JP2011102702A - Calibration method for radiation image photographing device, the radiation image photographing device, and radiation image photographing system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a calibration method for a radiation image photographing device which can reduce variations in image data caused by variations in components of each A/D conversion circuit, in radiation images. <P>SOLUTION: The calibration method for a radiation image photographing device includes: a radiation irradiation process for irradiating the radiation image photographing device 1 respectively with a predetermined plurality of doses Rhigh, Rlow of radiation on the lower dose side than half the maximum dose of radiation detectable by radiation detection elements 7; a read conversion process for converting image data of a plurality of analog values obtained by reading each charge accumulated in the radiation detection elements 7 by irradiation with the plurality of doses of radiation, into image data Gi of a plurality of digital values for each A/D conversion circuit 20; and a correction value calculation process for, based on the image data Gi of the plurality of digital values converted for each A/D conversion circuit 20, calculating correction values q with respect to the image data of the digital values for each A/D conversion circuit 20. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to a calibration method for a radiographic image capturing apparatus, a radiographic image capturing apparatus, and a radiographic image capturing system, and in particular, a calibration method for a radiographic image capturing apparatus provided with an A / D conversion circuit, and the same. The present invention relates to a radiographic imaging apparatus and a radiographic imaging system.

  A so-called direct type radiographic imaging device that generates electric charges by a detection element in accordance with the dose of irradiated radiation such as X-rays and converts it into an electrical signal, or other radiation such as visible light with a scintillator or the like. Various so-called indirect radiographic imaging devices have been developed that convert charges to electromagnetic waves after being converted into electrical signals by generating electric charges with photoelectric conversion elements such as photodiodes in accordance with the energy of the converted and irradiated electromagnetic waves. Yes. In the present invention, the detection element in the direct type radiographic imaging apparatus and the photoelectric conversion element in the indirect type radiographic imaging apparatus are collectively referred to as a radiation detection element.

  This type of radiographic imaging apparatus is known as an FPD (Flat Panel Detector), and conventionally formed integrally with a support base (or a bucky apparatus) (see, for example, Patent Document 1). A portable radiographic imaging device in which an element or the like is housed in a housing has been developed and put into practical use (see, for example, Patent Documents 2 and 3).

  In such a radiographic imaging device, for example, as shown in FIG. 7 to be described later, usually, a plurality of scanning lines 5 and a plurality of signal lines 6 are arranged on the detection part P of the glass substrate so as to intersect each other. In addition, the radiation detection elements 7 are arranged two-dimensionally (matrix) in each region partitioned by the scanning lines 5 and the signal lines 6. Each radiation detection element 7 is connected to a thin film transistor (hereinafter referred to as TFT) 8 which is a switching means, and is represented as a drain electrode 8d of each TFT 8 (D in FIG. 7). ) Are connected to the signal line 6 respectively.

  Each signal line 6 is connected to each readout circuit 17 including an amplifier circuit 18 and the like formed in the readout IC 16, and the output side of each readout circuit 17 is A / D converted via the multiplexer 21 and the like. Connected to the circuit 20. In many cases, the readout IC 16 is configured such that one A / D conversion circuit 20 is provided for every predetermined number of readout circuits 17 such as 128.

  The charge accumulated in each radiation detection element 7 after radiographic imaging is read out from each radiation detection element 7 to each readout circuit 17 via the signal line 6 and amplified during the readout process. After being converted into analog value image data by charge voltage conversion in the circuit 18 or the like, it is converted into digital value image data (that is, so-called raw data) by the A / D conversion circuit 20 and outputted, and storage means 40.

  Further, in the image processing on the raw data read out in this way, usually, after each of the raw data obtained from each radiation detection element 7 is subjected to correction processing such as offset correction and gain correction. The final image data for diagnosis or the like is created by performing logarithmic conversion processing on the corrected image data.

  In the present specification, image data after the correction processing such as offset correction and gain correction is performed on the raw data read from the radiation detection element 7 and A / D converted by the A / D conversion circuit 20. Are collectively referred to as an output value Gi from the A / D conversion circuit 20 in the sense of image data output from the A / D conversion circuit 20. That is, in the case of the output value Gi from the A / D conversion circuit 20, the output value (image data) Gi has already been subjected to correction processing such as offset correction and gain correction for each radiation detection element 7. It shall be.

  Further, in this specification, a value output by performing logarithmic conversion processing on the output value Gi from the A / D conversion circuit 20 is referred to as an output value Go after logarithmic conversion processing.

  By the way, in the A / D conversion circuit provided for each predetermined number of readout circuits 17 as described above, variations occur in each component constituting the A / D conversion circuit during the manufacturing process of the readout IC. . Therefore, in each A / D conversion circuit, the linearity of the output characteristic of the output value Gi from the A / D conversion circuit with respect to the value of the analog value image data Ga read by the reading circuit and input to the A / D conversion circuit. There is a problem that (linearity) varies.

  That is, for example, as illustrated in FIG. 18, when the analog value image data Ga input to the A / D conversion circuit is changed, the input image data Ga for each of the A / D conversion circuits 20 a to 20 c. There is a problem that the slope of the graph representing the linearity of the output value Gi from the A / D conversion circuit varies.

  Therefore, for example, as described in Patent Document 4 and Patent Document 5, calibration is usually performed for each A / D conversion circuit. Note that in Patent Document 4, output characteristics Gi of an output value Gi from any one A / D conversion circuit out of a plurality of A / D conversion circuits is output to an output value Gi from another A / D conversion circuit. A technique for correcting to match the characteristics is described. Patent Document 5 describes a technique for correcting output characteristics of output values Gi from a plurality of A / D conversion circuits according to reference output characteristics.

JP-A-9-73144 JP 2006-058124 A JP-A-6-342099 Japanese Patent Laid-Open No. 2005-210396 Japanese Patent Laying-Open No. 2005-210480

  However, according to the study by the inventors of the present application, even if the conventional calibration as described above is performed on the radiographic imaging apparatus, the final radiation for diagnosis or the like is based on the output value Go after the logarithmic conversion process. When an image is created, a radiation image is generated particularly in a pixel region where the output value Go after logarithmic conversion processing is small, that is, in a pixel region where the output value Gi from each A / D conversion circuit before logarithmic conversion is originally small. It was found that a striped pattern as shown in FIG.

  That is, the entire surface of a conventional radiographic imaging apparatus that has been calibrated is uniformly irradiated with a low dose of radiation, thereby reading out the charge accumulated in each radiation detection element and processing such as charge-voltage conversion Then, the output value Gi output from the A / D conversion circuit is logarithmically converted, and the output value Go after logarithmic conversion processing is displayed on the radiation image.

  Then, as shown in FIG. 19, in the pixel region Rn having a width of, for example, 128 pixels that is A / D converted and logarithmically converted by the same A / D converter circuit, the output value Go after logarithmic conversion processing of each pixel is The pixel regions Rm and Rn (m) in which the output values Go after the logarithmic conversion processing of each pixel are calculated by A / D conversion and logarithmic conversion by different A / D conversion circuits while the values are almost the same. ≠ n), it has been found that the output values Go after the logarithmic conversion processing of the pixels are different from each other, that is, a significant error occurs in the output value Go. In FIG. 19, the error of the output value Go of each pixel region is expressed so as to be emphasized.

