specifically, for example, the scintillator  22 A includes CsI (Tl) (cesium iodide having thallium added thereto) as a main component and the scintillator  22 B includes gadolinium oxysulfide (GOS) as a main component. GOS has a higher sensitivity to the high-energy radiation R than CsI. In addition, a combination of the composition of the scintillator  22 A and the composition of the scintillator  22 B is not limited to the above-mentioned example and may be a combination of other compositions or a combination of the same compositions. 
     The radiation limitation member  24  that limits the transmission of the radiation R is provided between the first radiation detector  20 A and the second radiation detector  20 B. An example of the radiation limitation member  24  is a metal plate made of, for example, copper or tin. It is preferable that a variation in the thickness of the radiation limitation member  24  in the incident direction of the radiation R is equal to or less than 1% in order to uniformize limitations (transmissivity) on the radiation. 
     An electronic circuit, such as an integrated control unit  71  (see  FIG. 3 ) which will be described below, is formed on the control board  25 . The control board  26 A is provided so as to correspond to the first radiation detector  20 A and electronic circuits, such as an image memory  56 A and a control unit  58 A which will be described below, are formed on the control board  26 A. The control board  26 B is provided so as to correspond to the second radiation detector  20 B and electronic circuits, such as an image memory  56 B and a control unit  58 B which will be described below, are formed on the control board  26 B. The control board  25 , the control board  26 A, and the control board  26 B are provided on the side of the second radiation detector  20 B which is opposite to the incident side of the radiation R. 
     As illustrated in  FIG. 2 , the case  28  is provided at a position (that is, outside the range of an imaging region) that does not overlap the radiation detector  20  at one end of the housing  21 . For example, a power supply unit  70  which will be described below is accommodated in the case  28 . The installation position of the case  28  is not particularly limited. For example, the case  28  may be provided at a position that overlaps the radiation detector  20  on the side of the second radiation detector  20 B which is opposite to the incident side of the radiation R. 
     Next, the structure of a main portion of an electric system of the radiography apparatus  16  according to this embodiment will be described with reference to  FIG. 3 . 
     As illustrated in  FIG. 3 , a plurality of pixels  32  are two-dimensionally provided in one direction (a row direction in  FIG. 3 ) and an intersection direction (a column direction in  FIG. 3 ) that intersects the one direction on the TFT substrate  30 A. The pixel  32  includes a sensor unit  32 A, a capacitor  32 B, and a field effect thin film transistor (TFT; hereinafter, simply referred to as a “thin film transistor”)  32 C. The sensor unit  32 A according to this embodiment is an example of a conversion element according to the invention. 
     The sensor unit  32 A includes, for example, an upper electrode, a lower electrode, and a photoelectric conversion film which are not illustrated, absorbs the light emitted from the scintillator  22 A, and generates charge. The capacitor  32 B accumulates the charge generated by the sensor unit  32 A. The thin film transistor  32 C reads the charge accumulated in the capacitor  32 B and outputs the charge in response to a control signal. The charge, of which the amount increases as the amount of radiation emitted increases, is accumulated in the pixel  32  according to this embodiment by the above-mentioned structure. 
     A plurality of gate lines  34  which extend in the one direction and are used to turn on and off each thin film transistor  32 C are provided on the TFT substrate  30 A. In addition, a plurality of data lines  36  which extend in the intersection direction and to which the charge read by the thin film transistors  32 C in an on state is output are provided on the TFT substrate  30 A. 
     A gate line driver  52 A is provided on one side of two adjacent sides of the TFT substrate  30 A and a signal processing unit  54 A is provided on the other side. Each gate line  34  of the TFT substrate  30 A is connected to the gate line driver  52 A and each data line  36  of the TFT substrate  30 A is connected to the signal processing unit  54 A. 
     The thin film transistors  32 C corresponding to each gate line  34  on the TFT substrate  30 A are sequentially turned on (in units of row illustrated in  FIG. 3  in this embodiment) by control signals which are supplied from the gate line driver  52 A through the gate lines  34 . The charge which is read by the thin film transistor  32 C in an on state is transmitted as an electric signal through the data line  36  and is input to the signal processing unit  54 A. In this way, charge is sequentially read from each gate line  34  (in units of row illustrated in  FIG. 3  in this embodiment) and image data indicating a two-dimensional radiographic image is generated by the signal processing unit  54 A. 
     As illustrated in  FIG. 4 , the signal processing unit  54 A includes a variable gain pre-amplifier (charge amplifier)  82  and a sample-and-hold circuit  84  which correspond to each data line  36 . 
     The variable gain pre-amplifier  82  includes an operational amplifier  82 A that has a positive input side grounded and a capacitor  82 B and a reset switch  82 C that are connected in parallel to each other between a negative input side and an output side of the operational amplifier  82 A. The reset switch  82 C is turned on and off by the control unit  58 A. The variable gain pre-amplifier  82  according to this embodiment is an example of an amplifier according to the invention. 
     In addition, the signal processing unit  54 A according to this embodiment includes a multiplexer  86  and an analog/digital (A/D) converter  88 . The sampling time of the sample-and-hold circuit  84  and the turn-on and turn-off of a switch  86 A provided in the multiplexer  86  are controlled by the control unit  58 A. 
     When a radiographic image is detected, first, the control unit  58 A maintains the reset switch  82 C of the variable gain pre-amplifier  82  in an on state for a predetermined period to release the charge accumulated in the capacitor  82 B. 
     When the connected thin film transistor  32 C is turned on, the charge that is accumulated in the capacitor  32 B of each pixel  32  irradiated with the radiation R is transmitted as an electric signal through the connected data line  36 . The electric signal transmitted through the data line  36  is amplified at a predetermined gain by the corresponding variable gain pre-amplifier  82 . 
     After the above-mentioned discharging is performed, the control unit  58 A drives the sample-and-hold circuit  84  for a predetermined period such that the level of the electric signal amplified by the variable gain pre-amplifier  82  is held and sampled by the sample-and-hold circuit  84 . 
     Then, the signal levels sampled by each sample-and-hold circuit  84  are sequentially selected by the multiplexer  86  and are then converted into digital signal levels by the A/D converter  88  under the control of the control unit  58 A. In this way, image data indicating the captured radiographic image is acquired. 
     The signal processing unit  54 B of the second radiation detector  20 B and the signal processing unit  54 A of the first radiation detector  20 A have the same structure except that the gains of the variable gain pre-amplifiers  82  are different from each other. The description of the same structure will not be repeated here. 
     In the radiography apparatus  16  according to this embodiment, since the first radiation detector  20 A and the radiation limitation member  24  absorb the radiation R, the amount of radiation that reaches the second radiation detector  20 B is less than the amount of radiation that reaches the first radiation detector  20 A. Therefore, the amount of charge generated in each pixel  32  of the second radiation detector  20 B is less than the amount of charge generated in each pixel  32  of the first radiation detector  20 A. 
     Therefore, in the radiography apparatus  16  according to this embodiment, the gain of the variable gain pre-amplifier  82  in the signal processing unit  54 B of the second radiation detector  20 B is higher than the gain of the variable gain pre-amplifier  82  in the signal processing unit  54 A of the first radiation detector  20 A. The amount of radiation R that is absorbed before reaching the second radiation detector  20 B varies depending on, for example, the material forming the radiation limitation member  24 . In a case in which the gain of the variable gain pre-amplifier  82  is too high, the capacitor  82 B is likely to be saturated. Therefore, specifically, the gain of the variable gain pre-amplifier  82  may be a value that is obtained in advance by, for example, experiments and is in the range in which the capacitor  82 B is not saturated. For example, it is preferable that the gain of the variable gain pre-amplifier  82  in the second radiation detector  20 B is 2 to 10 times higher than the gain of the variable gain pre-amplifier  82  in the first radiation detector  20 A, considering, for example, the material forming the radiation limitation member  24 . 
