Radiographic imaging device, radiographic imaging system, method of controlling radiation detection sensitivity and program storage medium

A radiographic imaging device including: a detector that detects an irradiation start of radiation irradiated in imaging of a radiographic image; a derivation unit that derives an irradiation amount of radiation that will be irradiated within a specific period of time based on input data; a controller that makes a power supply amount to the detector smaller and lowers detection sensitivity to radiation irradiation start in the detector the larger the radiation irradiation amount derived by the derivation unit; and an imaging unit that images the radiographic image after radiation irradiation start has been detected by the detector.

CROSS-REFERENCE TO RELATED APPLICATION

This application claims priority under 35 USC 119 from Japanese Patent Application No. 2012-158928 filed on Jul. 17, 2012, the disclosure of which is incorporated by reference herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a radiographic imaging device that captures a radiographic image expressing radiation that has passed through an imaging subject, a radiographic imaging system, a method of controlling detection sensitivity to radiation irradiation start and a storage medium stored with a program.

2. Description of the Related Art

Recently, radiation detectors such as Flat Panel Detectors (FPDs) are being implemented in which a radiation sensitive layer is disposed on a Thin Film Transistor (TFT) active matrix substrate and with which radiation can be converted directly into digital data. Radiographic imaging devices that employ such radiation detectors to capture radiographic images expressing irradiated radiation are also being implemented. Conversion methods for converting radiation into electric signals used by such radiation detectors include for example indirect conversion methods, in which radiation is first converted into light with a scintillator and then the converted light is converted into charge by a photodiode, or direct conversion methods in which radiation is converted into charge with a semiconductor layer containing for example amorphous selenium. There are various materials that may be used in the semiconductor layer for each method.

In radiographic imaging devices equipped with FPDs, it is necessary to perform synchronization control between the FPD and a radiation source in order to match the start of an accumulation operation, in which the FPD accumulates signal charge, to an irradiation timing of irradiation of radiation from the radiation source. In order to synchronize the timing for the start of radiation irradiation and the timing for the start of the accumulation operation of signal charge by the FPD, a controller such as a console that controls the radiographic imaging device receives an irradiation start signal generated by an irradiation switch connected to the radiation source and supplies this signal to the radiographic imaging device as a synchronization signal. The radiographic imaging device transitions to the accumulation operation and starts imaging on receipt of this synchronization signal.

However, in cases where an imaging system is configured including a radiographic imaging device and a radiation source, sometimes a synchronization control interface installed as standard in the radiographic imaging device or the console thereof (for example cable or connector standards, synchronization signal format) is not compatible with an interface of the radiation source. Due to such issues, radiographic imaging devices are being developed that include an automatic radiation detection function, with radiation irradiation start automatically detected by the radiographic imaging device itself, without the use of a synchronization signal.

For example, Japanese Patent Application Laid-Open (JP-A) No. 2011-185622 discloses a radiographic imaging device provided with: plural radiation detection elements arrayed in a 2D formation in each region of regions partitioned by plural scan lines and plural signal lines; current detection means that detects current flowing in a bias line for applying a bias voltage to the radiation detection elements; control means that detects radiation irradiation start based on a value of the current detected by the current detection means; and memory pre-stored with change profiles of the current detected by the current detection means during reset processing of each of the radiation detection elements. The control means detects radiation irradiation start based on a value ΔV that is the value of the current detected by the current detection means during the reset processing of each of the radiation detection elements reduced by a value corresponding to a value of the current in the change profile.

In radiographic imaging devices with automatic radiation detection functions such as those described above, since the FPD cannot be forewarned of the timing of radiation irradiation, in an irradiation standby state the power is constantly in an ON state and an alert state is maintained until radiation irradiation. There is accordingly significantly increased power consumption in comparison to techniques in which imaging is synchronized to a radiation source. In particular, portable radiographic imaging devices (electronic cassettes) are often driven by a rechargeable battery, with demand to suppress the power consumption and lengthen the operating time for each recharge. However, in radiographic imaging devices, for example, the dose of radiation irradiated onto the FPD differs, for example, according to exposure conditions such as tube current and tube voltage, and such factors as the body thickness of the imaging subject. The radiation irradiation amount irradiated onto the FPD becomes smaller for a larger body thickness of imaging subject than for a smaller body thickness, and hence a higher detection sensitivity to radiation needs to be set. However, if radiation detection sensitivity is set constantly high then, in addition to a large power consumption in the irradiation standby state, there is also concern that false detection of radiation irradiation start due to such influences as noise or vibration might occur.

SUMMARY

An aspect of the present invention provides a radiographic imaging device. The radiographic imaging device includes: a detector that detects an irradiation start of radiation irradiated during imaging of a radiographic image; a derivation unit that derives an irradiation amount of radiation that will be irradiated within a specific period of time based on input data; a controller that causes a power supply amount to the detector to become smaller and detection sensitivity to irradiation start in the detector to become lower, as the radiation irradiation amount derived by the derivation unit increases; and an imaging unit that images the radiographic image after radiation irradiation start has been detected by the detector.

DETAILED DESCRIPTION

First Exemplary Embodiment

Detailed explanation follows regarding an exemplary embodiment of the present invention, with reference to the drawings. Note that in the following explanation, an example is used of a case in which the present invention is applied to a radiology information system that is a system that performs comprehensive management of data used in a hospital radiology department. Moreover, substantially the same or equivalent configuration elements or portions are appended with the same reference numerals in each of the drawings.

FIG. 1illustrates a configuration of a radiology information system (referred to below as “RIS”)100according to an exemplary embodiment of the present invention.

The RIS100is a system for managing information such as medical appointments and diagnostic records in a radiology department and configures part of a hospital information system (referred to below as “HIS”).

The RIS100includes plural imaging request terminal devices140(referred to below as “terminal devices”), an RIS server150, and radiographic imaging systems (referred to below as “imaging systems”)104. The imaging systems are installed in individual radiographic imaging rooms (or operating rooms) in a hospital. The RIS100is configured by the terminal devices140, the RIS server150. The imaging systems104are respectively connected to an in-hospital network102configured by for example a wired or wireless local area network (LAN). The RIS100configures part of the HIS disposed in the same hospital, and an HIS server that manages the HIS overall is also connected to the in-hospital network102.

The terminal devices140are for doctors or radiographers to input and browse diagnostic information and facility reservations, and to make radiographic imaging requests and imaging reservations. Each of the terminal devices140includes a personal computer with a display device, and the terminal devices140are connected so as to be capable of communicating with each other through the RIS server150and the in-hospital network102.

The RIS server150receives imaging requests from each of the terminal devices140and manages radiographic imaging schedules in the imaging systems104. The RIS server150is configured including a database150A.

The database150A is configured including: data relating to patients (imaging subjects), such as patient attribute information (for example name, sex, date of birth, age, blood type, body weight, patient identification (ID)), medical history, consultation history, and previously captured radiographic images; data relating to electronic cassettes40, described later, that are used in the imaging systems104, such as identification number (ID data), model, size, sensitivity, date of first use, and numbers of times used; and environment data representing the environment in which radiographic images are captured using the electronic cassettes40, namely the environment in which the electronic cassettes40are used (for example radiographic imaging room, operating room).

A doctor or radiographer operates the imaging systems104to perform radiographic imaging in response to an instruction from the RIS server150. Each of the imaging systems104is equipped with a radiation generator120that irradiates the patient (imaging subject) with an amount of radiation X such as X-rays (see alsoFIG. 7) from a radiation source121(see alsoFIG. 9) according to exposure conditions. Each of the imaging systems104is also provided with the electronic cassettes40, each of which have a built-in radiation detector20(see alsoFIG. 7) that absorbs the radiation X that has passed through an imaging target site of the patient (imaging subject) and generates charge, and generates image data expressing a radiographic image based on the amount of generated charge. The imaging systems104are also provided with a cradle130that is built into the electronic cassette40and charges a battery, and a console110that controls the electronic cassette40and the radiation generator120.

The console110acquires various types of data included in the database150A from the RIS server150, stores the data in a HDD116, described later, (seeFIG. 9), and uses the data as needed to control the electronic cassette40and the radiation generator120.

FIG. 2shows an example of an installed state of each of the devices configuring the imaging system104of an exemplary embodiment of the present invention in a radiographic imaging room180.

As shown inFIG. 2, an upright stand160employed when performing radiographic imaging in a standing position, and a prone table164employed when performing radiographic imaging in a prone position, are installed in the radiographic imaging room180. The space in front of the upright stand160serves as a patient (imaging subject) imaging position170when performing radiographic imaging in a standing position. The space above the prone table164serves as a patient (imaging subject) imaging position172when performing radiographic imaging in a prone position.

A holder162that holds the electronic cassette40is provided to the upright stand160. The electronic cassette40is held by the holder162when capturing a radiographic image in the standing position. Similarly, a holder166that holds the electronic cassette40is provided to the prone table164. The electronic cassette40is held by the holder166when capturing a radiographic image in the prone position.

Further, a supporting and moving mechanism124is disposed in the radiographic imaging room180. The supporting and moving mechanism124supports the radiation source121in such a way that the radiation source121is rotatable about a horizontal axis (the direction of arrow a inFIG. 2), is movable in a vertical direction (the direction of arrow b inFIG. 2), and is movable in a horizontal direction (the direction of arrow c inFIG. 2). It is accordingly possible to employ the single radiation source121to perform radiographic imaging in a standing position and in a prone position.

The cradle130includes a housing portion130A capable of housing the electronic cassette40. When not in use, the electronic cassette40is housed in the housing portion130A of the cradle130, and the built-in battery of the electronic cassette40is charged with the electronic cassette40in a housed state in the housing portion130A of the cradle130.

In the imaging system104, various types of data are transmitted and received by wireless communication between the radiation generator120and the console110and between the electronic cassette40and the console110.

The electronic cassette40is not limited to being used only in a state held by the holder162of the upright stand160or the holder166of the prone table164. Due to its portability the electronic cassette40may also be employed in a state not held by a holder, for example when imaging arm or leg regions.

Explanation follows regarding the configuration of the radiation detector20that is built into the electronic cassette40.FIG. 3is a cross-section schematically illustrating the configuration of a portion including three pixels of the radiation detector20of an exemplary embodiment of the present invention.