  When the final radiographic image is used, for example, as a diagnosis for diagnosing the condition of a lesioned part of a patient, the lesioned part is lowered due to absorption or scattering when the radiation passes through the patient as a subject. In many cases, the image is taken in a pixel region irradiated with a low dose of radiation. Therefore, if a stripe pattern as described above appears in the pixel region irradiated with such a low dose of radiation, the doctor may misunderstand the state of the lesion, leading to misdiagnosis.

  There is also a highly accurate A / D conversion circuit that can suppress variations in the output value Gi to be output to a small value without depending on the output value Gi. In general, it is expensive and a large number of A / D conversion circuits are used in the radiographic imaging apparatus, which increases the manufacturing cost of the radiographic imaging apparatus. Therefore, even when an A / D conversion circuit with lower accuracy and lower cost is used, development of a calibration method for a radiographic imaging apparatus that can prevent the above-described phenomenon from occurring is required.

  The present invention has been made in view of the above-described problems, and the variation in image data (the output value Go after logarithmic conversion processing) generated due to the component variation in each A / D conversion circuit in a radiographic image is obtained. An object of the present invention is to provide a radiographic imaging apparatus calibration method, a radiographic imaging apparatus, and a radiographic imaging system that can be reduced.

In order to solve the above problem, the calibration method of the radiographic image capturing apparatus of the present invention,
A plurality of scanning lines and a plurality of signal lines arranged so as to intersect with each other; a plurality of radiation detecting elements arranged in a two-dimensional manner in each region partitioned by the plurality of scanning lines and the plurality of signal lines; A detector comprising:
A plurality of readout circuits for reading out charges from the radiation detection elements through the signal lines and outputting the charges as image data for each of the radiation detection elements;
With
Furthermore, in the calibration method of a radiographic imaging apparatus including a plurality of A / D conversion circuits that convert analog value image data output from the plurality of readout circuits into digital value image data, respectively,
A radiation irradiation step of irradiating the radiation imaging apparatus with a plurality of predetermined doses of radiation on a lower dose side than a half of the maximum dose of radiation that can be detected by the radiation detection element;
Each charge accumulated in the radiation detection element by irradiation with the plurality of doses of radiation is read as a plurality of analog value image data, and the read image data of the plurality of analog values is the A / D conversion circuit. A read conversion step of converting the image data into a plurality of digital values for each,
A correction value calculating step for calculating a correction value for the digital value image data for each of the A / D conversion circuits, based on the plurality of digital value image data converted for each of the A / D conversion circuits;
It is characterized by having.

Moreover, the radiographic imaging device of the present invention is
A plurality of scanning lines and a plurality of signal lines arranged so as to intersect with each other; a plurality of radiation detecting elements arranged in a two-dimensional manner in each region partitioned by the plurality of scanning lines and the plurality of signal lines; A detector comprising:
A plurality of readout circuits for reading out charges from the radiation detection elements through the signal lines and outputting the charges as image data for each of the radiation detection elements;
With
A plurality of A / D conversion circuits that convert analog value image data output from the plurality of readout circuits into digital value image data, respectively,
Furthermore, based on the digital value image data converted for each A / D conversion circuit, a correction value calculation means for calculating a correction value for the digital value image data for each A / D conversion circuit,
The correction value calculating means is configured to irradiate a plurality of predetermined doses of radiation on the dose side lower than half of the maximum dose of radiation that can be detected by the radiation detection element during calibration. When the accumulated charges are read out as a plurality of analog value image data and converted into a plurality of digital value image data for each of the A / D conversion circuits, each of the A / D conversion circuits A correction value for the digital image data for each A / D conversion circuit is calculated based on the converted digital image data.

Moreover, the radiographic imaging system of the present invention is
A plurality of scanning lines and a plurality of signal lines arranged so as to intersect with each other; a plurality of radiation detecting elements arranged in a two-dimensional manner in each region partitioned by the plurality of scanning lines and the plurality of signal lines; A detector comprising:
A plurality of readout circuits for reading out charges from the radiation detection elements through the signal lines and outputting the charges as image data for each of the radiation detection elements;
Transfer means for transferring the image data,
Furthermore, a radiographic imaging apparatus comprising a plurality of A / D conversion circuits for converting analog value image data output from the plurality of readout circuits into digital value image data, and
A console that performs logarithmic conversion processing on the image data transferred from the radiation image capturing apparatus to create final image data;
With
When the radiation image capturing apparatus is irradiated with a plurality of predetermined doses of radiation on a dose side lower than half of the maximum radiation dose that can be detected by the radiation detection element during calibration, the radiation detection element Each of the charges stored therein is read out as a plurality of analog value image data, converted into a plurality of digital value image data for each A / D conversion circuit, and transferred to the A / D conversion circuit Transferring image data of the plurality of digital values for each to the console;
When the console receives the image data of the plurality of digital values for each of the A / D conversion circuits from the radiographic imaging device, the image data of the digital values for each of the A / D conversion circuits based on the image data A correction value for is calculated.

  According to the method of calibrating a radiographic imaging apparatus, the radiographic imaging apparatus, and the radiographic imaging system according to the present invention, the maximum radiation dose that can be detected by the radiation detection element during calibration of the radiographic imaging apparatus. A plurality of predetermined doses of radiation on the dose side lower than half the dose of each of the charges and the respective charges accumulated in the radiation detection element are read out as a plurality of analog value image data, respectively, and an A / D conversion circuit When converted into image data having a plurality of digital values for each A / D conversion circuit, a correction value for the image data having a digital value for each A / D conversion circuit is obtained based on the image data having a plurality of digital values for each A / D conversion circuit. Calculated.

  With this configuration, even when a low-precision A / D conversion circuit is used, not only in a range where the output value Gi from the A / D conversion circuit is large, but also in a range where the output value Gi is small, The variation degree of the error of the output value Gi from each A / D conversion circuit is less than an allowable error. Further, even if the variation degree of the error ΔGi exceeds the allowable error, it is possible to perform calibration for each A / D conversion circuit so that the method of exceeding the tolerance ΔGi is significantly reduced as compared with the conventional calibration method. Become.

  In actual radiographic imaging, a logarithmic conversion process is performed on the output value Gi from each A / D conversion circuit to obtain a final output value Go. The final output value generated based on the logarithmic conversion process is obtained. In a radiological image, it is possible to reduce variations in image data (output value Go after logarithmic conversion processing) caused by component variations in each A / D conversion circuit.

  Further, in the radiographic image, a pixel region having a particularly small output value Gi from the A / D conversion circuit, that is, a value of the output value Go after logarithmic conversion processing obtained by logarithmically converting the output value Gi is small. In the pixel region, a striped pattern as shown in FIG. 19 is likely to appear on the radiographic image. However, the method of calibrating the radiographic image capturing device of the present invention, the radiographic image capturing device, and the radiographic image capturing system Accordingly, it is possible to accurately prevent the occurrence of such a striped pattern on the radiographic image.

  Then, since a striped pattern is prevented from being generated in a pixel region having a small output value Go in which a lesioned part of a patient on the radiographic image is captured in this way, the doctor can diagnose the lesioned part using the radiographic image. It is possible to accurately prevent a doctor's misdiagnosis due to the occurrence of a striped pattern.