     A method for setting the gain of the variable gain pre-amplifier  82  in the second radiation detector  20 B to be higher than the gain of the variable gain pre-amplifier  82  in the first radiation detector  20 A is not particularly limited. For example, since the gain of the variable gain pre-amplifier  82  increases as the capacitance of the capacitor  82 B increases, the capacitance of the capacitor  82 B in the variable gain pre-amplifier  82  of the second radiation detector  20 B may be higher than that in the first radiation detector  20 A. In addition, the gain of the variable gain pre-amplifier  82  in the second radiation detector  20 B may be variable. For example, a plurality of series circuits, each of which includes a switch and a capacitor, may be connected in parallel to the capacitor  82 B (operational amplifier  82 A) and the switches may be turned on and off to change the number of capacitors connected to the operational amplifier  82 A. In this way, the gain is changed. 
     The gain of the variable gain pre-amplifier  82  in the second radiation detector  20 B in a case in which the signal processing unit  54 B generates image data indicating the radiographic image captured by the second radiation detector  20 B may be higher than the gain of the variable gain pre-amplifier  82  in the first radiation detector  20 A in a case in which the signal processing unit  54 A generates image data indicating the radiographic image captured by the first radiation detector  20 A. In other cases, the gain is not particularly limited. 
     The image memory  56 A is connected to the signal processing unit  54 A through the control unit  58 A. The image data output from the A/D converter  88  of the signal processing unit  54 A is sequentially output to the control unit  58 A. The image memory  56 A is connected to the control unit  58 A. The image data sequentially output from the signal processing unit  54 A is sequentially stored in the image memory  56 A under the control of the control unit  58 A. The image memory  56 A has memory capacity that can store a predetermined amount of image data. Whenever a radiographic image is captured, captured image data is sequentially stored in the image memory  56 A. In addition, the image memory  56 A is connected to the control unit  58 A. 
     The control unit  58 A includes a central processing unit (CPU)  60 , a memory  62  including, for example, a read only memory (ROM) and a random access memory (RAM), and a non-volatile storage unit  64  such as a flash memory. An example of the control unit  58 A is a microcomputer. 
     The integrated control unit  71  includes a CPU  72 , a memory  74  including, for example, a ROM and a RAM, and a non-volatile storage unit  76  such as a flash memory. An example of the integrated control unit  71  is a microcomputer. The control unit  58 A and the integrated control unit  71  are connected such that they can communicate with each other. 
     The integrated control unit  71  according to this embodiment has a function that determines whether the emission of the radiation R has started on the basis of whether the value of the digital signal output from the control unit  58 A is equal to or greater than a predetermined threshold value and controls the control unit  58 A and the control unit  58 B such that the control unit  58 A and the control unit  58 B control an operation of accumulating charge in each pixel  32  and start the accumulation of the charge in a case in which it is determined that the emission of the radiation R has started, which will be described in detail below. 
     A communication unit  66  is connected to the control unit  58 A and the integrated control unit  71  and transmits and receives various kinds of information to and from external apparatuses, such as the radiation emitting apparatus  12  and the console  18 , using at least one of wireless communication or wired communication. The power supply unit  70  supplies power to each of the above-mentioned various circuits or elements (for example, the gate line driver  52 A, the signal processing unit  54 A, the image memory  56 A, the control unit  58 A, the communication unit  66 , and the integrated control unit  71 ). In  FIG. 3 , lines for connecting the power supply unit  70  to various circuits or elements are not illustrated in order to avoid complication. 
     Components of the TFT substrate  30 B, the gate line driver  52 B, the signal processing unit  54 B, the image memory  56 B, and the control unit  58 B of the second radiation detector  20 B have the same structures as the corresponding components of the first radiation detector  20 A and thus the description thereof will not be repeated here. The control unit  58 A and the control unit  58 B are connected such that they can communicate with each other. 
     According to the above-mentioned structure, the radiography apparatus  16  according to this embodiment captures radiographic images using the first radiation detector  20 A and the second radiation detector  20 B. 
     Next, the structure of the console  18  according to this embodiment will be described with reference to  FIG. 5 . As illustrated in  FIG. 5 , the console  18  includes a control unit  90 . The control unit  90  includes a CPU  90 A that controls the overall operation of the console  18 , a ROM  90 B in which, for example, various programs or various parameters are stored in advance, and a RAM  90 C that is used as, for example, a work area when the CPU  90 A executes various programs. 
     In addition, the console  18  includes a non-volatile storage unit  92  such as a hard disk drive (HDD). The storage unit  92  stores and holds image data indicating a radiographic image captured by the first radiation detector  20 A, image data indicating a radiographic image captured by the second radiation detector  20 B, and various other data. Hereinafter, the radiographic image captured by the first radiation detector  20 A is referred to as a “first radiographic image” and image data indicating the first radiographic image is referred to as “first radiographic image data”. In addition, hereinafter, the radiographic image captured by the second radiation detector  20 B is referred to as a “second radiographic image” and image data indicating the second radiographic image is referred to as “second radiographic image data”. In a case in which the “first radiographic image” and the “second radiographic image” are generically named, they are simply referred to as “radiographic images”. 
     The console  18  further includes a display unit  94 , an operation unit  96 , and a communication unit  98 . The display unit  94  displays, for example, information related to imaging and a captured radiographic image. The user uses the operation unit  96  to input, for example, a command to capture a radiographic image and a command related to image processing for a captured radiographic image. For example, the operation unit  96  may have the form of a keyboard or may have the form of a touch panel that is integrated with the display unit  94 . The communication unit  98  transmits and receives various kinds of information to and from the radiation emitting apparatus  12  and the radiography apparatus  16 , using at least one of wireless communication or wired communication. In addition, the communication unit  98  transmits and receives various kinds of information to and from external systems, such as a picture archiving and communication system (PACS) and a radiology information system (RIS), using at least one of wireless communication or wired communication. 
     The control unit  90 , the storage unit  92 , the display unit  94 , the operation unit  96 , and the communication unit  98  are connected to each other through a bus  99 . 
     As described above, in the radiography apparatus  16  according to this embodiment, the amount of radiation that reaches the second radiation detector  20 B is less than the amount of radiation that reaches the first radiation detector  20 A. In addition, the radiation limitation member  24  generally has the characteristic that it absorbs a larger number of low-energy components than high-energy components in energy forming the radiation R, which depends on the material forming the radiation limitation member  24 . Therefore, the energy distribution of the radiation R that reaches the second radiation detector  20 B has a larger number of high-energy components than the energy distribution of the radiation R that reaches the first radiation detector  20 A. 
     In this embodiment, for example, about 50% of the radiation R that has reached the first radiation detector  20 A is absorbed by the first radiation detector  20 A and is used to capture a radiographic image. In addition, about 60% of the radiation R that has passed through the first radiation detector  20 A and reached the radiation limitation member  24  is absorbed by the radiation limitation member  24 . About 50% of the radiation R that has passed through the first radiation detector  20 A and the radiation limitation member  24  and reached the second radiation detector  20 B is absorbed by the second radiation detector  20 B and is used to capture a radiographic image. 