As shown inFIG. 3, the radiation detector20is configured by forming a TFT substrate30by forming signal output portions14, sensor portions13and a transparent insulating film7in sequence on a substrate1, and adhering a scintillator8to the TFT substrate30using for example an adhesive resin with low light absorbance characteristics. A pixel is configured by each of the signal output portions14and each of the sensor portions13.

The scintillator8is formed on the sensor portions13with the transparent insulating film7interposed therebetween. The scintillator8includes a phosphor that converts incident radiation into light and emits the light. Namely, the scintillator8absorbs radiation that has passed through the patient (imaging subject) and emits light.

The wavelength region of the light emitted by the scintillator8is preferably in the visible light range (wavelengths of 360 nm to 830 nm). The wavelength region of the light emitted by the scintillator8more preferably includes the green wavelength region in order to enable monochrome imaging by the radiation detector20.

A phosphor including cesium iodide (CsI) is preferably employed as the phosphor in the scintillator8in a case in which imaging employs X-rays for the radiation. CsI(Tl) (thallium-doped cesium iodide) with a light emission spectrum of 420 nm to 700 nm when X-rays are applied is particularly preferably employed. The emission peak wavelength in the visible light range of CsI(Tl) is 565 nm.

The sensor portions13are each configured including an upper electrode6, a lower electrode2, and a photoelectric conversion layer4that is provided between the upper electrode6and the lower electrode2. The photoelectric conversion layer4is configured by an organic photoelectric conversion material that absorbs the light emitted by the scintillator8and generates charge.

The upper electrode6is preferably configured from a conducting material that is transparent at least with respect to the light emission wavelength of the scintillator8since it is necessary to allow the light produced by the scintillator8to be incident to the photoelectric conversion layer4. Specifically, a transparent conducting oxide (TCO) is preferably employed that has high transmittance with respect to visible light and has a small resistance value. A metal thin film of Au or the like can also be used as the upper electrode6, however TCO is more preferable since the resistance value increases readily when trying to obtain a transmittance of 90% or more. For example, ITO, IZO, AZO, FTO, SnO2, TiO2, and ZnO2can be preferably used, with ITO being the most preferred from the perspectives of ease of processing, low resistance, and transparency. The upper electrode6may be configured from a single sheet common to all the pixels or may be divided per pixel.

The photoelectric conversion layer4includes an organic photoelectric conversion material, absorbs the light emitted from the scintillator8, and generates charge corresponding to the amount of light absorbed. The photoelectric conversion layer4including the organic photoelectric conversion material has a sharp absorption spectrum in the visible range, and virtually no electromagnetic waves are absorbed by the photoelectric conversion layer4other than the light emitted by the scintillator8. Noise generated as a result of radiation such as X-rays being absorbed by the photoelectric conversion layer4can accordingly be effectively suppressed.

The absorption peak wavelength of the organic photoelectric conversion material configuring the photoelectric conversion layer4is preferably as close as possible to the emission peak wavelength of the scintillator8in order for the organic photoelectric conversion material to most efficiently absorb the light emitted by the scintillator8. Ideally, the absorption peak wavelength of the organic photoelectric conversion material matches the emission peak wavelength of the scintillator8. However as long as the difference between the two is small, the organic photoelectric conversion material can adequately absorb the light emitted from the scintillator8. Specifically, the difference between the absorption peak wavelength of the organic photoelectric conversion material and the emission peak wavelength of the scintillator8with respect to radiation is preferably 10 nm or below. The difference is even more preferably 5 nm or below.

Examples of organic photoelectric conversion materials that can satisfy this condition include quinacridone organic compounds and phthalocyanine organic compounds. For example, the absorption peak wavelength in the visible range of quinacridone is 560 nm. Therefore, if quinacridone is used as the organic photoelectric conversion material and CsI(Tl) is used as the material for the scintillator8, it is possible to make the difference between the peak wavelengths 5 nm or below, and the amount of charge generated in the photoelectric conversion layer4can be substantially maximized.

The signal output portions14are formed on the surface of the substrate1below the lower electrodes2.FIG. 4schematically illustrates the configuration of one of the signal output portions14.

As shown inFIG. 4, each of the signal output portions14include a capacitor9and a field-effect thin film transistor (TFT: also referred to below simply as a “thin film transistor”)10. The capacitor9accumulates charge that has moved to the lower electrode2. The thin film transistor10reads out the charge accumulated in the capacitor9into signal lines36, described later (seeFIG. 5). The capacitor9and the thin film transistor10are disposed so as to overlap with the lower electrode2in plan view. Namely, the signal output portion14and the sensor portion13overlap in the thickness direction in each of the pixels. In order to reduce the surface area of the radiation detector20(pixels), it is desirable for the region in which the capacitor9and the thin film transistor10are formed to be completely covered by the lower electrode2.

The capacitor9is electrically connected to the corresponding lower electrode2through a wire of a conductive material that is formed penetrating an insulating film11disposed between the substrate1and the lower electrode2. Charge collected in the lower electrode2can accordingly be moved to the capacitor9.

A gate electrode15, a gate insulating film16, and an active layer (channel layer)17are stacked in the thin film transistor10. A source electrode18and a drain electrode19are formed at a specific separation from each other on the active layer17.

The active layer17may, for example, be formed by a material such as amorphous silicon, an amorphous oxide, an organic semiconductor material or carbon nanotubes. Note that the material configuring the active layer17is not limited to the above.

As examples of amorphous oxides that may be used to configure the active layer17, oxides including at least one of In, Ga, and Zn (for example In—O amorphous oxides) are preferable, oxides including at least two of In, Ga, and Zn (for example In—Zn—O amorphous oxides, In—Ga—O amorphous oxides, or Ga—Zn—O amorphous oxides) are more preferable, and oxides including In, Ga, and Zn are particularly preferable. As an In—Ga—Zn—O amorphous oxide, an amorphous oxide whose composition in a crystalline state is expressed by InGaO3(ZnO)m(where m is a natural number less than 6) is preferable, with InGaZnO4being more preferable.

Examples of organic semiconductor materials capable of configuring the active layer17include phthalocyanine compounds, pentacene, and vanadyl phthalocyanine, however there is no limitation thereto. Configurations of phthalocyanine compounds are described in detail in JP-A No. 2009-212389, so descriptions thereof will be omitted here.

By forming the active layer17of the thin film transistor10from an amorphous oxide, an organic semiconductor material, or carbon nanotubes, the active layer17does not absorb radiation such as X-rays, or this is restricted to an extremely minute amount if radiation is absorbed, so the generation of noise in the signal output portion14can be effectively suppressed.

Further, in a case in which the active layer17is formed with carbon nanotubes, the switching speed of the thin film transistor10can be increased, and the thin film transistor10can be formed having a low degree of absorption of light in the visible light range. In a case in which the active layer17is formed with carbon nanotubes, the performance of the thin film transistor10drops significantly if even a tiny amount of metal impurity is incorporated into the active layer17, so it is necessary to separate, extract, and form extremely high-purity carbon nanotubes using centrifugal separation or the like.

Here, the amorphous oxide, organic semiconductor material, or carbon nanotubes configuring the active layer17of the thin film transistor10and the organic photoelectric conversion material configuring the photoelectric conversion layer4are all capable of being formed into films at a low temperature. Consequently, the substrate1is not limited to a substrate with high heat resistance, such as a semiconductor substrate, a quartz substrate, or a glass substrate, and a flexible substrate, such as plastic, with aramid or bionanofibers can also be used. Specific flexible substrates that can be used include polyesters, such as polyethylene terephthalate, polybutylene phthalate and polyethylene naphthalate, polystyrene, polycarbonate, polyethersulphone, polyarylate, polyimide, polycyclic olefin, norbornene resin, and poly(chloro-trifluoro-ethylene). Employing a flexible substrate made of plastic can achieve a reduction in weight, which is advantageous from the perspective of for example portability.

Further, for example an insulating layer for ensuring insulation, a gas barrier layer for preventing the transmission of moisture and/or oxygen, and an undercoat layer for improving flatness or adhesion to the electrodes, may also be disposed on the substrate1.

High-temperature processes of 200 degrees or higher can be applied to aramids, so a transparent electrode material can be cured at a high temperature and given a low resistance, and aramids are also compatible with automatic packaging of driver ICs including solder reflow processes. Aramids also have a thermal expansion coefficient that is close to that of indium tin oxide (ITO) or a glass substrate, so they have little warping after manufacture and do not break easily. Further, aramids can also form a thinner substrate compared to a glass substrate or the like. An ultrathin glass substrate and an aramid may also be stacked to form a substrate.

Further, bionanofibers are composites of cellulose microfibril bundles (bacterial cellulose) produced by a bacterium (Acetobacter xylinum) and a transparent resin. Cellulose microfibril bundles have a width of 50 nm, which is a size that is 1/10 visible wavelengths, and have high strength, high elasticity, and low thermal expansion. By impregnating and hardening a transparent resin such as an acrylic resin or an epoxy resin in bacterial cellulose, bionanofibers can be obtained that exhibit a light transmittance of about 90% at a wavelength of 500 nm while including fibers at 60 to 70%. Bionanofibers have a low thermal expansion coefficient (3 to 7 ppm) comparable to silicon crystal, a strength comparable to steel (460 MPa), high elasticity (30 GPa), and are flexible, thereby enabling the substrate1to be formed thinner compared for example to a glass substrate.

FIG. 5is a plan view illustrating a configuration of the TFT substrate30configuring the radiation detector20. As shown inFIG. 5, plural pixels32each configured including the sensor portion13, the capacitor9, and the thin film transistor10are disposed on the TFT substrate30in a two-dimensional pattern in one direction (the row direction inFIG. 5) and an direction intersecting the one direction (the column direction inFIG. 5).

The TFT substrate30is disposed with plural gate lines34that extend in the one direction (the row direction) and that switch each of the thin film transistors100N and OFF, and the plural signal lines36that extend in the intersecting direction (the column direction) and that read the charges through the thin film transistors10that are in an ON state.

Moreover, the respective sensor portions13are connected to bias lines37. The bias lines37are connected to a bias voltage generator71, described later. A bias voltage is supplied from the bias voltage generator71through the bias lines37to each of the sensor portions13.