It is a perspective view which shows the radiographic imaging apparatus common to each embodiment. It is sectional drawing which follows the AA line in FIG. It is a top view which shows the structure of the board | substrate of a radiographic imaging apparatus. FIG. 4 is an enlarged view showing a configuration of an imaging element, a thin film transistor, and the like formed in a small region on the substrate of FIG. 3. It is sectional drawing which follows the XX line in FIG. It is a side view explaining the board | substrate with which COF, a PCB board | substrate, etc. were attached. It is a figure showing the equivalent circuit schematic of the radiographic imaging apparatus which concerns on this embodiment. It is a block diagram showing the equivalent circuit about 1 pixel which comprises a detection part. It is a graph showing the relationship between the output value before logarithmic conversion processing and the output value after logarithmic conversion processing. It is a graph explaining the limit value of the error permitted in the output value after logarithmic conversion processing. 6 is a graph showing an allowable error allowed for an output value before logarithmic conversion processing. It is a graph showing the variation degree of the error of the output value from the A / D conversion circuit by the conventional calibration method. FIG. 13 is an enlarged view of a portion of Gi = 0 to 8192 [LSB] in FIG. 12. It is a graph showing the variation degree of the error of the output value from the A / D conversion circuit by the calibration method of this invention. It is an enlarged view of the part of Gi = 0-8192 [LSB] in FIG. It is a flowchart showing the procedure of the calibration method of a radiographic imaging apparatus. It is a figure which shows the whole structure of the radiographic imaging system which concerns on 2nd Embodiment. It is a graph which shows the example of the output value Gi from the A / D conversion circuit with respect to the image data Ga of the analog value input. It is a figure showing the striped pattern which appears in the pixel area | region where the output value on a radiographic image is small.

  DESCRIPTION OF THE PREFERRED EMBODIMENTS Embodiments of a radiation image capturing apparatus calibration method, a radiation image capturing apparatus, and a radiation image capturing system according to the present invention will be described below with reference to the drawings. However, the present invention is not limited to the following illustrated examples.

  In the following embodiments, a so-called indirect radiation that includes a scintillator or the like as a radiographic imaging device, converts the emitted radiation into electromagnetic waves of other wavelengths such as visible light, and obtains an electrical signal with a radiation detection element. Although an image capturing apparatus will be described, the present invention can also be applied to a so-called direct type radiation image capturing apparatus that directly detects radiation with a radiation detection element without using a scintillator or the like.

  Further, in the following embodiment, the case where the radiographic image capturing apparatus is a portable radiographic image capturing apparatus that can be loaded into a bucky apparatus or used alone is described. The present invention can also be applied to the case of a radiographic imaging device formed integrally with a support base or the like.

[First Embodiment]
In the first embodiment, a case will be described in which the radiographic imaging apparatus itself executes the calibration method for the radiographic imaging apparatus according to the present invention. Hereinafter, first, the configuration of the radiation image capturing apparatus will be described.

[Radiation imaging equipment]
FIG. 1 is an external perspective view of the radiographic image capturing apparatus according to the present embodiment, and FIG. 2 is a cross-sectional view taken along line AA in FIG. As shown in FIGS. 1 and 2, the radiation image capturing apparatus 1 according to the present embodiment is configured by housing a scintillator 3, a substrate 4, and the like in a housing 2.

  The housing 2 is formed of a material such as a carbon plate or plastic that at least the radiation incident surface R transmits radiation. 1 and 2 show a case in which the housing 2 is a so-called lunch box type formed by the frame plate 2A and the back plate 2B. However, the housing 2 is integrally formed in a rectangular tube shape. It is also possible to use a so-called monocoque type.

  As shown in FIG. 1, the side surface of the housing 2 is opened and closed for replacement of a power switch 36, an indicator 37 composed of LEDs and the like, and a battery 41 (not shown) (see FIG. 7 described later). A possible lid member 38 and the like are arranged. In the present embodiment, an antenna device 39 that is a transfer means for transferring image data or the like wirelessly to an external device is embedded in the side surface of the lid member 38.

  It is also possible to transfer the image data or the like to an external device in a wired manner. In this case, for example, as a transfer means, a connection terminal or the like for connecting by inserting a cable or the like is used as radiation. It is provided on the side surface of the image capturing apparatus 1 or the like. Further, the antenna device 39 is not necessarily provided in a state of being embedded in the lid member 38, but can be provided in a necessary number at an appropriate position of the housing 2.

  As shown in FIG. 2, a base 31 is disposed inside the housing 2 via a thin lead plate or the like (not shown) on the lower side of the substrate 4. The disposed PCB substrate 33, the buffer member 34, and the like are attached. In the present embodiment, a glass substrate 35 for protecting the substrate 4 and the radiation incident surface R of the scintillator 3 is disposed.

  The scintillator 3 is attached to a detection unit P, which will be described later, of the substrate 4. As the scintillator 3, for example, a scintillator 3 that has a phosphor as a main component and converts it into an electromagnetic wave having a wavelength of 300 to 800 nm, that is, an electromagnetic wave centered on visible light when it receives incident radiation, is used.

  In the present embodiment, the substrate 4 is formed of a glass substrate. As shown in FIG. 3, a plurality of scanning lines 5 and a plurality of signal lines are provided on a surface 4 a of the substrate 4 facing the scintillator 3. 6 are arranged so as to cross each other. In each small region r defined by the plurality of scanning lines 5 and the plurality of signal lines 6 on the surface 4 a of the substrate 4, radiation detection elements 7 are respectively provided.

  Thus, the entire region r in which a plurality of radiation detection elements 7 arranged in a two-dimensional manner are provided in each small region r partitioned by the scanning line 5 and the signal line 6, that is, shown by a one-dot chain line in FIG. The region is a detection unit P.

  In the present embodiment, a photodiode is used as the radiation detection element 7, but other than this, for example, a phototransistor or the like can also be used. Each radiation detection element 7 is connected to the source electrode 8s of the TFT 8 serving as a switch means, as shown in the enlarged views of FIGS. The drain electrode 8 d of the TFT 8 is connected to the signal line 6.

  The TFT 8 is turned on when a turn-on voltage is applied to the connected scanning line 5 by the scanning drive means 15 described later and applied to the gate electrode 8g, and is generated and accumulated in the radiation detection element 7. The charged electric charge is discharged to the signal line 6. Further, the TFT 8 is turned off when the off voltage is applied to the connected scanning line 5 and the off voltage is applied to the gate electrode 8g, and the emission of the charge from the radiation detecting element 7 to the signal line 6 is stopped. The charges generated in the radiation detection element 7 are held and accumulated in the radiation detection element 7.

  Here, the structure of the radiation detection element 7 and the TFT 8 in this embodiment will be briefly described with reference to the cross-sectional view shown in FIG. FIG. 5 is a sectional view taken along line XX in FIG.

A gate electrode 8g of a TFT 8 made of Al, Cr or the like is formed on the surface 4a of the substrate 4 so as to be integrally laminated with the scanning line 5, and silicon nitride (laminated on the gate electrode 8g and the surface 4a). The first electrode 74 of the radiation detecting element 7 is connected to the upper portion of the gate electrode 8g on the gate insulating layer 81 made of SiN _x ) or the like via the semiconductor layer 82 made of hydrogenated amorphous silicon (a-Si) or the like. The formed source electrode 8s and the drain electrode 8d formed integrally with the signal line 6 are laminated.

The source electrode 8s and the drain electrode 8d are divided by a first passivation layer 83 made of silicon nitride (SiN _x ) or the like, and the first passivation layer 83 covers both the electrodes 8s and 8d from above. In addition, ohmic contact layers 84a and 84b formed in an n-type by doping hydrogenated amorphous silicon with a group VI element are stacked between the semiconductor layer 82 and the source electrode 8s and the drain electrode 8d, respectively. The TFT 8 is formed as described above.