     That is, the amount of radiation (the amount of charge generated by the second radiation detector  20 B) used to capture a radiographic image by the second radiation detector  20 B is about 20% of the amount of radiation used to capture a radiographic image by the first radiation detector  20 A. In addition, the ratio of the amount of radiation used to capture a radiographic image by the second radiation detector  20 B to the amount of radiation used to capture a radiographic image by the first radiation detector  20 A is not limited to the above-mentioned ratio. However, it is preferable that the amount of radiation used to capture a radiographic image by the second radiation detector  20 B is equal to or greater than 10% of the amount of radiation used to capture a radiographic image by the first radiation detector  20 A in terms of diagnosis. 
     The radiation R is absorbed from a low-energy component. Therefore, for example, as illustrated in  FIG. 6 , the energy components of the radiation R that reaches the second radiation detector  20 B do not include the low-energy components of the energy components of the radiation R that reaches the first radiation detector  20 A. In  FIG. 6 , the vertical axis indicates the amount of radiation R absorbed per unit area and the horizontal axis indicates the energy of the radiation R in a case in which the tube voltage of the radiation source  14  is 80 kV. In addition, in  FIG. 6 , a solid line L 1  indicates the relationship between the energy of the radiation R absorbed by the first radiation detector  20 A and the amount of radiation R absorbed per unit area. In  FIG. 6 , a solid line L 2  indicates the relationship between the energy of the radiation R absorbed by the second radiation detector  20 B and the amount of radiation R absorbed per unit area. 
     Next, the operation of the radiography system  10  according to this embodiment will be described. 
     First, the operation of the console  18  will be described.  FIG. 7  is a flowchart illustrating an example of the flow of an overall imaging process performed by the control unit  90  of the console  18 . Specifically, the CPU  90 A of the control unit  90  executes an overall imaging processing program to perform the overall imaging process illustrated in  FIG. 7 . The control unit  90  executes the overall imaging processing program to function as an example of a derivation unit according to the invention. 
     In this embodiment, the overall imaging process illustrated in  FIG. 7  is performed in a case in which the control unit  90  of the console  18  acquires an imaging menu including, for example, the name of the subject W, an imaging part, and the emission conditions of the radiation R from the user through the operation unit  96 . The control unit  90  may acquire the imaging menu from an external system, such as an RIS, or may acquire the imaging menu input by the user through the operation unit  96 . 
     In Step S 100  of  FIG. 7 , the control unit  90  of the console  18  transmits information included in the imaging menu as an imaging start command to the radiography apparatus  16  through the communication unit  98  and transmits the emission conditions of the radiation R to the radiation emitting apparatus  12  through the communication unit  98 . 
     Then, in Step S 102 , the control unit  90  transmits a command to start the emission of the radiation R to the radiation emitting apparatus  12  through the communication unit  98 . When receiving the emission conditions and the emission start command transmitted from the console  18 , the radiation emitting apparatus  12  starts the emission of the radiation R according to the received emission conditions. The radiation emitting apparatus  12  may include an irradiation button. In this case, the radiation emitting apparatus  12  receives the emission conditions and the emission start command transmitted from the console  18  and starts the emission of the radiation R according to the received emission conditions in a case in which the irradiation button is pressed. 
     In the radiography apparatus  16 , the first radiation detector  20 A captures the first radiographic image and the second radiation detector  20 B captures the second radiographic image, on the basis of the information in the imaging menu transmitted from the console  18 , in response to the imaging start command, which will be described in detail below. In the radiography apparatus  16 , the control units  58 A and  58 B perform various correction processes, such as offset correction and gain correction, for first radiographic image data indicating the captured first radiographic image and second radiographic image data indicating the captured second radiographic image, respectively, and store the corrected radiographic image data in the storage unit  64 . 
     Then, in Step S 104 , the control unit  90  determines whether the capture of the radiographic images has ended in the radiography apparatus  16 . A method for determining whether the capture of the radiographic images has ended is not particularly limited. For example, each of the control units  58 A and  58 B of the radiography apparatus  16  transmits end information indicating that imaging has ended to the console  18  through the communication unit  66 . In a case in which the end information is received, the control unit  90  of the console  18  determines that the capture of the radiographic images has ended in the radiography apparatus  16 . 
     For example, each of the control units  58 A and  58 B transmits the first radiographic image data and the second radiographic image data to the console  18  through the communication unit  66  after imaging ends. In a case in which the first radiographic image data and the second radiographic image data are received, the control unit  90  determines that the capture of the radiographic images by the radiography apparatus  16  has ended. In addition, in a case in which the first radiographic image data and the second radiographic image data are received, the console  18  stores the received first radiographic image data and the received second radiographic image data in the storage unit  92 . 
     In a case in which the capture of the radiographic images by the radiography apparatus  16  has not ended, the determination result is “No” and the control unit  90  waits until the capture of the radiographic images by the radiography apparatus  16  ends. On the other hand, in a case in which the capture of the radiographic images by the radiography apparatus  16  has ended, the determination result is “Yes” and the control unit  90  proceeds to Step S 106 . 
     In Step S 106 , the control unit  90  performs an image generation process illustrated in  FIG. 8  and ends the overall imaging process. 
     Next, the image generation process performed in Step S 106  of the overall imaging process (see  FIG. 7 ) will be described with reference to  FIG. 8 . 
     In Step S 150  of  FIG. 8 , the control unit  90  of the console  18  acquires the first radiographic image data and the second radiographic image data. In a case in which the first radiographic image data and the second radiographic image data have been stored in the storage unit  92 , the control unit  90  reads and acquires the first radiographic image data and the second radiographic image data from the storage unit  92 . In a case in which the first radiographic image data and the second radiographic image data have not been stored in the storage unit  92 , the control unit  90  acquires the first radiographic image data from the first radiation detector  20 A and acquires the second radiographic image data from the second radiation detector  20 B. 
     Then, in Step S 152 , the control unit  90  generates image data indicating an energy subtraction image, using the first radiographic image data and the second radiographic image data. Hereinafter, the energy subtraction image is referred to as an “ES image” and the image data indicating the energy subtraction image is referred to as “ES image data”. 
     In this embodiment, the control unit  90  subtracts image data obtained by multiplying the first radiographic image data by a predetermined coefficient from image data obtained by multiplying the second radiographic image data by a predetermined coefficient for each corresponding pixel. The control unit  90  generates ES image data indicating an ES image in which soft tissues have been removed and bone tissues have been highlighted, using the subtraction. A method for determining the corresponding pixels of the first radiographic image data and the second radiographic image data is not particularly limited. For example, the amount of positional deviation between the first radiographic image data and the second radiographic image data, which are captured by the radiography apparatus  16  in a state in which a marker is put in advance, may be calculated from the difference between the positions of the marker in the first radiographic image data and the second radiographic image data. Then, the corresponding pixels of the first radiographic image data and the second radiographic image data may be determined on the basis of the calculated amount of positional deviation. 
     In this case, for example, the amount of positional deviation between the first radiographic image data and the second radiographic image data, which are obtained by capturing the image of both the subject W and the marker when the image of the subject W is captured, may be calculated from the difference between the positions of the marker in the first radiographic image data and the second radiographic image data. In addition, for example, the amount of positional deviation between the first radiographic image data and the second radiographic image data may be calculated on the basis of the structure of the subject W in the first radiographic image data and the second radiographic image data obtained by capturing the image of the subject W. 