The TFT substrate30is formed in flat plate shape, and in a quadrilateral shape having four sides on its outer edges in plan view. More specifically, the TFT substrate30is formed in a rectangular shape.

The TFT substrate30includes pixels32that are employed to detect the presence or absence of radiation irradiation, and pixels32that capture a radiographic image. In the following explanation, the pixels32that detect radiation will be referred to as radiation detection pixels32A, and the remaining pixels32will be referred to as radiographic imaging pixels32B. In the electronic cassette40of the present exemplary embodiment, the start of radiation irradiation is detected using the radiation detection pixels32A.

As illustrated inFIG. 5, the sources and drains of the thin film transistors10are shorted in the radiation detection pixels32A. Accordingly, in the radiation detection pixels32A the charge accumulated in the capacitors9flows out into the signal lines36irrespective of the switching state of the thin film transistors10.

Note that the radiation detection pixels32A may be disposed with uniform distribution on the TFT substrate30. Moreover, as shown in the example inFIG. 6, the radiation detection pixels32A may be disposed at a comparatively low density in a partial region (a rectangular region centered on a central portion of an imaging region of the radiation detector20in the present exemplary embodiment)20A that includes the central portion of the imaging region, and disposed at a comparatively high density at regions peripheral thereto.

In the TFT substrate30, it is not possible to obtain radiographic image pixel data for the positions where the radiation detection pixels32A are disposed within the imaging region. Accordingly, in the TFT substrate30the radiation detection pixels32A are disposed so as to be dispersed within the imaging region, and missing pixel correction processing is executed by the console110to interpolate radiographic image pixel data for the positions where the radiation detection pixels32A are disposed, by employing pixel data obtained from the radiographic imaging pixels32B positioned peripherally to the radiation detection pixels32A.

Explanation next follows regarding the configuration of the electronic cassette40according to the present exemplary embodiment.FIG. 7is a perspective view illustrating a configuration of the electronic cassette40of an exemplary embodiment of the present invention.

As shown inFIG. 7, the electronic cassette40is equipped with a housing41that is formed from a material that allows radiation to pass through, and the electronic cassette40is configured with a waterproof and airtight structure. There is a concern that blood or other contaminants may adhere to the electronic cassette40when the electronic cassette40is used for example in an operating room. Therefore, giving the electronic cassette40a waterproof and airtight structure enables a single electronic cassette40to be used repeatedly by disinfecting the electronic cassette40as required.

A space A that accommodates various components is formed inside the housing41. The radiation detector20that detects the radiation X that has passed through the patient (imaging subject), and a lead plate43that absorbs backscattered rays of the radiation X, are disposed inside the space A in this order from an irradiated face side of the housing41that is irradiated with the radiation X.

A region corresponding to the placement position of the radiation detector20configures an imaging region41A that is capable of detecting the radiation. The face of the housing41with the imaging region41A is configured as a top plate41B of the electronic cassette40. In the electronic cassette40of the present exemplary embodiment, the radiation detector20is disposed so that the TFT substrate30is on the top plate41B side, and in the housing41the TFT substrate30is adhered to the inside face of the top plate41B (the face of the top plate41B at the opposite side of the face to which radiation is incident).

As shown inFIG. 7, a case42that accommodates a cassette controller58, described later, and a power source unit70(seeFIG. 9for both), is placed at one end side of the interior of the housing41at a position that does not overlap with the radiation detector20(outside the range of the imaging region41A).

The housing41is for example configured from carbon fiber, aluminum, magnesium, bionanofibers (cellulose microfibrils), or a composite material, in order to achieve a reduction in weight for the electronic cassette40overall.

As a composite material, for example, a material including a reinforcement fiber resin is used, with for example carbon or cellulose incorporated in the reinforcement fiber resin. Specific examples of composite materials that may be used include carbon fiber reinforced plastic (CFRP), a composite material with a structure where a foam material is sandwiched by CFRP, or a composite material in which the surface of a foam material is coated with CFRP. In the present exemplary embodiment, a composite material with a structure in which a foam material is sandwiched by CFRP is used. The strength (rigidity) of the housing41can accordingly be raised compared to a case in which the housing41is configured by a carbon element.

FIG. 8is a cross-section illustrating a configuration of the electronic cassette40. As shown inFIG. 8, support members44are disposed inside the housing41on the inner face of a back face portion41C that faces the top plate41B. The radiation detector20and the lead plate43are arrayed in this order along the radiation X application direction between the support members44and the top plate41B. The support members44support the lead plate43and, from the perspective of weight reduction and the perspective of absorbing dimensional deviation, are configured by for example a foam material.

As shown inFIG. 8, adhesive members80are provided at the inner face of the top plate41B to detachably adhere the TFT substrate30of the radiation detector20. Double-sided tape, for example, can be employed for the adhesive members80. In this case, the double-sided tape is formed in such a way that the adhesive force of one adhesive face is stronger than that of the other adhesive face.

Specifically, the face with the weaker adhesive force (weak adhesive face) is set to have a 180-degree peel strength of 1.0 N/cm or lower. The face with the stronger adhesive force (strong adhesive face) contacts the top plate41B, and the weaker adhesive face contacts the TFT substrate30. The thickness of the electronic cassette40can accordingly be made thinner than in a case in which the radiation detector20is fixed to the top plate41B by, for example, fixing members such as screws. Moreover, even if the top plate41B deforms under impact or load, the radiation detector20follows the deformation of the top plate41B that has high rigidity, so only deformation of large radius of curvature (a gentle curve) arises, reducing the likelihood of the radiation detector20sustaining damage due to localized deformation of low radius of curvature. Moreover, the radiation detector20contributes to raising the rigidity of the top plate41B.

Thus in the electronic cassette40according to the present exemplary embodiment, since the radiation detector20is adhered at the inside of the top plate41B of the housing41, the housing41is separable into two between the top plate41B side and the back face portion41C side. The housing41is placed in a state divided into the two parts of the top plate41B side and the back face portion41C side in order to adhere the radiation detector20to the top plate41B or detach the radiation detector20from the top plate41B.

In the present exemplary embodiment, adhering the radiation detector20to the top plate41B does not have to be performed for example in a clean room. This is due to the fact that even if foreign objects such as metal fragments that absorb radiation where to be incorporated between the radiation detector20and the top plate41B, such foreign objects can be removed by detaching the radiation detector20from the top plate41B.

FIG. 9is a drawing illustrating a configuration of relevant portions of an electrical system of the imaging system104of the present exemplary embodiment. As shown inFIG. 9, in the TFT substrate30configuring the radiation detector20built into the electronic cassette40, a gate line driver52is disposed on one side of two adjacent sides, and a signal processor54is disposed on the other side. The individual gate lines34of the TFT substrate30configuring the radiation detector20are connected to the gate line driver52, and the individual signal lines36of the TFT substrate30are connected to the signal processor54.

An image memory56, the cassette controller58, a wireless communication unit60, the power source unit70and a bias voltage generator71are provided inside the housing41.

Each of the thin film transistors10of the TFT substrate30are switched ON in sequence in row units by signals supplied through the gate lines34from the gate line driver52, and the charges that have been read out by the thin film transistors10being switched to an ON state are transmitted through the signal lines36as electric signals and input to the signal processor54. The charges are thereby read out in sequence by row unit, and a two-dimensional radiographic image is acquired.

The signal processor54is configured including charge amplifiers, sample-and-hold circuits, a multiplexer and an analogue-to-digital (A/D) converter. The charge amplifiers generate electric signals with a voltage level corresponding to the amount of charge read out from the sensor portions13through each of the signal lines36. The signal levels of the electric signals generated by the charge amplifiers are held by the sample-and-hold circuits. Output terminals of the sample-and-hold circuits are connected to the common multiplexer. The multiplexer converts the signal levels held by the sample-and-hold circuits into serial data and supplies this serial data to the A/D converter. The A/D converter converts the analogue electric signals supplied from the multiplexer into image data as digital signals.

The image memory56is connected to the signal processor54. The image data output from the A/D converter of the signal processor54are sequentially stored in the image memory56. The image memory56has a storage capacity that is capable of storing a predetermined number of frames' worth of image data. The image data obtained by the imaging are sequentially stored in the image memory56each time radiographic imaging is performed. The image memory56is also connected to the cassette controller58.

The bias voltage from the bias voltage generator71is applied through the bias lines37to the sensor portions13. The sensor portions13generate more charge the greater the value of the bias voltage applied. Namely, the detection sensitivity to radiation rises the greater the value of the bias voltage applied to the sensor portions13. The bias voltage generator71is a variable voltage source in which the output voltage is variable, and a bias voltage of a voltage level corresponding to a control signal applied from a cassette controller58, described later, is supplied to the sensor portions13. Namely, the detection sensitivity to radiation of the sensor portions13is controlled by the cassette controller58.

The cassette controller58performs comprehensive control of the operation of the overall electronic cassette40including output voltage control of the bias voltage generator71described above. The cassette controller58is configured including a microcomputer, and is equipped with a central processing unit (CPU)58A, a memory58B including read-only memory (ROM) and random access memory (RAM), and a nonvolatile storage unit58C configured for example by flash memory. The wireless communication unit60is connected to the cassette controller58.

The wireless communication unit60conforms to a wireless local area network (LAN) standard such as typified by the Institute of Electrical and Electronics Engineers (IEEE) 802.11a/b/g and controls the transmission of various types of data to and from external devices by wireless communication. Through the wireless communication unit60, the cassette controller58enabled for wireless communication with external devices such as the console110that performs control relating to radiographic imaging and is enabled for transmitting and receiving various types of data to and from the console110, for example. Imaging subject data and exposure conditions supplied from the console110are received by the wireless communication unit60.

The electronic cassette40is provided with the power source unit70, with the mentioned various circuits and devices (the gate line driver52, the signal processor54, the image memory56, the wireless communication unit60, the microcomputer that functions as the cassette controller58, the bias voltage generator71and the like) actuated with power supplied from the power source unit70. The power source unit70has an inbuilt battery (a rechargeable secondary battery) so as not to affect the portability of the electronic cassette40, and power is supplied to the various circuits and devices from the charged battery. Note that wiring that connects the power source unit70to the various circuits and devices is omitted from illustration inFIG. 9.