  In the radiation detecting element 7, an auxiliary electrode 72 is formed by laminating Al, Cr, or the like on the insulating layer 71 formed integrally with the gate insulating layer 81 on the surface 4 a of the substrate 4. A first electrode 74 made of Al, Cr, Mo or the like is laminated on the auxiliary electrode 72 with an insulating layer 73 formed integrally with the first passivation layer 83 interposed therebetween. The first electrode 74 is connected to the source electrode 8 s of the TFT 8 through the hole H formed in the first passivation layer 83.

  On the first electrode 74, an n layer 75 formed in an n-type by doping a hydrogenated amorphous silicon with a group VI element, an i layer 76 which is a conversion layer formed of hydrogenated amorphous silicon, and a hydrogenated amorphous A p layer 77 formed by doping a group III element into silicon and forming a p-type layer is formed by laminating sequentially from below.

  When radiation enters from the radiation incident surface R of the housing 2 of the radiographic imaging apparatus 1 and is converted into an electromagnetic wave such as visible light by the scintillator 3, and the converted electromagnetic wave is irradiated from above in the figure, the electromagnetic wave is detected by radiation. The electron hole pair is generated in the i layer 76 by reaching the i layer 76 of the element 7. In this way, the radiation detection element 7 converts the electromagnetic waves irradiated from the scintillator 3 into electric charges.

  On the p layer 77, a second electrode 78 made of a transparent electrode such as ITO is laminated and formed so that the irradiated electromagnetic wave reaches the i layer 76 and the like. In the present embodiment, the radiation detection element 7 is formed as described above. The order of stacking the p layer 77, the i layer 76, and the n layer 75 may be reversed. Further, in the present embodiment, a case where a so-called pin-type radiation detection element formed by sequentially stacking the p layer 77, the i layer 76, and the n layer 75 as described above is used as the radiation detection element 7. However, it is not limited to this.

A bias line 9 for applying a bias voltage to the radiation detection element 7 is connected to the upper surface of the second electrode 78 of the radiation detection element 7 via the second electrode 78. The second electrode 78 and the bias line 9 of the radiation detection element 7, the first electrode 74 extended to the TFT 8 side, the first passivation layer 83 of the TFT 8, that is, the upper surfaces of the radiation detection element 7 and the TFT 8 are A second passivation layer 79 made of silicon nitride (SiN _x ) or the like is covered from above.

  As shown in FIGS. 3 and 4, in this embodiment, one bias line 9 is connected to a plurality of radiation detection elements 7 arranged in rows, and each bias line 9 is connected to a signal line 6. Are arranged in parallel with each other. Further, each bias line 9 is bound to the connection 10 at a position outside the detection portion P of the substrate 4.

  In this embodiment, as shown in FIG. 3, each scanning line 5, each signal line 6, and connection 10 of the bias line 9 are input / output terminals (also referred to as pads) provided near the edge of the substrate 4. 11 is connected. As shown in FIG. 6, a COF (Chip On Film) 12 in which a chip such as an IC 12a is incorporated in each input / output terminal 11 is an anisotropic conductive adhesive film (Anisotropic Conductive Film) or anisotropic conductive paste (Anisotropic paste). It is connected via an anisotropic conductive adhesive material 13 such as Conductive Paste).

  The COF 12 is routed to the back surface 4b side of the substrate 4 and connected to the PCB substrate 33 described above on the back surface 4b side. Thus, the board | substrate 4 part of the radiographic imaging apparatus 1 is formed. In FIG. 6, illustration of the electronic component 32 and the like is omitted.

  Here, the circuit configuration of the radiation image capturing apparatus 1 will be described. FIG. 7 is a block diagram showing an equivalent circuit of the radiographic imaging apparatus 1 according to the present embodiment, and FIG. 8 is a block diagram showing an equivalent circuit for one pixel constituting the detection unit P.

  As described above, each radiation detection element 7 of the detection unit P of the substrate 4 has the bias line 9 connected to the second electrode 78, and each bias line 9 is bound to the connection 10 to the bias power supply 14. It is connected. The bias power supply 14 applies a bias voltage to the second electrode 78 of each radiation detection element 7 via the connection 10 and each bias line 9. The bias power source 14 is connected to a control unit 22 described later, and the control unit 22 controls a bias voltage applied to each radiation detection element 7 from the bias power source 14.

  In the present embodiment, the current detection means 43 for detecting the amount of current flowing through the connection 10 (bias line 9) is provided in the connection 10 of the bias line 9. As described above, when the radiation imaging apparatus 1 is irradiated with radiation, electron-hole pairs are generated in the i layer 76 (see FIG. 5) of each radiation detection element 7, and this is the bias line 9 or the connection. The current detection means 43 can detect the start and end of radiation irradiation by detecting the increase or decrease of the current flowing through the connection 10. In the present invention, the current detection means 43 is not necessarily provided.

  As shown in FIGS. 7 and 8, in this embodiment, it can be seen that the bias line 9 is connected via the second electrode 78 to the p-layer 77 side (see FIG. 5) of the radiation detection element 7. In addition, the bias power supply 14 supplies a voltage equal to or lower than a voltage applied to the second electrode 78 of the radiation detection element 7 via the bias line 9 as a bias voltage on the first electrode 74 side of the radiation detection element 7 (that is, a so-called reverse bias voltage). Is applied.

  The first electrode 74 of each radiation detection element 7 is connected to the source electrode 8s of the TFT 8 (indicated as S in FIGS. 7 and 8), and the gate electrode 8g of each TFT 8 (FIGS. 7 and 8). Are respectively connected to the lines L1 to Lx of the scanning line 5 extending from the gate driver 15b of the scanning driving means 15 to be described later. Further, the drain electrode 8 d (denoted as D in FIGS. 7 and 8) of each TFT 8 is connected to each signal line 6.

  In the present embodiment, the scanning drive unit 15 includes a power supply circuit 15a that supplies an on voltage and an off voltage to the gate driver 15b, and a voltage applied to each of the lines L1 to Lx of the scanning line 5 between the on voltage and the off voltage. A gate driver 15b that switches between the on state and the off state of each TFT 8 is provided.

  Each signal line 6 is connected to each readout circuit 17 formed in the readout IC 16. Note that the readout IC 16 is provided with one readout circuit 17 for each signal line 6.

  The readout circuit 17 includes an amplifier circuit 18 and a correlated double sampling circuit 19. In addition to the respective readout circuits 17, an analog multiplexer 21 and an A / D conversion circuit 20 are provided in the readout IC 16. 7 and 8, the correlated double sampling circuit 19 is represented as CDS. In FIG. 8, the analog multiplexer 21 is omitted.

  In the present embodiment, the amplifier circuit 18 is configured by a charge amplifier circuit, and is configured by connecting a capacitor 18b and a charge reset switch 18c in parallel to the operational amplifier 18a and the operational amplifier 18a. In addition, a power supply unit 18 d for supplying power to the amplifier circuit 18 is connected to the amplifier circuit 18.

Further, the signal line 6 is connected to the inverting input terminal on the input side of the operational amplifier 18 a of the amplifier circuit 18, and the reference potential V ₀ is applied to the non-inverting input terminal on the input side of the amplifier circuit 18. ing. Note that the reference potential V ₀ is set to an appropriate value, and in this embodiment, for example, 0 [V] is applied.