     Then, in Step S 154 , the control unit  90  determines a bone tissue region (hereinafter, referred to as a “bone region”) in the ES image that is indicated by the ES image data generated in Step S 152 . In this embodiment, for example, the control unit  90  estimates the approximate range of the bone region on the basis of the imaging part included in the imaging menu. Then, the control unit  90  detects pixels that are disposed in the vicinity of the pixels, of which the differential values are equal to or greater than a predetermined value, as the pixels forming the edge (end) of the bone region in the estimated range to determine the bone region. 
     For example, as illustrated in  FIG. 9 , in Step S 154 , the control unit  90  detects the edge E of a bone region B and determines a region in the edge E as the bone region B. For example,  FIG. 9  illustrates an ES image in a case in which the image of a backbone part of the upper half of the body of the subject W is captured. 
     A method for determining the bone region B is not limited to the above-mentioned example. For example, the control unit  90  displays the ES image that is indicated by the ES image data generated in Step S 152  on the display unit  94 . The user designates the edge E of the bone region B in the ES image displayed on the display unit  94  through the operation unit  96 . Then, the control unit  90  may determine a region in the edge E designated by the user as the bone region B. 
     The control unit  90  may display an image in which the ES image and the edge E detected in Step S 154  overlap each other on the display unit  94 . In this case, in a case in which it is necessary to correct the edge E displayed on the display unit  94 , the user corrects the position of the edge E through the operation unit  96 . Then, the control unit  90  may determine a region in the edge E corrected by the user as the bone region B. 
     Then, in Step S 156 , the control unit  90  determines a soft tissue region (hereinafter, referred to as a “soft region”) in the ES image that is indicated by the ES image data generated in Step S 152 . In this embodiment, for example, the control unit  90  determines a region, which is other than the bone region B and has a predetermined area including pixels that are separated from the edge E by a distance corresponding to a predetermined number of pixels in a predetermined direction, as the soft region. For example, as illustrated in  FIG. 9 , in Step S 156 , the control unit  90  determines a plurality of (in the example illustrated in  FIG. 9 , six) soft regions S. 
     The predetermined direction and the predetermined number of pixels may be predetermined by, for example, experiments using the actual radiography apparatus  16  according to the imaging part. The predetermined area may be predetermined or may be designated by the user. In addition, for example, the control unit  90  may determine, as the soft region S, the pixels with pixel values in a predetermined range having the minimum pixel value (a pixel value corresponding to a position where the body thickness of the subject W is the maximum except the bone region B) as the lower limit in the ES image data. In addition, it goes without saying that the number of soft regions S determined in Step S 156  is not limited to that illustrated in  FIG. 9 . 
     Then, in Step S 158 , the control unit  90  corrects the ES image data generated in Step S 152  such that a variation in the ES image in each imaging operation is within an allowable range. In this embodiment, for example, the control unit  90  performs a correction process of removing image blur in the entire frequency band of the ES image data. The image data corrected in Step S 158  is used to calculate bone density in a process from Step S 160  to Step S 164  which will be described below. Therefore, hereinafter, the corrected image data is referred to as “dual-energy X-ray absorptiometry (DXA) image data”. 
     Then, in Step S 160 , the control unit  90  calculates an average value A1 of the pixel values of the bone region B in the DXA image data. Then, in Step S 162 , the control unit  90  calculates an average value A2 of the pixel values of all of the soft regions S in the DXA image data. Here, in this embodiment, for example, the control unit  90  performs weighting such that the soft region S which is further away from the edge E has a smaller pixel value and calculates the average value A2. Before the average values A1 and A2 are calculated in Step S 160  and Step S 162 , respectively, abnormal values of the pixel values of the bone region B and the pixel values of the soft region S may be removed by, for example, a median filter. 
     Then, in Step S 164 , the control unit  90  calculates the bone density of the imaging part of the subject W. In this embodiment, for example, the control unit  90  calculates the difference between the average value A1 calculated in Step S 160  and the average value A2 calculated in Step S 162 . In addition, the control unit  90  multiplies the calculated difference by a conversion coefficient for converting the pixel value into bone mass [g] to calculate the bone mass. Then, the control unit  90  divides the calculated bone mass by the area [cm 2 ] of the bone region B to calculate bone density [g/cm 2 ]. The conversion coefficient may be predetermined by, for example, experiments using the actual radiography apparatus  16  according to the imaging part. 
     Then, in Step S 166 , the control unit  90  stores the ES image data generated in Step S 152  and the bone density calculated in Step S 164  in the storage unit  92  so as to be associated with information for identifying the subject W. In addition, for example, the control unit  90  may store the ES image data generated in Step S 152 , the bone density calculated in Step S 164 , the first radiographic image data, and the second radiographic image data in the storage unit  92  so as to be associated with the information for identifying the subject W. 
     Then, in Step S 168 , the control unit  90  displays the ES image indicated by the ES image data generated in Step S 152  and the bone density calculated in Step S 164  on the display unit  94  and then ends the image generation process. 
     Next, the operation of the radiography apparatus  16  according to this embodiment will be described. 
     As described above, the radiography apparatus  16  according to this embodiment captures the first radiographic image, using the first radiation detector  20 A, and captures the second radiographic image, using the second radiation detector  20 B, in response to the imaging start command received from the console  18 . First, the entire flow of a radiographic image capture process performed by the radiography apparatus  16  will be described. 
     When the imaging start command is received, the control unit  58 A and the control unit  58 B direct the first radiation detector  20 A and the second radiation detector  20 B to perform a reset operation, respectively. Even in a state in which the first radiation detector  20 A and the second radiation detector  20 B are not irradiated with the radiation R, charge is accumulated in the pixels  32  by a dark current. For this reason, a reset operation for sweeping the accumulated charge is performed. The reset operation that is performed in the first radiation detector  20 A according to this embodiment is an example of a first reset operation according to the invention and the reset operation that is performed in the second radiation detector  20 B is an example of a second reset operation according to the invention. 
     In this embodiment, for example, as illustrated in  FIG. 10 , in a reset period, the control unit  58 A controls the gate line driver  52 A such that the gate line driver  52 A sequentially outputs an on signal to each gate line  34  of the first radiation detector  20 A from a gate line  34   1  for a predetermined period H 1 . In addition, in the reset period, the control unit  58 B controls the gate line driver  52 B such that the gate line driver  52 B sequentially outputs an on signal to each gate line  34  of the second radiation detector  20 B from a gate line  34   1  for the predetermined period H 1 .  FIG. 10  illustrates a case in which each of the first radiation detector  20 A and the second radiation detector  20 B includes n gate lines  34 . 
     In the first radiation detector  20 A, an electric signal, of which the level increases as the amount of charge output from the pixel  32  increases, is output to the integrated control unit  71  by the reset operation. The integrated control unit  71  detects the time when the emission of the radiation R has started, using the electric signal output by the reset operation. When the time when the emission of the radiation R has started is detected, the integrated control unit  71  outputs an accumulation start command to start a charge accumulation operation for generating a radiographic image to the control unit  58 A and the control unit  58 B. 