As shown inFIG. 9, the console110is configured by a server/computer, and is equipped with a display111that displays for example an operation menu and captured radiographic images, and an operation panel112that is configured including plural keys and is input with various types of information and operation instructions.

Moreover, the console110according to the present exemplary embodiment is equipped with: a CPU113that controls operation of the overall apparatus; ROM114that is pre-stored with for example various programs including a control program; RAM115that temporarily stores various data; a hard disk drive (HDD)116that stores and holds various data; a display driver117that controls the display of various information on the display111; and an operation input detector118that detects an operation state of the operation panel112. The console110is further equipped with a wireless communication unit119that employs wireless communication to transmit and receive various data such as exposure conditions, described later, between the console110and the radiation generator120, as well as transmitting and receiving various data such as image data between the console110and the electronic cassette40.

The CPU113, the ROM114, the RAM115, the HDD116, the display driver117, the operation input detector118and the wireless communication unit119are connected together through a system bus BUS. The CPU113can accordingly access the ROM114, the RAM115and the HDD116, and the CPU113can also control the display of various data on the display111through the display driver117, and control the transmission and reception through the wireless communication unit119of various data to and from the radiation generator120and the electronic cassette40. The CPU113can also ascertain the operation state of the operation panel112by a user through the operation input detector118.

The radiation generator120is equipped with the radiation source121, a wireless communication unit123that transmits and receives various data such as exposure conditions between the radiation generator120and the console110, and a controller122that controls the radiation source121based on received exposure conditions.

The controller122is also configured including a microcomputer, and stores received exposure conditions. These exposure conditions received from the console110include data such as tube voltage, tube current, and exposure duration. The controller122causes the radiation X to be irradiated from the radiation source121based on the received exposure conditions.

The electronic cassette40includes a power adjustment function that, in preparation for imaging a radiographic image, adjusts the detection sensitivity during detection of radiation irradiation start based on the imaging subject data and the exposure conditions notified from the console110by adjusting the power supply amount from the battery configuring the power source unit70. Explanation follows regarding this power adjustment function.

The cassette controller58of the electronic cassette40, in preparation for imaging a radiographic image, is notified with the imaging subject data and the exposure conditions through the wireless communication unit119of the console110. The imaging subject data includes for example data such as the imaging target side and imaging orientation of the imaging subject. The exposure conditions include for example data such as tube voltage, tube current, and exposure duration. The cassette controller58, based on the imaging subject data and the exposure conditions, derives an irradiation amount of radiation that will be irradiated from the radiation source121onto the electronic cassette40through the imaging subject within a specific period of time during imaging of a radiographic image.

The cassette controller58is equipped with a first reference table500such as that illustrated inFIG. 10inside its own storage unit58C. The cassette controller58derives the irradiation amount of radiation to be irradiated onto the radiation irradiation face (imaging face) of the electronic cassette40within the specific period of time by searching the first reference table500using as keys the imaging subject data and the exposure conditions acquired from the console110. A to D inFIG. 10indicate imaging target sites, a1to a4, b1to b4, c1to c4and d1to d4indicate radiation intensity determined by settable tube currents and tube voltages for each of the imaging target sites. X1to X8are estimated values of the approximate irradiation amount of radiation that will be irradiated onto the electronic cassette40in the specific period of time corresponding to the imaging target sites A to D and to the radiation intensities a1to d4. The irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time is smaller for imaging target sites that are sites that do not readily transmit radiation (namely that have a high attenuation coefficient) and for sites of great thickness, whereas the irradiation amount is large for opposite cases. The irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time is also smaller the smaller the value of the tube current and tube voltage set for the radiation generator120. It is accordingly possible to estimate the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time based on input data such as the imaging target site and the tube voltage and tube current. The first reference table500is accordingly stored in the storage unit58C with the radiation irradiation amounts X1to X8derived by testing or simulation, associated with the imaging subject data and the exposure conditions. Note that the first reference table500may include other parameters to those described above (parameters such as for example the imaging subject gender, age, height, weight, body fat index, and thickness of imaging target site). The cassette controller58may also compute the intensity of radiation attenuated by passing through the imaging subject based on the imaging subject data and the exposure conditions, and then compute the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time from the computed radiation intensity. The first reference table500becomes redundant in such cases, however processing time is required to compute the radiation irradiation amount.

The cassette controller58derives a value of a bias voltage to be applied to the sensor portions13based on the derived radiation irradiation amount. The cassette controller58is also equipped in its own storage unit58C with a second reference table501such as the example illustrated inFIG. 11. The second reference table501includes the irradiation amounts X1to X8of radiation that will be irradiated onto the electronic cassette40within the specific period of time associated with bias voltage values V1to V8that should be applied to the sensor portions13. The cassette controller58derives values of bias voltages that should be applied to the sensor portions13by searching the second reference table501using as keys the radiation irradiation amounts derived based on the imaging subject data and the exposure conditions.

During imaging of a radiographic image, the value of the bias voltage applied to the sensor portions13needs to be made comparatively large to raise the detection sensitivity of the sensor portions13in cases in which the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time is comparatively small. The radiation irradiation start cannot be detected accurately in cases in which the detection sensitivity is insufficient, and transition cannot be made to imaging operation despite radiation being irradiated onto the imaging subject. It is however possible to detect the radiation irradiation start even when the detection sensitivity of the sensor portions13is somewhat lowered in cases in which the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time is comparatively large. The value of the bias voltage applied to the sensor portions13in such cases can be made comparatively small.

The second reference table501is configured from irradiation amounts associated with bias voltages such that the bias voltage to be applied to the sensor portions13is smaller the larger the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time. Namely, during detection of radiation irradiation start the cassette controller58identifies the irradiation amount of radiation that will be irradiated onto the electronic cassette40and then controls the bias voltage generator71such that the bias voltage applied to the sensor portions13is smaller the larger the identified irradiation amount. The power supply amount from the battery configuring the power source unit70is thereby reduced. According to such a power adjustment function of the electronic cassette40, it is possible to achieve a reduction in power consumption in an irradiation standby state of the electronic cassette40, in comparison to cases in which a relatively high bias voltage is applied in a fixed manner irrespective of the radiation irradiation amount.

Note that it is possible to appropriately increase or decrease the number of bias voltage adjustment steps by re-writing the second reference table501. Moreover, configuration may be made such that a relationship equation between the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time and the bias voltage is pre-stored in the storage unit58C, and a cassette controller38computes the bias voltage to be applied to the sensor portions13from the relationship equation. The second reference table501becomes redundant in such cases.

Explanation next follows regarding operation of the imaging system104of the present exemplary embodiment.

First, explanation follows regarding the operation of the console110when capturing a radiographic image, with reference toFIG. 12.FIG. 12is a flow chart showing a flow of processing by a radiographic imaging processing program that is executed by the CPU113of the console110when input with an instruction to execute radiographic imaging through the operation panel112. This program is pre-stored in a predetermined region of the ROM114.

At step300inFIG. 12, the CPU113controls the display driver117so as to cause the display111to display a predetermined initial information input screen. At the next step302, the CPU113is on standby for input of specific information.

FIG. 13shows an example of the initial information input screen that is displayed on the display111by the processing of step300. As shown inFIG. 13, in the initial information input screen according to the present exemplary embodiment, a message is displayed prompting input of the name of the patient (imaging subject) on whom radiographic imaging is to be performed, the imaging target site, the posture during imaging, and the exposure conditions of the radiation X during imaging (in the present exemplary embodiment, the tube voltage, the tube current and exposure duration during radiation X exposure). Input fields for these items of information are also displayed.

After the initial information input screen shown inFIG. 13is displayed on the display111, the radiographer inputs the name of the patient (imaging subject) to be imaged, the imaging target site, the posture during imaging, and the exposure conditions into the corresponding input fields through the operation panel112.

The radiographer enters the radiographic imaging room180with the patient (imaging subject). When performing image capture in a standing position or prone position, the radiographer positions the patient (imaging subject) at a specific imaging position (performs positioning) after the electronic cassette40has been held by holder162of the upright stand160or the holder166of the prone table164as appropriate and the radiation source121has been positioned correspondingly. However in order to perform radiographic imaging with the electronic cassette40not held by a holder, such as when the imaging target site is a region of an arm or leg, the radiographer positions the patient (imaging subject) in a specific imaging position (performs positioning). However, when capturing a radiographic image of an imaging target site such as an arm or a leg without the electronic cassette40being held in the holders, the radiographer positions (performs positioning of) the patient (imaging subject), the electronic cassette40and the radiation source121in a state that allows imaging of the imaging target site.

The radiographer then exits the radiographic imaging room180, and uses the operation panel112to select the INPUT COMPLETE button displayed in the vicinity of the bottom edge of the initial information input screen. Step302is determined in the affirmative when the radiographer has selected the INPUT COMPLETE button and processing then transitions to step304.

At step304the CPU113transmits the data input to the initial information input screen (referred to below as “initial information”) to the electronic cassette40through the wireless communication unit119. Then at the next step306the exposure conditions are set by transmitting the exposure conditions included in the initial information to the radiation generator120through the wireless communication unit119. The controller122of the radiation generator120then performs preparation for exposure according to the received exposure conditions.

At the next step308, the CPU113transmits instruction data instructing the start of exposure to the radiation generator120and the electronic cassette40through the wireless communication unit119.

In response the radiation source121starts emitting the radiation X with the tube voltage and tube current corresponding to the exposure conditions the radiation generator120has received from the console110. The radiation X emitted from the radiation source121reaches the electronic cassette40after passing through the patient (imaging subject).

The cassette controller58of the electronic cassette40receives the instruction data instructing the start of exposure, and remains on standby until the radiation amount detected by the radiation detection pixels32A reaches a predetermined threshold value or greater that serves as a value for detecting that radiation irradiation has started. The electronic cassette40starts radiographic imaging operation when determination is made that the radiation amount detected by the radiation detection pixels32A has reached the threshold value or greater. The electronic cassette40ends the radiographic imaging operation after a specific accumulation duration has elapsed since the start of radiation irradiation, and then transmits the thus obtained image data to the console110.