  The charge reset switch 18c of the amplifier circuit 18 is connected to the control means 22 described later, and is controlled to be turned on / off by the control means 22. When the image data is read from each radiation detection element 7, when the charge reset switch 18c is turned off and the TFT 8 of the radiation detection element 7 is turned on (that is, the scanning line 5 is applied to the gate electrode 8g of the TFT 8). When an on-voltage for signal readout is applied through the charge detection device 7), the electric charge released from the radiation detection element 7 flows into the capacitor 18 b and is accumulated.

  A voltage value corresponding to the accumulated charge amount is output from the output side of the operational amplifier 18a. In this way, the amplifier circuit 18 outputs a voltage value according to the amount of charge output from each radiation detection element 7 and converts the charge voltage. Note that the amplifier circuit 18 may be configured to output a current in accordance with the charge output from the radiation detection element 7.

  During the process of reading image data from each radiation detection element 7, the charge is read from each radiation detection element 7, and the voltage value that has been subjected to charge-to-voltage conversion by the amplifier circuit 18 is correlated by the correlated double sampling circuit 19. Double sampling processing is performed, and analog value image data is output downstream.

  The image data for each radiation detection element 7 output from the correlated double sampling circuit 19 is transmitted to the analog multiplexer 21 (see FIG. 7), and is sequentially transmitted from the analog multiplexer 21 to the A / D conversion circuit 20. . Then, the A / D conversion circuit 20 sequentially converts the image data into digital values, outputs them to the storage means 40, and sequentially stores them.

  In the present embodiment, in the process of reading the image data from the radiation detection element 7, the lines L1 to Lx of the scanning line 5 to which the ON voltage is applied are sequentially switched, and each of the lines is performed as described above. A process for reading image data from the radiation detection element 7 is performed.

  In the present embodiment, one A / D conversion circuit 20 and one analog multiplexer 21 are provided in the read IC 16, and one A / D conversion circuit 20 (that is, one read IC 16) has 128. The signal lines 6 and 128 readout circuits 17 provided for each signal line 6 are connected.

  The control means 22 includes a CPU (Central Processing Unit), a ROM (Read Only Memory), a RAM (Random Access Memory), an input / output interface connected to the bus, an FPGA (Field Programmable Gate Array), or the like (not shown). It is configured. It may be configured by a dedicated control circuit.

  And the control means 22 controls operation | movement etc. of each member of the radiographic imaging apparatus 1. Further, as shown in FIG. 7 and the like, the control means 22 is connected with a storage means 40 constituted by a DRAM (Dynamic RAM) or the like.

  In the present embodiment, the above-described antenna device 39 is connected to the control unit 22, and each member such as the detection unit P, the scanning drive unit 15, the readout circuit 17, the storage unit 40, the bias power supply 14, and the like. A battery 41 for supplying electric power is connected. The battery 41 is provided with a connection terminal 42 for charging the battery 41 by supplying power to the battery 41 from a charging device (not shown) such as a cradle.

  As described above, the control means 22 controls the bias power supply 14 to set a bias voltage to be applied to each radiation detection element 7 from the bias power supply 14, or the charge reset switch 18 c of the amplification circuit 18 of the readout circuit 17. Various processes such as on / off control and transmission of a pulse signal to the correlated double sampling circuit 19 to control on / off of the sample hold function are executed.

  Further, the control means 22 performs scanning from the scanning driving means 15 to the scanning driving means 15 at the time of reset processing of each radiation detecting element 7 or reading of image data from each radiation detecting element 7 after radiographic imaging. Processing such as transmitting a pulse signal for switching the voltage applied to the gate electrode 8g of each TFT 8 between the on voltage and the off voltage via the line 5 is performed.

  On the other hand, in the present embodiment, the control means 22 uses the A / D conversion circuit based on the digital value image data converted and output for each A / D conversion circuit 20 during calibration of the radiation image capturing apparatus 1. It is configured to function as correction value calculation means for calculating correction values for every 20 digital values of image data.

  Hereinafter, the calibration method of the radiographic imaging apparatus 1 having the above-described configuration and the configuration of processing in the control unit 22 as the correction value calculation unit will be described. The operation of the radiographic image capturing apparatus 1 and the calibration method of the radiographic image capturing apparatus 1 according to the present embodiment will also be described.

  In this embodiment, it is assumed that the offset correction value and the gain correction value for each radiation detection element 7 have already been determined, but the conventional calibration for obtaining the offset correction value and the gain correction value after the present calibration method. Correction may be performed or may be performed in conjunction with this calibration. It is preferable to obtain and correct the offset correction value and the gain correction value by the conventional calibration correction before the main calibration because the accuracy is further improved.

[Calibration method for radiation imaging equipment]
Hereinafter, first, the principle of the calibration method of the radiographic image capturing apparatus 1 according to the present invention, more precisely the principle of the calibration method of each A / D conversion circuit 20 in the radiographic image capturing apparatus 1 will be described.

  As described above, in the radiographic image capturing apparatus 1, each image data of analog values that has been read out for each radiation detection element 7 by the readout circuit 17 and subjected to charge-voltage conversion in the readout process of image data after radiographic imaging has been obtained. Each image is converted into digital image data (ie, so-called raw data) by a predetermined A / D conversion circuit 20 and output.

  Further, each digital image data (raw data) output from the A / D conversion circuit 20 includes an offset correction value, a gain correction value, and the like determined for each radiation detection element 7 as described above. After performing correction processing such as offset correction and gain correction, logarithmic conversion processing is performed on each image data subjected to the correction processing. Then, final image data for diagnosis and the like is formed based on the logarithmically converted image data.

  As described above, a plurality of (for example, 128) signal lines 6 are connected to the A / D conversion circuit 20, and a large number (for example, 1000) of radiation detection elements 7 are connected to each signal line 6 via the TFT 8. Connected. Therefore, when calibrating each A / D conversion circuit 20, an average value of image data of each radiation detection element 7 subjected to correction processing such as offset correction and gain correction is calculated, and the average value is calculated. Calibration is performed.

  Even if the same dose of radiation is irradiated, fluctuations occur in the image data read from each radiation detection element 7, but the fluctuations are canceled out by adopting the average value of the image data as described above. There is an advantage that a value with reduced fluctuation can be used.

  For example, among the 128 signal lines 6 connected to the individual A / D conversion circuits 20, the radiation detection elements 7 connected to a predetermined number of signal lines 6 and the predetermined number of scanning lines 5 are connected. It is also possible to use the image data read from the radiation detection element 7 that has been read.

  Further, as described above, the average value of the image data after the correction processing such as offset correction and gain correction is performed on the raw data is meant as the image data output from the A / D conversion circuit 20. These are collectively referred to as an output value Gi from the A / D conversion circuit 20. Further, a value output by performing logarithmic conversion processing on the output value Gi from the A / D conversion circuit 20 is referred to as an output value Go after logarithmic conversion processing.

Hereinafter, a case where the output value Gi from the A / D conversion circuit 20 is configured to take a value in the range of 0 to 65535 (= 2 ¹⁶ −1) [LSB] will be described. In this case, the image resolution Ri of the output value Gi before logarithmic conversion processing is 65536. Further, when the image resolution of the output value Ro after logarithmic conversion processing is Ro (for example, 4096 (= 2 ¹² )) and the dynamic range is D (for example, 4 digits), the logarithmic conversion processing is expressed by the following equation (1). It is assumed that logarithmic conversion processing is performed in accordance with the expressed logarithmic conversion formula.
Go = Ro + (Ro / D) × log (Gi / Ri) (1)

  In this case, the relationship between the output value Gi [LSB] from the A / D conversion circuit 20 before the logarithmic conversion processing and the output value Go [LSB] after the logarithmic conversion processing is as shown in the graph of FIG. .