     The time when the emission of the radiation R has started according to this embodiment is an example of a time related to the emission of radiation according to the invention. As illustrated in  FIG. 11 , the amount of radiation R emitted from the radiation source  14  of the radiation emitting apparatus  12  varies depending on the irradiation time. In the radiography apparatus  16  according to this embodiment, the period from a time T 1  to a time T 2  illustrated in  FIG. 11 , which depends on the amount of radiation R that is emitted from the radiation source  14  to the radiography apparatus  16 , is an accumulation period which will be described below. Therefore, the time T 1  is detected as the time when the emission of the radiation R has started. The time when the radiation source  14  actually starts the emission of the radiation R is different from the time when the radiography apparatus  16  starts to be irradiated with the radiation R. The time T 1  is determined in terms of, for example, an error in the detection of the time. 
     For example, as illustrated in  FIG. 10 , when the accumulation start command is input, the control unit  58 A ends the reset operation and performs the accumulation operation for the accumulation period. Specifically, the control unit  58 A controls the gate line driver  52 A such that an off signal is output from the gate line driver  52 A to each gate line  34  of the first radiation detector  20 A. Then, all of the thin film transistors  32 C of the pixels  32  of the first radiation detector  20 A are turned off. Similarly, when the accumulation start command is input, the control unit  58 B ends the reset operation and controls the gate line driver  52 B such that an off signal is output from the gate line driver  52 B to each gate line  34  of the second radiation detector  20 B for the accumulation period. Then, all of the thin film transistors  32 C of the pixels  32  of the second radiation detector  20 B are turned off. 
     When the accumulation period elapses, for example, as illustrated in  FIG. 10 , for a read period, the control unit  58 A directs the gate line driver  52 A to sequentially output the on signal to each gate line  34  of the first radiation detector  20 A from the gate line  34   1  for a predetermined period H 2  which is a read time per pixel. Similarly, when the accumulation period elapses, for the read period, the control unit  58 B directs the gate line driver  52 B to sequentially output the on signal to each gate line  34  of the second radiation detector  20 B from the gate line  34   1  for a predetermined period H 3  which is a read time per pixel. 
     In this embodiment, the predetermined periods H 2  and H 3  for which the on signal is output to the gate line  34  in the read period are longer than the predetermined period H 1  for which the on signal is output to the gate line  34  in the reset period of each of the first radiation detector  20 A and the second radiation detector  20 B, which will be described in detail below. In addition, the predetermined period H 3  for which the on signal is output to the gate line  34  of the second radiation detector  20 B in the read period is longer than the predetermined period H 2  for which the on signal is output to the gate line  34  of the first radiation detector  20 A in the reset period of each of the first radiation detector  20 A and the second radiation detector  20 B. 
     In the radiography apparatus  16  according to this embodiment, the signal processing unit  54 A and the signal processing unit  54 B generate the first radiographic image data and the second radiographic image data, using the electric signals output from each pixel  32  for the read period, respectively. 
     Next, the operation of each of the integrated control unit  71 , the control unit  58 A, and the control unit  58 B will be described in detail.  FIG. 12  is a flowchart illustrating an example of the flow of an accumulation synchronization process performed by the integrated control unit  71 . Specifically, when the imaging start command is received from the console  18 , the CPU  72  of the integrated control unit  71  executes an accumulation synchronization processing program that is stored in the ROM of the memory  74  in advance to perform the accumulation synchronization process illustrated in  FIG. 12 . The accumulation synchronization processing program is an example of a program including a radiography program according to the invention. 
     In this embodiment, a case in which the integrated control unit  71  controls the start of a charge accumulation operation in the first radiation detector  20 A and the second radiation detector  20 B as an example of a charge accumulation operation control process to synchronize charge accumulation will be described. 
     In Step S 200  of  FIG. 12 , the integrated control unit  71  determines whether the digital signal (hereinafter, referred to as a “reset digital signal”) obtained by converting the electric signal output from the pixel  32  of the first radiation detector  20 A by the reset operation, using the signal processing unit  54 A, has been received from the control unit  58 A. In a case in which the reset digital signal has not been received, the determination result is “No” and the integrated control unit  71  waits until the reset digital signal is received. On the other hand, in a case in which the reset digital signal has been received, the determination result is “Yes” and the process proceeds to Step S 202 . 
     In Step S 202 , the integrated control unit  71  determines whether the value of the reset digital signal received in Step S 200  is equal to or greater than a predetermined threshold value for detecting the start of the emission of the radiation R. In a case in which the value of the reset digital signal is less than the threshold value, the determination result is “No” and the process returns to Step S 200 . On the other hand, in a case in which the value of the reset digital signal is equal to or greater than the threshold value, the determination result is “Yes” and the process proceeds to Step S 204 . As such, the integrated control unit  71  according to this embodiment uses the method that detects the time when the reset digital signal is equal to or greater than the threshold value as the time when the emission of the radiation R has started. However, a method for detecting the time when the emission of the radiation R has started is not limited thereto. For example, the time when the reset digital signal is greater than the threshold value may be detected as the time when the emission of the radiation R has started or the time when a variation in the reset digital signal per unit time is equal to or greater than a predetermined threshold value may be detected as the time when the emission of the radiation R has started. 
     In Step S 204 , the integrated control unit  71  outputs an accumulation start command to the control unit  58 A and the control unit  58 B and ends the accumulation synchronization process. 
       FIG. 13  is a flowchart illustrating an example of the flow of a first imaging process performed by the control unit  58 A of the radiography apparatus  16 . Specifically, when the imaging start command is received from the console  18 , the CPU  60  of the control unit  58 A executes a first imaging processing program that is stored in the ROM of the memory  62  in advance to perform the first imaging process illustrated in  FIG. 13 . 
     In Step S 230  of  FIG. 13 , the control unit  58 A determines whether a charge accumulation start command has been received from the integrated control unit  71 . In a case in which the accumulation start command has not been received, the determination result is “No” and the process proceeds to Step S 232 . 
     In Step S 232 , the control unit  58 A determines whether it is time to perform the reset operation. The time when the reset operation is performed is not particularly limited. For example, the reset operation may be performed whenever a predetermined period of time has elapsed since the imaging start command has been received from the console  18 . In a case in which it is not time to perform the reset operation, the determination result is “No” and the process returns to Step S 230 . On the other hand, in a case in which it is time to perform the reset operation, the determination result is “Yes” and the process proceeds to Step S 234 . 
     In Step S 234 , the control unit  58 A starts the reset operation. The electric signal generated by the charge flowing to each data line  36  in the reset operation is input to the signal processing unit  54 A, is amplified by the variable gain pre-amplifier  82 , and is converted into the reset digital signal by the A/D converter  88 . The reset digital signal is input to the control unit  58 A through the image memory  56 A. 
     Then, in Step S 236 , the control unit  58 A outputs the input reset digital signal to the integrated control unit  71  and the process returns to Step S 230 . 
     As described above, the reset digital signal that is output from the control unit  58 A to the integrated control unit  71  by the above-mentioned reset operation is used to detect the start of the emission of the radiation R. Here, the reset digital signal that is generated by the charge output from all of the pixels  32  of the first radiation detector  20 A may be output from the control unit  58 A to the integrated control unit  71  or the reset digital signal that is generated by the charge output from the pixels  32  corresponding to at least one of the gate line  34  or the data line  36  predetermined to detect the start of the emission of the radiation R may be output. 