At the next step310, the CPU113enters standby until the image data is received from the electronic cassette40, and at the next step312, image processing is performed on the received image data to perform various corrections such as shading correction after the missing pixel correction processing described above has been performed.

Then at the next step314the CPU113stores in the HDD116the image data that has been subject to image processing (referred to below as “corrected image data”). Then at the next step316the display driver117is controlled so as to display a radiographic image expressed by the corrected image data on the display111, in order for example to perform verification.

At the next step318the CPU113transmits the corrected image data to the RIS server150over the in-hospital network102, after which the radiographic imaging processing program is ended. The corrected image data transmitted to the RIS server150is stored in the database150A, thereby enabling a medical doctor to read the captured radiographic image and perform diagnostics.

Explanation follows regarding operation of the electronic cassette40when the initial information is received from the console110, with reference toFIG. 14.FIG. 14is a flow chart illustrating the flow of processing of a cassette imaging processing program executed by the CPU58A of the cassette controller58in the electronic cassette40when initial information is received from the console110. The cassette imaging program is pre-stored in a specific region of the storage unit58C of the cassette controller58.

At step400, the CPU58A awaits receipt of instruction data instructing exposure start, described above, from the console110. At the next step402the following detection sensitivity setting processing program is executed.

FIG. 15is a flow chart indicating a flow of processing of the detection sensitivity setting processing program executed in the CPU58A. This program is also pre-stored in a specific region of the storage unit58C of the cassette controller58.

At step450ofFIG. 15, the CPU58A derives the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time by searching the first reference table500(seeFIG. 10) stored in the storage unit58C using as keys the imaging target site as the imaging subject data, and the tube current and the tube voltage as the exposure conditions, that are contained in initial information supplied from the console110.

In the next step452, the CPU58A derives a value of the bias voltage that should be applied to the sensor portions13by searching the second reference table501(seeFIG. 11) stored in the storage unit58C using as a key the radiation irradiation amount derived at step450. Note that the second reference table501is configured to derive a smaller value bias voltage the larger the radiation irradiation amount derived at step450.

In the next step454, the CPU58A supplies to the bias voltage generator71the control signal required to apply the bias voltage derived at step452to the sensor portions13, so as to control the output voltage of the bias voltage generator71. The present detection sensitivity setting processing program is then ended. On receipt of the control signal, the bias voltage generator71outputs the bias voltage derived by the CPU58A, and applies this bias voltage to each of the sensor portions13.

Transition is made to step404of a cassette imaging processing program (main routine) when the detection sensitivity setting processing program has ended. At step404, the CPU58A supplies a control signal to the signal processor54and processing to read charge from the radiation detection pixels32A is executed in the signal processor54. The amount of charge read from the radiation detection pixels32A is temporarily stored in the memory58B of the cassette controller58as radiation amount data (a detection signal) indicating the dose of radiation that has been irradiated. The CPU58A acquires the radiation amount data (a detection signal) by accessing the memory58B.

In the next step406, the CPU58A determines whether or not a radiation dose indicated in the radiation amount data (a detection signal) previously acquired at step404is a predetermined specific threshold value for detecting radiation irradiation start or greater. Processing returns to step404when determination is negative at step406. However, affirmative determination is taken to mean that emission of radiation from the radiation source121has started, and processing transitions to step408. Note that configuration may be made such that at this point in time the CPU58A supplies the gate line driver52with a control signal required to perform a reset operation for discharging dark current that has accumulated in the radiation detection pixels32A and the radiographic imaging pixels32B up until radiation irradiation start has been detected. On receipt of such a control signal, the gate line driver52then supplies a drive signal sequentially to the gate lines34so as to switch the thin film transistors100N one line at a time. Any dark current that has accumulated in the radiation detection pixels32A and the radiographic imaging pixels32B is thereby discharged to the signal lines36, resetting each of the pixels.

In the next step408, the CPU58A supplies the gate line driver52with a control signal required to switch all the thin film transistors10OFF. Accordingly, accumulation starts of charge generated in the radiographic imaging pixels32B according to the irradiation of radiation, and transition is made to the imaging operation of a radiographic image. When this is performed, the CPU58A may be configured to set the bias voltage to be applied to the sensor portions13of the radiographic imaging pixels32B to a different voltage to the bias voltage derived in step452(seeFIG. 15). For example, the CPU58A may at the present step408supply the bias voltage generator71with a control signal required to set an upper limit value to a bias voltage setting range. Namely, the detection sensitivity of the sensor portions13during imaging of a radiographic image may be set higher than the detection sensitivity during radiation irradiation start detection. The CPU58A thereby controls the detection sensitivity to radiation during imaging of a radiographic image independently from the detection sensitivity during radiation irradiation start detection.

In the next step410, the CPU58A determines whether or not a specific accumulation duration has elapsed since transition to the accumulation operation. The accumulation duration is set according to the exposure duration of radiation set as the exposure conditions. The CPU58A transitions to the processing of step412when determined that the specific accumulation duration has elapsed since transition to the accumulation operation.

At the next step412, the CPU58A causes an ON signal to be output from the gate line driver52to the gate lines34in sequence one line at a time by supplying a control signal to the gate line driver52. Each of the thin film transistors10connected to each of the gate lines34is thereby switched ON in sequence one line at a time. The charge that has been accumulated in each of the capacitors9is thereby read into each of the signal lines36, and converted by the signal processor54into digital image data which is then stored in the image memory56.

At the next step414, the CPU58A reads the image data stored in the image memory56and the present cassette imaging processing program is ended after transmitting the read image data to the console110through the wireless communication unit60.

In the electronic cassette40of the present exemplary embodiment, as shown inFIG. 8, the radiation detector20is disposed such that radiation X is irradiated from the TFT substrate30side of the electronic cassette40.

In cases using what is referred to as a Penetration Side Sampling (PSS) method in which the radiation detector20is irradiated with radiation from the side on which the scintillator8is formed, as shown inFIG. 16, and radiographic images are read by the TFT substrate30provided on the opposite side to the radiation incident face, light is emitted with higher intensity from the top face side of the scintillator8inFIG. 18(the opposite side to the face joined to the TFT substrate30). However, in cases using what is referred to as an Irradiation Side Sampling (ISS) method in which radiation is irradiated from the TFT substrate30side and radiographic images are read by the TFT substrate30provided on the radiation incident face side, radiation that has passed through the TFT substrate30is incident to the scintillator8and light is emitted with higher intensity from the side of the scintillator8of the face joined to the TFT substrate30. Each of the sensor portions13provided to the TFT substrate30generates charge according to the light generated in the scintillator8. The radiographic images captured are accordingly of higher resolution when an ISS method is employed than when a PSS method is employed since the light emission position of the scintillator8is closer to the TFT substrate30.

The radiation detector20is also configured with the photoelectric conversion layer4formed from an organic photoelectric conversion material and so radiation is barely absorbed by the photoelectric conversion layer4. The radiation detector20of the present exemplary embodiment is accordingly capable of suppressing deterioration in sensitivity to radiation, since the amount of radiation absorbed by the photoelectric conversion layer4is smaller even when radiation passes through the TFT substrate30due to employing an ISS method. In an ISS method the radiation has passed through the TFT substrate30to reach the scintillator8. However application may be made to an ISS method when the photoelectric conversion layer4of the TFT substrate30is thus configured from an organic photoelectric conversion material, since there is hardly any radiation absorption in the photoelectric conversion layer4and radiation attenuation can be suppressed to a small amount.

It is also possible to form both the amorphous oxide configuring the active layer17of the thin film transistors10and the organic photoelectric conversion material configuring the photoelectric conversion layer4using film forming at low temperature. The substrate1can accordingly be formed from plastic resin with aramid and/or bionanofibers, having low absorptivity to radiation. Since the amount of radiation absorbed by the thus formed substrate1is small, sensitivity to radiation can be suppressed from deteriorating even when radiation passes through the TFT substrate30due to employing an ISS method.

According to the present exemplary embodiment, as shown inFIG. 8, the radiation detector20is attached inside the housing41to the top plate41B so that the TFT substrate30is on the top plate41B side. Moreover, the top plate41B of the housing41can be formed thinner in cases in which the substrate1is formed with high rigidity from a plastic resin with aramid and/or bionanofibers, since the rigidity of the radiation detector20itself is high. The radiation detector20is also not easily damaged in cases in which the substrate1is formed with high rigidity from a plastic resin with aramid and/or bionanofibers, even when the imaging region41A incurs an impact since the radiation detector20itself is flexible.

As explained in the above, based on the imaging subject data and the exposure conditions supplied from the console110, the electronic cassette40according to an exemplary embodiment of the present invention derives an irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time prior to radiation actually being irradiated from the radiation source121. The electronic cassette40then derives a value for the bias voltage that should be applied to the sensor portions13based on the derived radiation irradiation amount, thereby setting the detection sensitivity to radiation irradiation start. When doing so, the electronic cassette40makes the power supply amount from the power source unit70smaller the larger the derived radiation irradiation amount, thereby lowering the detection sensitivity to radiation irradiation start. Namely, the cassette controller58controls such that the value of the bias voltage is smaller the larger the derived radiation irradiation amount. The electronic cassette40thus changes the bias voltage to be applied to the sensor portions13to appropriately adjust the detection sensitivity according to the anticipated radiation irradiation amount. It is accordingly possible to prevent application of a bias voltage excessively larger than the bias voltage needed to detect the radiation irradiation start. A reduction in power consumption in the irradiation standby state is accordingly possible in comparison to cases in which an excessive bias voltage is applied in a fixed manner irrespective of the radiation irradiation amount. The electronic cassette40appropriately adjusts the detection sensitivity to radiation according to the anticipated radiation irradiation amount, and so false detection of radiation irradiation start can be prevented in comparison to cases in which a constant high sensitivity is set.