  Then, in the radiographic imaging device 1, the error ΔGo of the output value Go after the logarithmic conversion process is not dependent on the output value Gi from the A / D conversion circuit 20 before the logarithmic conversion process, as shown in FIG. The output value Go is required to be within a certain numerical range, that is, within a range of Go−δGo to Go + δGo. Note that δGo is an allowable error representing a limit value of the error ΔGo allowed for the output value Go after the logarithmic conversion process.

  When the predetermined error tolerance δGo as described above is allowed for the output value Go after the logarithmic conversion processing, how much is the output value Gi from the A / D conversion circuit 20 before the logarithmic conversion processing Is calculated on the basis of the above equation (1). That is, when the limit value of the error ΔGi allowed for the output value Gi from the A / D conversion circuit 20 before the logarithmic conversion processing is expressed as the allowable error δGi, how the allowable error δGi is expressed (1 ) Calculated based on the formula.

When the above equation (1) is transformed,
Gi = Ri × 10 ^ (Go × D / Ro-D) (2)
Transformed into Therefore, when Gi ± δGi is substituted for Gi in the equation (2) and Go ± δGo is substituted for Go,
Gi ± δGi = Ri × 10 ^ {(Go ± δGo) × D / Ro-D}
It becomes.

And solving this, the tolerance δGi is
± δGi = Ri × 10 ^ {(Go ± δGo) × D / Ro-D} -Gi
= Ri × 10 ^ {(Go ± δGo) × D / Ro-D}
−Ri × 10 ^ (Go × D / Ro-D) (3)
It is expressed.

  Considering that the output value Go after the logarithmic conversion process is a digital value, the error ΔGo is normally allowed if it is 0.5 [LSB] or less of the quantization error, so that it is the limit value of the error range. The error δGo is preferably 0.5 [LSB], but experience has shown that it is more preferable to set the allowable error δGo to a stricter value of 0.25 [LSB].

Therefore, substituting δGo = 0.25, Gi = Ri = 65536, and Go = 4096 into the above equation (3),
δGi = 36.8 [LSB]
It becomes. When δGo = 0.5 under the same conditions, δGi = 73.6 [LSB]. Based on this, when the allowable error δGi allowed when δGo is 0.25 [LSB] and 0.5 [LSB] is represented in a graph, the allowable error δGi is expressed as in the graph of FIG. The

  By the way, as described above, a highly accurate A / D conversion circuit capable of suppressing the variation (error) of the output value Gi to be output to a small value such as 1 [LSB] without depending on the output value Gi. However, such a high-precision A / D conversion circuit is generally expensive, and it is not practical to use it in the radiographic imaging apparatus 1.

  Therefore, in the radiographic imaging device 1 of the present embodiment, as the A / D conversion circuit 20, for example, a low-precision A / D conversion circuit that is a sequential conversion type and includes a D / A converter and a comparator is used. . In such an A / D conversion circuit, for example, 65536 parts constituting 1 [LSB] are connected in series.

  In the conventional calibration method of the A / D conversion circuit 20, for example, it corresponds to image data (Gi = 65535) read out from the radiation detection element 7 when the maximum dose of radiation that can be detected by the radiation detection element 7 is irradiated. ), The output value Gi (65535) output from the A / D conversion circuit 20 as described above is recorded, and is read out from the radiation detection element 7 when the radiation image capturing apparatus 1 is not irradiated with radiation. Based on the image data (corresponding to Gi = 0), the output value Gi (0) output from the A / D conversion circuit 20 is recorded.

  The A / D conversion circuit 20 is calibrated based on the output values Gi (65535) and Gi (0), and the output value Gi of the A / D conversion circuit 20 at 0 [LSB] and 65535 [LSB]. Adjustment (calibration) is performed so as to suppress variations in the image quality. Note that, instead of using the output value Gi (65535), image data read from the radiation detection element 7 when radiation of half the maximum dose that can be detected by the radiation detection element 7 is irradiated (corresponding to Gi = 32767). In some cases, the output value Gi (32767) output from the A / D conversion circuit 20 based on the above is used.

  Here, it is assumed that the relative accuracy variation σ of the components (for example, resistors) constituting 1 [LSB] of the A / D conversion circuit 20 is 1.25%, and the output value due to the component variations of each A / D conversion circuit 20. In order to evaluate the variation of the Gi error ΔGi (INL error), the accuracy of the low-precision A / D conversion circuit 20 will be considered. At that time, the accuracy of the A / D conversion circuit 20 is evaluated by the variation degree ΔGi (3σ) at 3σ.

The variation degree ΔGi (3σ) at 3σ can be calculated according to the following equation (4).
ΔGi (3σ) = 3 × σ × {(Ri−Gi) × Gi / Ri} ^1/2 (4)

In this case, since σ = 0.0125 and Ri = 65536, the above equation (4) is
ΔGi (3σ) = 3 × 0.0125 × {(65536−Gi) × Gi / 65536} ^1/2 (5)
It becomes.

  When the function ΔGi (3σ) expressed by the equation (5) is illustrated together with the allowable error δGi (0.25) in the case of δGo = 0.25 shown in FIG. 11, it is expressed as a graph of FIG. FIG. 13 is an enlarged view of a portion where Gi = 0 to 8191 [LSB] in FIG.

  As shown in FIG. 12, according to the conventional calibration method of the A / D conversion circuit 20, the variation ΔGi (3σ) of the error ΔGi is almost in the range of the output value Gi from the A / D conversion circuit 20. As shown in FIG. 13, when the output value Gi from the A / D conversion circuit 20 is about 4095 [LSB] or less, the variation degree ΔGi (3σ) is the allowable error δGi. It exceeds (0.25).

  This is because the error ΔGi of the output value Gi out of the plurality of A / D conversion circuits 20 is within the allowable error δGi ( This means that an A / D conversion circuit 20 exceeding 0.25) appears, or an A / D conversion circuit 20 where the error ΔGi of the output value Gi falls below the allowable error δGi (0.25) appears.

  Therefore, in a radiographic image, a pixel region having a small output value Gi from the A / D conversion circuit 20, that is, a pixel having a small output value Go after logarithmic conversion processing obtained by logarithmically converting the output value Gi. In the area, as shown in FIG. 19, a striped pattern appears on the image.

  Therefore, in the calibration method of the radiographic image capturing apparatus 1 according to the present invention, based on the above consideration, even when the output value Gi from the A / D conversion circuit 20 is small, the variation degree ΔGi ( 3σ) is equal to or less than the allowable error δGi (0.25), or at least each A / A is reduced so that the deviation rate ΔGi (3σ) exceeds the allowable error δGi (0.25). Calibration for the D conversion circuit 20 is performed.

  And in order to implement | achieve it, in the calibration method of the radiographic imaging device 1 which concerns on this invention, like the conventional calibration method, the radiation of the maximum dose which the radiation detection element 7 can detect (or the half) Instead of using the output value Gi (65535) (or output value Gi (32767)) based on the image data when the radiation imaging apparatus 1 is irradiated with the radiation (dose radiation), the dose on the lower dose side, that is, radiation Calibration is performed by irradiating the radiation image capturing apparatus 1 with radiation having a dose lower than half the maximum dose that can be detected by the detection element 7.