     In a case in which the accumulation start command has been received in Step S 230 , the determination result is “Yes” and the process proceeds to Step S 238 . In a case in which the accumulation start command has been received even though the on signal has not been output to the gate line  34   n  in the reset operation started in Step S 234 , the control unit  58 A ends the reset operation, proceeds from the reset period to the accumulation period, and turns off all of the thin film transistors  32 C of the pixels  32  of the first radiation detector  20 A. 
       FIG. 10  illustrates a case in which the reset operation is performed for the pixels  32  including the thin film transistors  32 C controlled by the control signal flowing through the gate line  34   1  and then the accumulation start command is received. In this case, the on signal is not output to the gate line  34   2  and the subsequent gate lines  34 . 
     In Step S 238 , the control unit  58 A determines whether to end the accumulation of charge. A method for determining whether to end the accumulation of charge is not particularly limited. For example, in a case in which a predetermined accumulation period has elapsed since the accumulation start command has been received, the control unit  58 A may determine to end the accumulation of charge. In a case in which the predetermined accumulation period has not elapsed, the determination result is “No” and the control unit  58 A waits until the predetermined accumulation period elapses. On the other hand, in a case in which the predetermined accumulation period has elapsed, the determination result is “Yes” and the process proceeds to Step S 240 . 
     In Step S 240 , the control unit  58 A proceeds from the accumulation period to the read period and controls the gate line driver  52 A such that the gate line driver  52 A sequentially outputs the on signal to each gate line  34  of the first radiation detector  20 A for the predetermined period H 2 . Then, the lines of the thin film transistors  32 C connected to each gate line  34  are sequentially turned on and charge accumulated in each line of the capacitors  32 B flows as an electric signal to each data line  36 . Then, the electric signal that has flowed to each data line  36  is converted into digital image data by the signal processing unit  54 A and is then stored in the image memory  56 A. 
     For the read period, the charge that has been generated by irradiation with the radiation R and then accumulated is output from the pixel  32 . For the reset period, the charge generated by, for example, a dark current in a state in which the radiation R is not emitted is output from the pixel  32 . Therefore, the amount of charge output from the pixel  32  for the read period is more than that for the reset period. For this reason, in this embodiment, as illustrated in  FIG. 10 , the predetermined period H 2  in the read period is longer than the predetermined period H 1  in the reset period. It is preferable that the time required for the reset operation is short. Therefore, it is preferable that the predetermined period H 1  is short. 
     Then, in Step S 242 , the control unit  58 A performs image processing including various correction processes, such as offset correction and gain correction, for the image data stored in the image memory  56 A in Step S 240 . Then, in Step S 244 , the control unit  58 A transmits the image data (first radiographic image data) processed in Step S 242  to the integrated control unit  71  and ends the first imaging process. 
       FIG. 14  is a flowchart illustrating an example of the flow of a second imaging process performed by the control unit  58 B of the radiography apparatus  16 . Specifically, when the imaging start command is received from the console  18 , the CPU  60  of the control unit  58 B executes a second imaging processing program that is stored in the ROM of the memory  62  in advance to perform the second imaging process illustrated in  FIG. 14 . 
     In Step S 250  of  FIG. 14 , the control unit  58 B suppresses power supplied from the power supply unit  70  to the signal processing unit  54 B to change the signal processing unit  54 B to a power saving mode. In the power saving mode, power that is supplied to the entire signal processing unit  54 B may be suppressed or power that is supplied to each unit (see  FIG. 4 ) of the signal processing unit  54 B may be suppressed. Since the A/D converter  88  consumes a large amount of power, it is preferable to stop the driving of the A/D converter  88 . 
     In the power saving mode according to this embodiment, the control unit  58 B suppresses power supplied to the signal processing unit  54 B. However, the invention is not limited thereto. For example, the control unit  58 B may output a control signal for controlling the driving of each unit of the signal processing unit  54 B and some or all of the units of the signal processing unit  54 B may stop their driving or may be driven at a low speed in response to the control signal. As a result, the power supplied may be suppressed. 
     In this embodiment, the case in which the signal processing unit  54 B is changed to the power saving mode has been described. However, the invention is not limited thereto. For example, each unit, such as the image memory  56 B, that does not need to be driven in the reset operation and is considered not to have an effect on the generation of a radiographic image may be changed to the power saving mode. 
     Then, in Step S 252 , the control unit  58 B determines whether the charge accumulation start command has been received from the integrated control unit  71 . In case in which the accumulation start command has not been received, the determination result is “No” and the process proceeds to Step S 254 . 
     In Step S 254 , the control unit  58 B determines whether it is time to perform the reset operation. The time when the reset operation is performed is not particularly limited. For example, the reset operation may be performed whenever a predetermined period of time has elapsed since the imaging start command has been received from the console  18 . The reset operation in the first radiation detector  20 A may be asynchronous with the reset operation in the second radiation detector  20 B. In a case in which it is not time to perform the reset operation, the determination result is “No” and the process returns to Step S 250 . On the other hand, in a case in which it is time to perform the reset operation, the determination result is “Yes” and the process proceeds to Step S 256 . 
     In Step S 256 , the control unit  58 B starts the above-mentioned reset operation and returns to Step S 250 . In the second radiation detector  20 B, the electric signal generated by the charge that has flowed to each data line  36  in the reset operation is swept, without being converted into the reset digital signal, since the signal processing unit  54 B is in the power saving mode. Therefore, the reset digital signal is not output from the control unit  58 B to the integrated control unit  71 . 
     In a case in which the accumulation start command has been received in Step S 252 , the determination result is “Yes” and the process proceeds to Step S 258 . In a case in which the accumulation start command has been received even though the on signal has not yet been output to the gate line  34   n  in the reset operation started in Step S 256 , the control unit  58 B ends the reset operation, proceeds from the reset period to the accumulation period, and turns off all of the thin film transistors  32 C of the pixels  32  of the second radiation detector  20 B. 
     Then, in Step S 258 , the control unit  58 B stops the suppression of the supply of the power from the power supply unit  70  to the signal processing unit  54 B and returns the signal processing unit  54 B from the power saving mode. 
     Then, in Step S 260 , the control unit  58 B determines whether to end the accumulation of charge. A method for determining whether to end the accumulation of charge is not particularly limited. For example, in a case in which a predetermined accumulation period has elapsed since the accumulation start command has been received, the control unit  58 B may determine to end the accumulation of charge. In a case in which the predetermined accumulation period has not elapsed, the determination result is “No” and the control unit  58 B waits until the predetermined accumulation period elapses. On the other hand, in a case in which the predetermined accumulation period has elapsed, the determination result is “Yes” and the process proceeds to Step S 262 . 
     Then, in Step S 262 , the control unit  58 B controls the gate line driver  52 B such that the gate line driver  52 B sequentially outputs the on signal to each gate line  34  of the second radiation detector  20 B for the predetermined period H 3 . Then, the lines of the thin film transistors  32 C connected to each gate line  34  are sequentially turned on and charge accumulated in each line of the capacitors  32 B flows as an electric signal to each data line  36 . Then, the electric signal that has flowed to each data line  36  is converted into digital image data by the signal processing unit  54 B and is then stored in the image memory  56 B. 
     As described above, the amount of charge generated in each pixel  32  of the second radiation detector  20 B is less than the amount of charge generated in each corresponding pixel  32  of the first radiation detector  20 A. Therefore, in the radiography apparatus  16  according to this embodiment, so-called oversampling in which a read time per pixel for which the charge accumulated in the pixel  32  of the second radiation detector  20 B is read is longer than that in the first radiation detector  20 A. In this embodiment, for example, as illustrated in  FIG. 10 , the predetermined period H 3  is longer than the predetermined period H 2  in the first radiation detector  20 A. 