Note that in the present exemplary embodiment, an example has been illustrated in which the sensor portions13of the radiation detection pixels32A and the sensor portions13of the radiographic imaging pixels32B are connected to the common bias voltage generator71through the bias lines37, however there is no limitation thereto. Namely, as illustrated inFIG. 17, a first bias voltage generator71A for supplying a bias voltage to the sensor portions13of the radiation detection pixels32A and a second bias voltage generator71B for supplying a bias voltage to the sensor portions13of the radiographic imaging pixels32B may be separately provided. In such cases, during detection of radiation irradiation start, the output voltage of the first bias voltage generator71A that applies bias voltage to the sensor portions13of the radiation detection pixels32A is controlled according to the anticipated radiation irradiation amount similarly to in the exemplary embodiment described above. Moreover, in such cases control may be performed to stop the supply of bias voltage from the first bias voltage generator71A after radiation irradiation start has been detected by the radiation detection pixels32A. It is thereby possible to achieve an even greater reduction in power consumption. However, the output voltage of the second bias voltage generator71B for applying bias voltage to the sensor portions13of the radiographic imaging pixels32B is controlled independently from the first bias voltage generator71A so as to be a voltage appropriate to imaging a radiographic image.

Second Exemplary Embodiment

Explanation follows regarding an electronic cassette according to a second exemplary embodiment of the present invention. The electronic cassette40according to the first exemplary embodiment derives an irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time, and adjusts the power supply amount from the power source unit70by adjusting the value of the bias voltage applied to the sensor portions13according to the derived radiation irradiation amount. In the electronic cassette40according to the second exemplary embodiment, however, the power supply amount from a power source unit70is adjusted by adjusting the number of drive times of a charge amplifier configuring a signal processor54according to the derived radiation irradiation amount.

FIG. 18is a diagram illustrating a configuration of the signal processor54configuring the electronic cassette40. Note that since configuration portions other than the signal processor54are similar to those of the first exemplary embodiment, explanation thereof is omitted.

As illustrated inFIG. 18, the signal processor54includes charge amplifiers92respectively connected to each of the signal lines36. Note that one or more radiation detection pixels32A is connected to each of the signal lines36. Each of the charge amplifiers92includes: an operational amplifier (operation amplification circuit)92A with inverting input terminal connected to the respective signal line36and non-inverting input terminal connected to a ground potential; a capacitor92B with one terminal connected to the inverting input terminal of the operational amplifier92A and the other terminal connected to the output terminal of the operational amplifier92A; and a reset switch92C that is connected in parallel to the capacitor92B. Each of the charge amplifiers92also includes a switch92E between a power line92D and a power input terminal of the operational amplifier92A. Namely, the operational amplifiers92are each placed in an operational state supplied with drive voltage from the power line92D by switching the switch92E ON. ON/OFF switching of the switches92E is performed according to a control signal supplied from the cassette controller58.

The charges generated in each of the radiation detection pixels32A during radiation irradiation start detection are accumulated in the capacitors92B of the charge amplifiers92through the signal lines36. The charge amplifiers92generate electric signals with a signal level corresponding to the charge amount accumulated in the capacitors92B. These electric signals are supplied to sample-and-hold circuits93. The electric signals output from the charge amplifiers92are reset every fixed cycle by the reset switches92C being placed in an ON state in response to a control signal supplied from the cassette controller58. The output terminal of each of the charge amplifiers92is connected to the input terminal of the sample-and-hold circuits93.

The sample-and-hold circuits93hold the signal level of the electric signals supplied from the charge amplifiers92in response to a control signal supplied from the cassette controller58. Namely, the sample-and-hold circuits93perform sampling of the signal levels of the electric signals output from the charge amplifiers92at a specific sampling cycle. The output terminals of each of the sample-and-hold circuits93are connected to a common multiplexer94.

The multiplexer94sequentially selects and outputs the signal levels held in the sample-and-hold circuits93according to a control signal supplied from the cassette controller58. Namely, the multiplexer94converts electrical signals from the sample-and-hold circuits93into serial data and supplies the serial data to an A/D (analogue-digital) converter95.

The A/D converter95converts the signal levels of the electrical signals sequentially supplied from the multiplexer94into digital signals, and temporarily stores digital values obtained thereby as radiation amount data (detection signals) in the memory58B of the cassette controller58.

During radiation irradiation standby the cassette controller58executes processing in the signal processor54to read charges accumulated in the radiation detection pixels32A, and accesses the memory58B to acquire radiation amount data (a detection signal) that indicates the amount of charge read from the radiation detection pixels32A, namely the dose of radiation irradiated. The cassette controller58for example sums the radiation amount data (a detection signal) for each of the signal lines36sequentially output from the A/D converter95, and determines that emission of radiation from the radiation source121has started when determined that the summed value thereby obtained is a predetermined specific threshold value for detecting radiation irradiation start or greater.

The electronic cassette40according to the present exemplary embodiment also, in preparation for imaging a radiographic image, has a power adjustment function that adjusts the power supply amount from the battery configuring the power source unit70by adjusting the detection sensitivity during radiation irradiation start detection based on the imaging subject data and the exposure conditions notified from the console110. The power adjustment function according to the present exemplary embodiment is explained below.

The cassette controller58, similarly to in the first exemplary embodiment, is equipped with a first reference table500such as that illustrated inFIG. 10inside its own storage unit58C. The cassette controller58derives the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time by searching the first reference table500using as keys the imaging subject data and the exposure conditions acquired from the console110. A to D inFIG. 10indicate imaging target sites, a1to a4, b1to b4, c1to c4and d1to d4indicate radiation intensity determined by settable tube currents and tube voltages for each of the imaging target sites. Note that the cassette controller58may also compute the intensity of radiation attenuated by passing through the imaging subject based on the imaging subject data and the exposure conditions, and then compute the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time from the computed radiation intensity. The first reference table500becomes redundant in such cases, however processing time is required to compute the radiation irradiation amount.

The cassette controller58derives the number of times for driving the charge amplifiers92of the signal processor54based on the derived radiation irradiation amount. The cassette controller58is equipped with a third reference table502such as that illustrated inFIG. 19stored in its own storage unit58C. The third reference table502includes the irradiation amounts X1to X8of radiation that will be irradiated onto the electronic cassette40within the specific period of time associated with numbers N1to N8of the charge amplifiers92that should be driven. The cassette controller58derives the number of the charge amplifiers92that should be driven (drive number) by searching the third reference table502using as keys the derived radiation irradiation amounts.

As explained above, the cassette controller58sums the digital values corresponding to each of the signal lines36sequentially output from the A/D converter95, and detects radiation irradiation start based on the summed value thereby obtained. During imaging of a radiographic image, in cases in which the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time is relatively small, the detection sensitivity to radiation irradiation start needs to be raised by making the charge amplifier92drive number relatively large, thereby increasing the summed value obtained above. However, when the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time is relatively large, a sufficiently large summed value can be obtained even with a reduced charge amplifier92drive number and lower detection sensitivity to radiation irradiation start. In the electronic cassette40according to the present exemplary embodiment the detection sensitivity to radiation irradiation start is adjusted by controlling the charge amplifier92drive number.

In the third reference table502irradiation amounts and charge amplifier92drive numbers are associated with each other such that the charge amplifier92drive number gets smaller the larger the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time. Namely, during radiation irradiation start detection, the cassette controller58identifies the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time, and controls the switches92E of the signal processor54such that the charge amplifier drive number is smaller the larger the identified irradiation amount. It is accordingly possible to achieve a reduction in power consumption in the irradiation standby state compared to cases in which all the charge amplifiers92are always driven and a high detection sensitivity is always maintained irrespective of the radiation irradiation amount.

Note that it is possible to appropriately increase or decrease the number of drive number adjustment steps by re-writing the third reference table502. Moreover, configuration may be made such that a relationship equation between the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time and the charge amplifier92drive number is pre-stored in the storage unit58C, and a cassette controller38computes the charge amplifier92drive number from the relationship equation.

Moreover, configuration may be made such that instead of the charge amplifier92drive number, or as well as the drive number, the cassette controller58derives which of the charge amplifiers92connected to which of the signal lines36to drive. In such cases, determination is made as to which of the radiation detection pixels32A disposed at which position on the TFT substrate30to make active. The charge amplifiers92to be driven are preferably derived by the cassette controller58such that the active radiation detection pixels32A are not unevenly distributed on the TFT substrate30. In other words, preferably the charge amplifiers92to be driven are derived such that the active radiation detection pixels32A are uniformly distributed on the TFT substrate30. For example, the cassette controller58may be set to control such that the charge amplifiers92to be driven are disposed on every other line, or on one line every plural number of lines.

Explanation follows regarding a detection sensitivity setting processing program according to the present exemplary embodiment executed at step402of the cassette imaging processing program illustrated inFIG. 14, with reference toFIG. 20.FIG. 20is a flow chart illustrating a flow of processing of the detection sensitivity setting processing program according to the present exemplary embodiment executed by the CPU58A of the cassette controller58, and corresponds to the flow chart according to the detection sensitivity setting processing program according to the first exemplary embodiment illustrated inFIG. 15.

At step460ofFIG. 20, the CPU58A derives the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time by searching the first reference table500(seeFIG. 10) stored in the storage unit58C using as keys the imaging target site as the imaging subject data and the tube current and tube voltage as the exposure conditions contained in the initial information supplied from the console110.

In the next step462, the CPU58A derives the charge amplifier92drive number corresponding to this irradiation amount by searching the third reference table502(seeFIG. 19) stored in the storage unit58C using as a key the radiation irradiation amount derived at step460. Note that the third reference table502is configured such that drive numbers with smaller values are derived the larger the radiation irradiation amount derived at step460.

At the next step464, the CPU58A selectively switches the switches92E provided between the power input terminal of the charge amplifiers92and the power line to the ON state to achieve the drive number derived at step462. The present detection sensitivity setting processing program is then ended. In the signal processor54, only the switches92E of the charge amplifiers92selected by the cassette controller58are placed in the ON state, and power is only supplied to the drive number of the charge amplifiers92derived at step462.

Note that transition is made to the radiographic image imaging operation after radiation irradiation start has been detected and charges generated according to irradiated radiation start to accumulate in the radiographic imaging pixels32B. When this is performed, the CPU58A may be configured to set the charge amplifier92drive number to a different drive number to that derived in step462. For example, the CPU58A may control the switches92E required to drive all the charge amplifiers92. Namely, the radiation detection sensitivity during imaging a radiographic image may be set higher than the radiation detection sensitivity during radiation irradiation start detection. The CPU58A thereby controls the detection sensitivity to radiation during imaging the radiographic image independently to the radiation detection sensitivity during radiation irradiation start detection.