  In addition, the radiation image capturing apparatus 1 is irradiated with two predetermined doses of radiation on a lower dose side than the half of the maximum dose that can be detected by such a radiation detection element 7, and each A / D conversion circuit 20. Calibration is to be performed.

When the radiation image capturing apparatus 1 is irradiated with the above two doses of radiation, the output values output from each A / D conversion circuit 20 are output values Ch and Cl (h is high and l is low, respectively). Then, as can be understood from the above equation (4), the variation degree ΔGi (3σ) of the error ΔGi in this case can be expressed as the following equation (6).
ΔGi (3σ) = 3 × σ × | {(Ch-Gi) × (Gi-Cl) / (Ch-Cl)} | ^1/2 (6)

  Then, by varying the output values Ch and Cl from the A / D conversion circuit 20, the variation degree ΔGi (3σ) of the error ΔGi can be maintained even in a range where the output value Gi from the A / D conversion circuit 20 is small. The output values Ch and Cl are set so that the allowable error δGi (0.25) is less than or equal to at least the deviation ΔGi (3σ) of the error ΔGi exceeds the allowable error δGi (0.25). Just decide.

  As can be inferred from the graph of the variation degree ΔGi (3σ) of the error ΔGi shown in FIG. 12 and FIG. 13, among the output values Ch and Cl from the A / D conversion circuit 20, the output value on the high output side. It is understood that the way of exceeding the variation degree ΔGi (3σ) with respect to the permissible error δGi (0.25) is reduced as Ch is further reduced from 65535 [LSB] or half that of 32767 [LSB]. This corresponds to shifting the two doses of radiation irradiated to the radiographic imaging apparatus 1 to the lower dose side.

  Therefore, for example, a graph of the function ΔGi (3σ) represented by the above equation (6) when the output value Ch from the A / D conversion circuit 20 is 4800 [LSB] and the output value Cl is 700 [LSB]. FIG. 14 and FIG. 15 which is an enlarged view of a portion where Gi = 0 to 8191 [LSB] is shown. In addition, the broken line in each figure is a graph showing the variation degree ΔGi (3σ) of the error ΔGi represented by the above equation (5) shown in FIGS.

  Actually, in the above equation (6), σ = 0.0125, Gi = 0 [LSB] and ΔGi (3σ) = 1 [LSB] are substituted into the above equation (6), and the variables Ch, Cl Further, Gi = 65535 [LSB] and 36.8 [LSB] which is a value of Gi = 65535 [LSB] of δGi (0.25) are substituted into the above equation (6) to obtain the variable Ch, If another equation of Cl is established and those simultaneous equations are solved, Ch≈4800 [LSB] and Cl≈700 [LSB] are obtained.

  Compared to the conventional calibration method shown in FIGS. 12 and 13, the calibration method of the present invention shown in FIG. 14 and FIG. The degree of variation ΔGi (3σ) of the error ΔGi of the output value Gi is much smaller than the allowable error δGi (0.25), and can be reduced to a slight extent within a small range of the output value Gi.

  Compared to the conventional calibration method, in the calibration method of the present invention, the variation degree ΔGi (3σ) of the error ΔGi on the side where the output value Gi is high (the side close to Gi = 65535 [LSB]) is the allowable error δGi ( However, if the output value Gi is logarithmically converted as shown in FIG. 10, the error ΔGo of the output value Go after logarithmic conversion processing is a certain numerical range defined by the limit value δGo. Because it fits in, there is no problem.

  Further, in the calibration method of the present invention, when the output value Gi is a very small value near 0 [LSB], the variation degree ΔGi (3σ) of the error ΔGi exceeds the allowable error δGi (0.25), and 1 [LSB However, when the output value Gi is in the vicinity of 0 [LSB], the output value Gi output from each A / D conversion circuit 20 is the A / D conversion circuit 20 itself. And noise such as shot noise caused by the amplifier circuit 18 and dark charges.

  Therefore, even when the output value Gi is a value in the vicinity of 0 [LSB], even if the variation degree ΔGi (3σ) of the error ΔGi exceeds the allowable error δGi (0.25) and becomes a value of about 1 [LSB], no matter what. There is no problem.

  As described above, in the calibration method of the radiographic image capturing apparatus 1 according to the present invention, two predetermined doses (that is, for example, lower doses than the maximum dose that can be detected by the radiation detection element 7) (ie, for example, Calibration of each A / D conversion circuit 20 of the radiographic imaging apparatus 1 by irradiating the radiographic imaging apparatus 1 with radiation of the above (dose corresponding to Ch = 4800 [LSB] and Cl = 700 [LSB]). Perform

  With this configuration, the degree of variation ΔGi (3σ) of the error ΔGi of the output value Gi from each A / D conversion circuit 20 is much smaller than the allowable error δGi (0.25), or The variation degree ΔGi (3σ) of the error ΔGi can be made equal to or less than the allowable error δGi (0.25).

  Therefore, in a radiographic image, a pixel region having a small output value Gi from the A / D conversion circuit 20, that is, a pixel having a small output value Go after logarithmic conversion processing obtained by logarithmically converting the output value Gi. Also in the region, it is possible to prevent the striped pattern as shown in FIG. 19 from appearing on the radiation image.

  In the study by the present inventors, in the above formula (6), about 1/8 of the maximum dose (corresponding to 65535 [LSB]) that the radiation detection element 7 can detect the high output value Ch. If the value is set to about 8191 [LSB] or less corresponding to the following dose, and the low output value Cl is set to a value of 0 to 1000 [LSB] corresponding to the weak dose, it is calculated by the above formula (6). The variation degree ΔGi (3σ) of the error ΔGi of the output value Gi from the A / D conversion circuit 20 is less than the allowable error δGi (0.25), or the variation degree ΔGi (3σ) of the error ΔGi is the allowable error δGi ( Even if it exceeds 0.25), it has been found that the degree of increase can be reduced to a very small value as shown in FIG.

  However, the output values Ch and Cl and the specific two doses of radiation to be irradiated to the radiographic imaging apparatus 1 during calibration corresponding to the output values Ch and Cl are specifically determined. Whether the value should be set may vary depending on the type and performance of the A / D conversion circuit 20 used in the radiation image capturing apparatus 1. Therefore, two predetermined doses of radiation to be irradiated to the radiographic imaging apparatus 1 during calibration are set in advance to appropriate doses according to the specifications of the radiographic imaging apparatus 1.

  Next, based on the principle of the calibration method of the radiographic imaging apparatus 1 according to the present invention described above, the control means as the correction value calculation means according to the present embodiment is performed according to the flow of actual calibration for realizing the calibration method. The configuration of the processing at 22 will be described. FIG. 16 is a flowchart showing the procedure of the calibration method of the radiation image capturing apparatus 1.

  At the time of calibration of the radiographic imaging apparatus 1, first, a radiation irradiation step (step S <b> 1 in FIG. 16) is executed, and the maximum radiation that can be detected by the radiation detection element 7 as described above is detected for the radiographic imaging apparatus 1. Two predetermined doses of radiation on the lower dose side than half the dose, that is, two predetermined doses of radiation corresponding to the output values Ch and Cl are respectively irradiated.

  Subsequently, the readout conversion step (step S2) is executed, and the respective charges accumulated in the radiation detecting element 7 by the irradiation of the above-mentioned two doses of radiation are converted into two analog value images by the readout circuit 17, respectively. The data is read as data, and the read image data of two analog values is converted into image data of two digital values for each A / D conversion circuit 20.