     A method for performing the oversampling is not limited to the method illustrated in  FIG. 10 . For example, as illustrated in  FIG. 15 , the control unit  58 B may continuously output the on signal to each gate line  34  for a predetermined period H 4  a plurality of times (two times in  FIG. 15 ) to perform the oversampling. In this case, the predetermined period H 2  and the predetermined period H 4  may be the same or different from each other. In addition, for example, as illustrated in  FIG. 16 , the control unit  58 B may repeatedly perform a process that sequentially outputs the on signal to all of the gate lines  34  from the gate line  34   1  to the gate line  34   n  for a predetermined period H 5  and sequentially outputs the on signal to the gate lines  34   1  for the predetermined period H 5  again. In this case, the predetermined period H 2  and the predetermined period H 5  may be the same or different from each other. 
     In this embodiment, the case in which the lines of the thin film transistors  32 C connected to each gate line  34  are sequentially turned on and charge accumulated in each line of the capacitors  32 B flows as an electric signal to each data line  36  has been described. However, a method for reading charge from the pixels  32  of the second radiation detector  20 B (for outputting the electric signal) is not limited thereto. For example, since the amount of charge generated in each pixel  32  of the second radiation detector  20 B is less than the amount of charge generated in each corresponding pixel  32  of the first radiation detector  20 A, charge may be collectively read from a plurality of adjacent pixels  32  of the second radiation detector  20 B. For example, as illustrated in  FIG. 17 , charge may be collectively read from the pixels  32  connected to each group of a plurality of adjacent gate lines  34 . For example,  FIG. 17  illustrates a case in which every two lines of the thin film transistors  32 C connected to each gate line  34  are sequentially turned on and charge accumulated in every two lines of the capacitors  32 B sequentially flows as an electric signal to each data line  36 . 
     For example, as illustrated in  FIG. 18 , charge may be collectively read from the pixels  32  to each group of a plurality of adjacent data lines  36 . For example,  FIG. 18  illustrates a case in which m data lines  36  are provided, the sample-and-hold circuit  84  of the signal processing unit  54 B samples the electric signals in every two data lines including a data line  36   1+2k  and a data line  36   2+2k  (k is an integer in the range of 0 to m/2) and the signals are selected by the switches  86 A of the multiplexer  86  and are converted into digital signal by the A/D converter  88 . 
     As such, in a case in which charge is collectively read from a plurality of adjacent pixels  32 , for example, the quality of the generated second radiographic image, for example, the resolution of the generated second radiographic image is lower than that in a case in which charge is read from each pixel  32 . However, as described above, in a case in which bone density is derived, bone density is preferably derived, not using the image indicated by DXA image data, but using the pixel value. Therefore, the influence of the reduction in image quality is small. 
     The amount of charge generated in each pixel  32  of the second radiation detector  20 B is less than the amount of charge generated in each pixel  32  of the first radiation detector  20 A and image quality is likely to be affected by noise. Therefore, the control unit  58 B may adjust the gain of the variable gain pre-amplifier  82  of the signal processing unit  54 B to reduce the influence of noise. In general, noise is generated due to a dark current in both stages before and behind the variable gain pre-amplifier  82  and the influence of noise caused by the radiation R overlaps noise generated in the stage before the variable gain pre-amplifier  82 . Therefore, the gain of the variable gain pre-amplifier  82  is adjusted to adjust the ratio of noise generated in the stage before the variable gain pre-amplifier  82  and noise generated in the stage behind the variable gain pre-amplifier  82 . As a result, it is possible to adjust the influence of noise in the stages before and behind the variable gain pre-amplifier  82 . For example, when the gain of the variable gain pre-amplifier  82  increases, the influence of noise generated in the stage behind the variable gain pre-amplifier  82  is reduced. In addition, it goes without saying that the gain is adjusted in the range in which the capacitor  82 B of the variable gain pre-amplifier  82  is not saturated. 
     Then, in Step S 264 , the control unit  58 B performs image processing including various correction processes, such as offset correction and gain correction, for the image data stored in the image memory  56 B in Step S 262 . Then, in Step S 268 , the control unit  58 B transmits the image data (second radiographic image data) processed in Step S 264  to the integrated control unit  71  and ends the second imaging process. 
     As described above, the radiography system  10  according to this embodiment includes: the radiography apparatus  16  including the first radiation detector  20 A in which a plurality of pixels  32 , each of which includes the sensor unit  32 A that generates a larger amount of charge as it is irradiated with a larger amount of radiation R, are two-dimensionally arranged and the second radiation detector  20 B which is provided so as to be stacked on the side of the first radiation detector  20 A from which the radiation R is transmitted and emitted and in which a plurality of pixels  32 , each of which includes the sensor unit  32 A that generates a larger amount of charge as it is irradiated with a larger amount of radiation R, are two-dimensionally arranged; and the integrated control unit  71  that controls a charge accumulation operation in the plurality of pixels  32  of the first radiation detector  20 A and a charge accumulation operation in the plurality of pixels  32  of the second radiation detector  20 B, on the basis of the detection result of the time related to the emission of the radiation R using an electric signal which is obtained by converting charge generated in the pixels  32  of the first radiation detector  20 A and of which the level increases as the amount of charge generated increases. 
     In the radiography apparatus  16  according to this embodiment, the amount of radiation that reaches the second radiation detector  20 B is less than the amount of radiation that reaches the first radiation detector  20 A. Therefore, the detection results of the time related to the emission of radiation in the first radiation detector  20 A and the second radiation detector  20 B are different from each other and the accumulation of charge in each pixel  32  of each of the radiation detectors is likely to be asynchronous. For this reason, in the radiography apparatus  16  according to this embodiment, in a case in which the time when the emission of the radiation R starts is detected by the electric signal output from the pixel  32  of the first radiation detector  20 A, the accumulation start command is output to the first radiation detector  20 A and the second radiation detector  20 B to control the accumulation operation in the first radiation detector  20 A and the second radiation detector  20 B. 
     Therefore, according to the radiography system  10  according to each of the above-described embodiments, even when the amount of radiation R emitted to the second radiation detector  20 B is less than the amount of radiation R emitted to the first radiation detector  20 A, it is possible to synchronize the accumulation of charge. 
     In this embodiment, since the electric signal output from the pixel  32  of the second radiation detector  20 B is not used to detect the start of the emission of radiation, it is possible to suppress power supplied from the power supply unit  70  for the reset period and to change the signal processing unit  54 B to the power saving mode. Therefore, according to the radiography apparatus  16  according to this embodiment, it is possible to reduce power consumption. In particular, since the driving of the A/D converter  88  with a large amount of power consumption is stopped, it is possible to further reduce power consumption. In a case in which the A/D converter  88  is driven, power consumption increases and the amount of heat generated increases, which results in an increase in temperature around the A/D converter  88 . Therefore, there is a concern that noise will be generated. However, in this embodiment, since the driving of the A/D converter  88  is stopped, it is possible to suppress the generation of noise caused by an increase in temperature. 