As explained above, the electronic cassette40according to an exemplary embodiment of the present invention derives the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time prior to radiation actually being irradiated from the radiation source121based on the imaging subject data and exposure conditions supplied from the console110. The electronic cassette40then derives a value for the charge amplifier92drive number based on the derived radiation irradiation amount, thereby setting the detection sensitivity to radiation irradiation start. When doing so, the electronic cassette40makes the power supply amount from the power source unit70smaller the larger the derived radiation irradiation amount, thereby lowering the detection sensitivity to radiation irradiation start. Namely, the electronic cassette40is controlled such that the charge amplifier92drive number is smaller the larger the derived radiation irradiation amount. The cassette controller58accordingly changes the charge amplifier92drive number to appropriately adjust the detection sensitivity according to the anticipated radiation irradiation amount. It is accordingly possible to prevent use of more of the charge amplifiers92than are needed to detect the radiation irradiation start. A reduction in power consumption in the irradiation standby state is accordingly possible in comparison to cases in which all the charge amplifiers92are driven irrespective of the radiation irradiation amount. The electronic cassette40appropriately adjusts the detection sensitivity to radiation according to the anticipated radiation irradiation amount, and so false detection of radiation irradiation start can be prevented in comparison to cases in which a constant high sensitivity is set.

Thus in the present exemplary embodiment, the number of radiation detection pixels32A that are active (active number) is adjusted by the charge amplifier92drive number based on the derived radiation irradiation amount to adjust the power supply amount from the power source unit70. Another example of a method for adjusting the number of radiation detection pixels32A that are active (active number) is a method of adjusting the number of sensor portions13to which the bias voltage is applied. Consequently, the cassette controller58may be configured to adjust the individual number of the sensor portions13to which the bias voltage is applied according to the derived radiation irradiation amount, either in place of adjusting the charge amplifier92drive number, or as well as adjusting the charge amplifier92drive number.

The cassette controller58of the electronic cassette40may execute the following control to control the signal processor54to lower the detection sensitivity to radiation irradiation start by making the power supply amount to the signal processor54smaller the larger the derived radiation irradiation amount.

Namely, the cassette controller58derives a first value as a setting value for gain of the charge amplifiers92in cases in which the irradiation amount of radiation that will be irradiated within the specific period of time derived based on the imaging subject data and the exposure conditions is smaller than a specific threshold value. However, the cassette controller58derives a second value smaller than the first value as the setting value for gain of the charge amplifiers92in cases in which the derived radiation irradiation amount is greater than the specific threshold value. Consequently, the electronic cassette40controls the gain of the charge amplifiers92according to the derived radiation irradiation amount, and so setting gain excessively larger than the appropriate gain needed for radiation irradiation start detection can be prevented. It is accordingly possible to achieve a reduction in power consumption in the irradiation standby state compared to cases in which a comparatively high gain is set in a fixed manner irrespective of the radiation irradiation amount. Note that to make the gain of the charge amplifiers92variable, configuration may be made such that plural capacitors with different capacitance values to each other are selectively connected between the input terminals and the output terminals of the operational amplifiers92A. Moreover, in the above explanation an example is given of a case in which gain is adjusted to two levels employing a single threshold value, however a configuration is possible in which the gain can be varied to three or more levels employing two or more threshold values.

Moreover, the cassette controller58of the electronic cassette40may execute the following control. Namely, the cassette controller58executes correction processing to remove noise components from the radiation amount data (a detection signal) generated in the signal processor54in cases in which the radiation irradiation amount derived based on the imaging subject data and the exposure conditions is smaller than a specific threshold value. In such cases, noise data that is employed in the correction processing is acquired. The noise data is, for example, acquired by storing radiation amount data (a detection signal) in a radiation non-irradiation state in the memory58B. Although executing such correction processing raises the detection sensitivity to radiation irradiation start, attempting to store noise data in the memory58B and retain this noise data consumes power. However, the cassette controller58omits the above described correction processing in cases in which the radiation irradiation amount derived based on the imaging subject data and the exposure conditions is larger than the specific threshold value. Although detection sensitivity to radiation irradiation start is thereby lowered compared to cases in which the correction processing is executed, since the processing to acquire noise data and the processing to hold the noise data in the memory58B is not required, it is possible to achieve a saving in the power consumption of the amount that would have been needed for such processing. Due to omitting execution of the correction processing according to the anticipated radiation irradiation amount, the electronic cassette40is accordingly able to achieve a reduction in power consumption in the irradiation standby state compared to cases in which the correction processing is always executed irrespective of the radiation irradiation amount. Note that in the above exemplary embodiment an example is given of a case in which the correction processing is omitted according to the radiation irradiation amount, however there is no limitation thereto and other processing that contributes to raising the detection sensitivity during radiation irradiation start detection may be employed.

Third Exemplary Embodiment

Explanation follows regarding an electronic cassette according to a third exemplary embodiment of the present invention. In the electronic cassette40according to the first exemplary embodiment and the second exemplary embodiment, automatic adjustment of the power supply amount from the power source unit70is performed by deriving the irradiation amount of radiation that will be irradiated onto the electronic cassette40itself within the specific period of time, and deriving the bias voltage to be applied to the sensor portions13, or the charge amplifier drive number (namely the radiation detection sensitivity), according to the derived radiation irradiation amount. In contrast thereto, in the electronic cassette according to the third exemplary embodiment, the irradiation amount of radiation that will be irradiated onto the electronic cassette40itself within the specific period of time is derived, and the derived irradiation amount is notified to a radiographer in a configuration enabling manual adjustment of radiation detection sensitivity.

FIG. 21is a perspective view of an electronic cassette40aaccording to the third exemplary embodiment of the present invention, andFIG. 22is a diagram illustrating configuration of relevant portions of an electrical system of the electronic cassette40a. The electronic cassette40aaccording to the present exemplary embodiment includes a notification unit45and a sensitivity setting input section46on the front face of a housing41.

The notification unit45is a unit for notifying the radiographer of the irradiation amount of radiation that will be irradiated onto the electronic cassette40awithin the specific period of time derived by the cassette controller58based on the imaging subject data and the exposure conditions. The notification unit45is, for example, a display that displays the radiation irradiation amount derived by the cassette controller58. A radiographer is able to ascertain the irradiation amount of radiation that will be irradiated onto the electronic cassette40aprior to emission of radiation from the radiation source121by a display of the notification unit45. Note that the notification unit45may be configured to include a speaker so as to notify the radiation irradiation amount by voice.

The sensitivity setting input section46is for manual adjustment of the magnitude of the bias voltage output from the bias voltage generator71and, for example, includes an adjustment knob configuration. For example, configuration may be made such that the bias voltage output from the bias voltage generator71is made larger by rotating the adjustment knob configuring the sensitivity setting input section46to the right, and the bias voltage is made smaller by rotating to the left. A radiographer is thereby able to determine the bias voltage (radiation detection sensitivity) needed to detect the radiation irradiation start based on the radiation irradiation amount notified by the notification unit45, and able to set the desired bias voltage (radiation detection sensitivity) by operating the sensitivity setting input section46. The power supply amount from the power source unit70is thereby set accompanying setting of the bias voltage. Note that the sensitivity setting input section46may be any configuration capable of adjusting the bias voltage manually, and may include a configuration other than an adjustment knob.

Explanation follows regarding a detection sensitivity setting processing program according to the present exemplary embodiment executed at step402in the cassette imaging processing program illustrated inFIG. 14, with reference toFIG. 23.FIG. 23is a flow chart illustrating a flow of processing of the detection sensitivity setting processing program according to the present exemplary embodiment executed by the CPU58A of the cassette controller58, and corresponds to the flow chart according to the detection sensitivity setting processing program according to the first exemplary embodiment illustrated inFIG. 15.

At step470, the CPU58A derives the irradiation amount of radiation that will be irradiated onto the electronic cassette40awithin the specific period of time by searching the first reference table500(seeFIG. 10) stored in the storage unit58C using as keys the imaging target site as the imaging subject data, and the tube current and the tube voltage as the exposure conditions, that are contained in initial information supplied from the console110.

At the next step472, the CPU58A supplies the radiation irradiation amount derived at step470to the notification unit45. The notification unit45accordingly displays the radiation irradiation amount derived by the CPU58A at step470.

At the next step474the CPU58A awaits input of a setting operation of the bias voltage. On display of the radiation irradiation amount on the notification unit45, the radiographer can decide on the bias voltage needed to detect radiation irradiation start based on the radiation irradiation amount displayed on the notification unit45, and the decided bias voltage can be set by operation of the sensitivity setting input section46. The radiographer performs an operation such as depressing a setting complete button provided to the electronic cassette40a, and then processing transitions to the next step476.

At the next step476, the CPU58A supplies a control signal needed to apply to the sensor portions13the bias voltage set through the sensitivity setting input section46at step474, controls the output voltage of the bias voltage generator71, and then ends the present detection sensitivity setting processing program. The bias voltage generator71, on receipt of this control signal, outputs the bias voltage set by operation of the sensitivity setting input section46, and this bias voltage is applied to each of the sensor portions13.

In this manner, the electronic cassette40aaccording to the present exemplary embodiment derives, based on the imaging subject data and the exposure conditions supplied from the console110, the irradiation amount of radiation that will be irradiated onto the electronic cassette40awithin the specific period of time prior to emission of radiation from the radiation source121actually being performed, and displays the derived radiation irradiation amount on the notification unit45. The radiographer is thereby able to ascertain the irradiation amount of radiation to be irradiated to the electronic cassette40aprior to emission of radiation from the radiation source121actually being performed. The magnitude of the bias voltage applied to the sensor portions13can accordingly be adjusted to the desired magnitude by operating the detection sensitivity setting input section46.

Moreover, according to the electronic cassette40aof the present exemplary embodiment, the radiographer is able to determine the bias voltage (radiation detection sensitivity) required to detect radiation irradiation start from the radiation irradiation amount displayed on the notification unit45, enabling the radiographer to set a given bias voltage (radiation detection sensitivity) decided upon by the radiographer by operating the sensitivity setting input section46. Namely, the electronic cassette40aof the present exemplary embodiment enables sensitivity setting that utilizes the experience and skills of the radiographer.

Note that configuration may be made such that the CPU58A derives a radiation irradiation amount based on the imaging subject data and the exposure conditions and derives recommended setting values of bias voltage according to the derived irradiation amount, and displays these derived recommended setting values on the notification unit45. In order to derive the recommended setting values of bias voltage, for example, the second reference table501of the first exemplary embodiment described above may be employed.

The electronic cassette40amay be configured to be capable of being switched between a manual mode in which the bias voltage is set manually as in the present exemplary embodiment, or an automatic mode performed as in the first exemplary embodiment.

Moreover, in the above exemplary embodiment an example is given of a case in which the detection sensitivity to radiation irradiation start is set in the electronic cassette40by setting the bias voltage, however configuration may be made such that the radiation detection sensitivity is changed according to the power supply amount by another parameter, such as for example by setting the drive number or the gain of the charge amplifiers.

Fourth Exemplary Embodiment

Explanation follows regarding a radiographic imaging system according to a fourth exemplary embodiment of the present invention. Explanation has been given in the third exemplary embodiment of an example in which the notification unit45and the sensitivity setting input section46are provided to the electronic cassette40a, however in the present exemplary embodiment the functions of a notification unit45and a sensitivity setting input section46are provided to a console110.

FIG. 24is a flow chart illustrating a flow of processing of a radiographic imaging processing program executed by the CPU113of the console110according to the present exemplary embodiment. Note that in the present exemplary embodiment a first reference table such as the one illustrated inFIG. 10is pre-installed in the ROM114of the console110.

At step320, the CPU113controls the display driver117such that an initial information input screen such as the example illustrated inFIG. 13is displayed by a display111of the console110, and then awaits input of specific data at the next step322.

The radiographer inputs as initial information imaging subject data, including the name of the patient subject to imaging (the imaging subject), the imaging target site, the posture during imaging, and exposure conditions, through an operation panel112. Then determination is affirmative at step322when the radiographer selects an INPUT COMPLETE button displayed in the vicinity of the bottom edge of the initial information input screen, and processing transitions to step324.

At step324the CPU113searches a first reference table500stored in the ROM114using as keys the imaging target site as the imaging subject data, and the tube current and the tube voltage as the exposure conditions, that are contained in initial information that has been input, and derives an irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time.

In the next step326, the CPU113controls the display driver117such that the radiation irradiation amount derived at step324is displayed on the display111. The radiation irradiation amount derived at step324is thereby displayed on the display111.

In the next step328, the CPU113awaits input of setting values for the bias voltage. The irradiation amount of radiation that will be irradiated onto the electronic cassette is displayed on the display111, and the radiographer decides on the bias voltage the electronic cassette40needs to detect radiation irradiation start based on the irradiation amount displayed on the display111, and is able to input the setting value of the decided bias voltage by operating the operation panel112. Processing transitions to the next step330when the radiographer has performed bias voltage setting value input through the operation panel112.

In the next step330, the CPU113transmits the setting value of the bias voltage input by the radiographer at step328to the electronic cassette through the wireless communication unit119. When the electronic cassette has received the bias voltage setting value from the console110, the CPU58A of the electronic cassette40supplies the bias voltage generator71with a control signal to apply the received setting value for the bias voltage to the sensor portions13, thereby controlling the output voltage of the bias voltage generator71.

The processing from step332onwards is similar to the processing in step306to step318of the flow chart of the first exemplary embodiment (FIG. 12) described above and so explanation thereof is omitted.

Thus according to the radiographic imaging system of the present exemplary embodiment, the processing load to the electronic cassette40can be reduced due to executing the processing to derive the irradiation amount of radiation that will be irradiated onto the electronic cassette40within the specific period of time and the input operation to set the bias voltage on the console110side. As a result, the power consumption of the electronic cassette40can be reduced. Moreover, the notification unit45and the sensitivity setting input section46of the third exemplary embodiment do not need to be provided to the electronic cassette even in cases in which manual setting of the detection sensitivity to radiation irradiation start is performed according to the radiographic imaging system of the present exemplary embodiment. Consequently, the configuration for the electronic cassette40can be simplified in comparison to that of the third exemplary embodiment.

Note that in the above the bias voltage is manually set by the radiographer operating the operation panel112, however configuration may be made such that the CPU113of the console110derives a setting value of the bias voltage based on the irradiation amount of radiation that will be irradiated onto the electronic cassette40derived at step324, and transmits the derived bias voltage setting value to the electronic cassette40.FIG. 25is a flow chart illustrating a flow of processing in the console110in such a case. In the flow chart ofFIG. 25, the processing of step326and step328illustrated inFIG. 24is eliminated, and processing to derive the bias voltage setting value based on the irradiation amount of radiation that will be irradiated onto the electronic cassette40derived at step324is added at step327. To derive the bias voltage setting value, the CPU113may, for example, employ the second reference table501of the first exemplary embodiment. Namely, the CPU113derives a bias voltage with a value that is smaller the larger the radiation irradiation amount derived at step324. Moreover, the setting value of the bias voltage derived by the CPU113may be displayed as a recommended setting value on the display111. Moreover, configuration may be made such that the processing program of the console110can switch between an automatic mode in which the bias voltage setting is derived by the CPU113, and a manual mode in which it is decided by the radiographer.

Moreover, an example has been explained above of a case in which the detection sensitivity to radiation irradiation start of the electronic cassette40is set by performing bias voltage setting, however configuration may be made with other parameters that change the radiation detection sensitivity according to the power supply amount, such as for example by setting the drive number or the gain of the charge amplifiers.

Moreover, in each of the exemplary embodiments examples have been given of cases in which the radiation irradiation amount is derived based on both the imaging subject data and the exposure conditions, however the radiation irradiation amount may be derived based on any one of the imaging subject data or the exposure conditions.

Moreover, in each of the exemplary embodiments explanation has been given of cases in which the radiation detection pixels32A are configured by shorting the sources and drains of the thin film transistors10, however the present invention is not limited thereto and may, for example, be configured with direct read lines connected to the capacitors9, and with processing in the signal processor54to read charges accumulated in the capacitors9from the direct read lines.

In each of the exemplary embodiments described above, explanation has been given of cases in which some of the pixels32provided to the radiation detector20are employed for the radiation detection pixels32A, however the present invention is not limited thereto. For example, the radiation detector20may have a stacked configuration with the radiation detection pixels32A in a separate layer to the pixels32. In such cases, the quality of radiographic images can be raised in comparison to the above exemplary embodiments since there are no missing pixels.

Moreover, in the above exemplary embodiments explanation has been given of cases in which some of the radiographic imaging pixels32B are applied as the radiation detection pixels32A, as shown in the example inFIG. 26A, however the present invention is not limited thereto and the radiation detection pixels32A may be provided in gaps between the radiographic imaging pixels32B, for example as shown in the example inFIG. 26B. In such cases, the sensitivity of the radiographic imaging pixels32B provided at positions corresponding to the radiation detection pixels32A decreases since the surface area of these radiographic imaging pixels32B decreases, however the quality of radiographic images can be increased since these pixels can also be used for radiographic image detection.

The sensors for detecting radiation do not necessarily have to be applied to the pixels of the radiation detector20, and configuration may be made such that radiation irradiation start is detected by designated radiation detection sensors that generate charge on irradiation with radiation, provided for example between each row of pixels in the radiation detector20, or at predetermined positions in peripheral positions. In such cases, such sensors do not necessarily have to be provided to the radiation detector20, and may be disposed as a separate body to the radiation detector20.

In the above exemplary embodiments, explanation has been given of cases in which the radiation detection pixels32A and the radiographic imaging pixels32B are provided separately to one another, however the present invention is not limited thereto. Configuration may be made wherein the radiographic imaging pixels32B are applied as sensors that determine whether or not radiation has been detected, without providing the radiation detection pixels32A. Namely, configuration may be made with the sensors that determine whether or not radiation has been detected being common to the radiographic imaging pixels32B. In such cases, the present invention can be achieved simply, without the need to provide extra sensors.

In the above exemplary embodiments, explanation has been given of cases in which the sensor portions13are configured including an organic photoelectric conversion material that generates charge upon receiving light generated by the scintillator8. The present invention is not limited thereto, and configuration may be made wherein the sensor portions13do not include an organic photoelectric conversion material. For example, the sensor portions13may employ a semiconductor such as amorphous selenium, in a configuration wherein radiation is converted directly into charge.

In the above exemplary embodiments, explanation has been given of cases in which the case42that houses the cassette controller58and the power source unit70is disposed inside the housing41of the electronic cassette40so as not to overlap with the radiation detector20, however there is no limitation thereto. The radiation detector20may for example be disposed so as to overlap with the cassette controller58and/or the power source unit70.

In the above exemplary embodiments, explanation has been given of cases in which wireless communication is performed between the electronic cassette40and the console110, and between the radiation generator120and the console110, however the present invention is not limited thereto, and wired communication may be performed between the electronic cassette40and the console110and/or between the radiation generator120and the console110.

In the above exemplary embodiments, explanation has been given of cases in which X-rays are applied as the radiation, however the present invention is not limited thereto and other radiation such as gamma rays may be applied as the radiation.

Other configurations of the RIS100(seeFIG. 1), the radiographic imaging room180(seeFIG. 2), the electronic cassette40(seeFIG. 3toFIG. 8) and the imaging systems104(seeFIG. 9) described in the above exemplary embodiments are merely examples thereof. Obviously, for example, unnecessary portions may be omitted, new portions added, and connection states changed within a scope not departing from the spirit of the present invention.

Moreover, the flow of processing in each of the programs described in the above exemplary embodiments (seeFIG. 12,FIG. 14,FIG. 15,FIG. 20) are also merely examples thereof, and obviously unnecessary steps may be omitted, new steps added, and processing sequences varied within a scope not departing from the spirit of the present invention.

Each of the controls for adjusting the detection sensitivity in the detection of radiation irradiation start illustrated in each of the above exemplary embodiments may be combined as appropriate. For example, it is possible to execute a combination of control to adjust the bias voltage applied to the sensor portions13illustrated in the first exemplary embodiment and control to adjust the drive number of the charge amplifiers92illustrated in the second exemplary embodiment.