     In this embodiment, the case in which the integrated control unit  71  detects the time when the emission of the radiation R starts as the time related to the emission of the radiation R has been described. However, the invention is not limited thereto. For example, the integrated control unit  71  may detect the time when the emission of the radiation R is stopped like the time T 2  illustrated in  FIG. 11 . In this case, for example, the integrated control unit  71  compares the value of the reset digital signal with a predetermined threshold value for detecting the stop of the emission of the radiation R. In a case in which the value of the reset digital signal is less than the threshold value, the integrated control unit  71  may determine that it is time to stop the emission of the radiation R. In addition, in a case in which the time when the emission of the radiation R is stopped is detected in this way, the integrated control unit  71  may output a command to end the charge accumulation operation to the control unit  58 A and the control unit  58 B. In this case, in a case in which the command is input, the control unit  58 A and the control unit  58 B end the accumulation period and proceed to the read period. Therefore, it is possible to synchronize the end of the accumulation period. 
     In this embodiment, the case in which an indirect-conversion-type radiation detector that converts radiation into light and converts the converted light into charge is applied to both the first radiation detector  20 A and the second radiation detector  20 B has been described. However, the invention is not limited thereto. For example, a direct-conversion-type radiation detector that directly converts radiation into charge may be applied to at least one of the first radiation detector  20 A or the second radiation detector  20 B. 
     In the radiography apparatus  16  according to this embodiment, the aspect in which the reset digital signal output from the signal processing unit  54 A in the reset operation is used as the electric signal output from the pixel  32  of the first radiation detector  20 A has been described. However, the electric signal used to detect the time related to the emission of the radiation R is not limited thereto. For example, a radiation detection pixel  32  including a thin film transistor  32 C in which a source and a drain are short-circuited may be provided in the first radiation detector  20 A and an electric signal generated by charge output from the radiation detection pixel  32  may be used. 
     In this embodiment, the aspect in which, in the second radiation detector  20 B, every two lines of the thin film transistors  32 C connected to each gate line  34  are sequentially turned on for the read period and charge accumulated in each line of the capacitors  32 B sequentially flows as the electric signal to each data line  36  has been described. However, the invention is not limited thereto. For example, in both the first radiation detector  20 A and the second radiation detector  20 B, for the reset operation, as described with reference to  FIGS. 17 and 18 , charge may be collectively read from the pixels  32  connected to every group of a plurality of adjacent gate lines  34  or charge may be collectively read from the pixels  32  connected to every group of a plurality of adjacent data lines  36 . 
     In this embodiment, the case in which the irradiation side sampling radiation detectors in which the radiation R is incident from the TFT substrates  30 A and  30 B are applied to the first radiation detector  20 A and the second radiation detector  20 B, respectively, has been described. However, the invention is not limited thereto. For example, a so-called penetration side sampling (PSS) radiation detector in which the radiation R is incident from the scintillator  22 A or  22 B may be applied to at least one of the first radiation detector  20 A or the second radiation detector  20 B. 
     In this embodiment, the case in which the radiography apparatus  16  is controlled by three control units (control units  58 A,  58 B, and  71 ) has been described. However, the invention is not limited thereto. For example, the control unit  58 A may have the functions of the integrated control unit  71  or the radiography apparatus  16  may be controlled by one control unit. 
     In this embodiment, the case in which bone density is derived using the first radiographic image and the second radiographic image has been described. However, the invention is not limited thereto. For example, bone mineral content or both bone density and bone mineral content may be derived using the first radiographic image and the second radiographic image. 
     In this embodiment, the aspect in which the overall imaging processing program is stored (installed) in the ROM  90 B in advance, the accumulation synchronization processing program is stored in the memory  74  in advance, the first imaging processing program is stored in the memory  62  in advance, and the second imaging processing program is stored in the memory  62  in advance has been described. However, the invention is not limited thereto. Each of the overall imaging processing program, the accumulation synchronization process program, the first imaging processing program, and the second imaging processing program may be recorded in a recording medium, such as a compact disk read only memory (CD-ROM), a digital versatile disk read only memory (DVD-ROM), or a universal serial bus (USB) memory, and then provided. In addition, each of the overall imaging processing program, the accumulation synchronization process program, the first imaging processing program, and the second imaging processing program may be downloaded from an external apparatus through a network. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the control unit may detect a start of emission of the radiation as the time related to the emission of the radiation. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the controller may detect a time when the electric signal becomes equal to or greater than a predetermined threshold as the start of the emission of the radiation. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the controller may detect a time when a variation in the electric signal per unit time becomes equal to or greater than a predetermined threshold as the start of the emission of the radiation. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the control unit may further perform control such that a first reset operation which resets the charge accumulated in the first plural pixels and a second reset operation which resets the charge accumulated in the first plural pixels are performed at a predetermined time before the emission of the radiation starts. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the first reset operation and the second reset operation may collectively reset at least one of the charge in each pixel in a plurality of adjacent rows or the charge in each pixel in a plurality of adjacent columns. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, each of the first radiation detector and the second radiation detector may further include a signal processing unit that includes an amplifier to which the charge accumulated in the plural pixels is input as the electric signal and which amplifies the input electric signal, a sample-and-hold circuit that holds the electric signal amplified by the amplifier, and an analog/digital converter that converts the electric signal output from the sample-and-hold circuit into a digital signal, and performs a process of generating image data of a radiographic image from the input electric signal. A gain of the amplifier in the second radiation detector may be higher than a gain of the amplifier in the first radiation detector. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the second radiation detector may further include: a signal processing unit to which the charge accumulated in the first plural pixels is input as the electric signal and which performs a process of generating image data of a radiographic image from the electric signal; and a power control unit that controls the supply of power from a power supply unit which supplies power for driving the second radiation detector. The power control unit may suppress the supply of power from the power supply unit to the signal processing unit until the second radiation detector starts the accumulation of charge in the first plural pixels under the control of the control unit. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the signal processing unit may include an amplifier that amplifies the input electric signal, a sample-and-hold circuit that holds the electric signal amplified by the amplifier, and an analog/digital converter that converts the electric signal output from the sample-and-hold circuit into a digital signal. The power control unit may perform control such that the supply of power from the power supply unit to the analog digital converter is suppressed. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, after controlling the charge accumulation operation, the control unit may perform a control operation that reads the charge accumulated in the first plural pixels and a control operation that sets a read time per pixel in the second radiation detector to be longer than a read time per pixel in the first radiation detector and reads the charge accumulated in the first plural pixels. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the control unit may collectively read at least one of the charge accumulated in each pixel in a plurality of adjacent rows or the charge accumulated in each pixel in a plurality of adjacent columns. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the control unit may control at least one of the start of the charge accumulation operation or the end of the charge accumulation operation as a control process for the charge accumulation operation. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, each of the first radiation detector and the second radiation detector may include a light emitting layer that is irradiated with radiation and emits light. The plural pixels of each of the first radiation detector and the second radiation detector may receive the light, generate the charge, and accumulate the charge. The light emitting layer of the first radiation detector and the light emitting layer of the second radiation detector may have different compositions. 
     In the radiography system according to the above-mentioned aspect of the present disclosure, the light emitting layer of the first radiation detector may include CsI and the light emitting layer of the second radiation detector may include GOS. 
     The radiography system according to the above-mentioned aspect of the present disclosure may further includes a derivation unit that derives at least one of bone mineral content or bone density, using a first radiographic image captured by the first radiation detector and a second radiographic image captured by the second radiation detector. 
     According to the present disclosure, it is possible to synchronize the accumulation of charge even when the amount of radiation emitted to the second radiation detector is less than the amount of radiation emitted to the first radiation detector.