  At that time, as described above, the control means 22 performs correction processing such as offset correction and gain correction on the two types of image data converted into digital values, respectively, and all or all connected to the A / D conversion circuit 20. The average value of the image data from the predetermined radiation detection element 7 is calculated, and each output value Gi from the A / D conversion circuit 20 is generated.

  As described above, in the present embodiment, since the low-precision A / D conversion circuit 20 is used, each A / D conversion circuit 20 causes each variation due to variations in components constituting the A / D conversion circuit 20. The output value Gi varies. Therefore, even if the radiation image capturing apparatus 1 is irradiated with two predetermined doses of radiation corresponding to the output values Ch and Cl, the high-dose side from each actual A / D conversion circuit 20 is irradiated. The output value Gi (high) and the output value Gi (low) on the low dose side do not necessarily match the output values Ch and Cl.

  Therefore, the control means 22 serving as the correction value calculating means has two digital value image data converted for each A / D conversion circuit 20, that is, the high dose side from each of the actual A / D conversion circuits 20 described above. The correction value q for the digital image data is calculated for each A / D conversion circuit 20 on the basis of the output value Gi (high) and the output value Gi (low) on the low dose side (step). S3: Correction value calculation step).

  In the present embodiment, in the correction value calculation step (step S3), the control means 22 performs image data of two digital values for each A / D conversion circuit, that is, a high dose with respect to the difference between the two predetermined doses. The change rate of the difference between the output value Gi (high) on the side and the output value Gi (low) on the low dose side is a preset reference change rate, that is, the output value Ch for two predetermined dose differences, The correction value q for the digital image data for each A / D conversion circuit 20 is calculated so as to be equal to the change rate of the difference in Cl.

That is, assuming that two predetermined doses of irradiated radiation are Rhigh and Rlow respectively,
(Ch-Cl) / (Rhigh-Rlow) = q × (Gi (high) -Gi (low)) / (Rhigh-Rlow)
Therefore, the correction value q is eventually set to the preset output values Ch and Cl and the actual high-dose-side output value Gi (high) output from each A / D conversion circuit 20 or low. Based on the output value Gi (low) on the dose side,
q = (Ch-Cl) / (Gi (high) -Gi (low)) (7)
Can be calculated according to

  For each A / D conversion circuit 20, the control means 22 serving as a correction value calculation means outputs the actual high-dose side output value Gi (high) output from each A / D conversion circuit 20 and the low-dose side output. By substituting the value Gi (low) into the above equation (7), the correction value q for the digital image data for each A / D conversion circuit 20 is calculated.

  As described above, according to the radiographic image capturing apparatus 1 and the calibration method for the radiographic image capturing apparatus according to the present embodiment, when the radiographic image capturing apparatus 1 is calibrated, the radiation detecting element is compared with the radiographic image capturing apparatus 1. Two predetermined doses of radiation on the lower dose side than the maximum dose of radiation 7 can be detected are respectively irradiated.

  Then, each charge accumulated in the radiation detection element 7 is read out as two analog value image data, and converted into two digital value image data for each A / D conversion circuit 20. Two kinds of digital value image data converted for each D conversion circuit 20, that is, a correction value q for the digital value image data for each A / D conversion circuit 20 based on each output value Gi from the A / D conversion circuit 20. Is calculated.

  With this configuration, even when the low-precision A / D conversion circuit 20 is used, not only in the range where the output value Gi from the A / D conversion circuit 20 is large, but also in the range where the output value Gi is small. The variation degree ΔGi (3σ) of the error ΔGi of the output value Gi from each A / D conversion circuit 20 is logarithmic because the output value Go after the logarithmic conversion of the output value Gi falls within the predetermined error range δGo. It is possible to perform calibration on each A / D conversion circuit 20 so that the allowable error δGi (0.25) allowed for the output value Gi before the conversion processing is less than or equal to.

  Further, even if the variation degree ΔGi (3σ) of the error ΔGi exceeds the allowable error δGi (0.25), each A / D conversion circuit 20 is so greatly reduced as compared with the conventional calibration method. Can be calibrated.

  In actual radiographic imaging, a logarithmic conversion process is performed on the output value Gi from each A / D conversion circuit 20 to obtain a final output value Go. In a typical radiographic image, it is possible to reduce variations in image data (output value Go after logarithmic conversion processing) caused by component variations in each A / D conversion circuit 20.

  Further, as described above, in such a radiographic image, in particular, a pixel region where the output value Gi from the A / D conversion circuit 20 is small, that is, a logarithmic conversion process obtained by logarithmically converting the output value Gi. In the pixel region where the output value Go is small, a striped pattern as shown in FIG. 19 is likely to appear on the radiographic image. However, the radiographic image capturing device 1 and the radiographic image capturing device according to the present embodiment are calibrated. According to the method, it is possible to accurately prevent the occurrence of such a striped pattern on the radiographic image.

  Then, since a striped pattern is prevented from being generated in a pixel region having a small output value Go in which a lesioned part of a patient on the radiographic image is captured in this way, the doctor can diagnose the lesioned part using the radiographic image. It is possible to accurately prevent a doctor's misdiagnosis due to the occurrence of a striped pattern.

  In the present embodiment, the radiation image capturing apparatus 1 is irradiated with two predetermined doses of radiation on the dose side lower than the half of the maximum dose that can be detected by the radiation detection element 7. Although the case of performing calibration of the / D conversion circuit 20 has been described, the radiation image capturing apparatus 1 has three or more predetermined doses on the dose side lower than the half dose of the maximum dose that can be detected by the radiation detection element 7. It is also possible to configure so that each A / D conversion circuit 20 is calibrated by irradiating with radiation.

  For example, the radiation image capturing apparatus 1 is irradiated with three predetermined doses of radiation on a lower dose side than the half of the maximum dose that can be detected by the radiation detection element 7. When calibration is performed, output values Ch, Cm, and Cl (m represents middle) are set in advance as reference values for the output values Gi from the A / D conversion circuits 20, and the output values Ch, Cm, and Cl are set. Radiation of the corresponding doses Rhigh, Rmiddle, and Rlow is applied to the radiographic image capturing apparatus 1 to calibrate each A / D conversion circuit 20.

At this time, the correction value q is set in two ways, and when the output value Gi from the A / D conversion circuit 20 is in the range of 0 to Cm [LSB], as estimated from the above equation (7),
q0 _-m = (Cm-Cl) / (Gi (middle) -Gi (low)) (8)
The correction value q _0-m calculated in ( ₁₎ is applied.

Further, when the output value Gi from the A / D conversion circuit 20 is in the range of Cm to 65535 [LSB],
q _m−65535 = (Ch−Cm) / (Gi (high) −Gi (middle)) (9)
The correction value q _m-65535 calculated in (1) is applied.

Calibration of each A / D conversion circuit 20 is performed by irradiating the radiation image capturing apparatus 1 with predetermined four or more doses of radiation on a lower dose side than half the maximum dose that can be detected by the radiation detection element 7. In this case, the output values C1, C2, C3,... (C1 <C2 <C3 <...) are set in advance as reference values for the output values Gi from the A / D conversion circuits 20. In the same manner as described above, correction values q _{0 -c 2} , q _{c 2 -c 3} ,... _Are calculated and applied to the output values Gi of each range.

  Note that in the present embodiment and the above-described modification, the reference low-dose-side output value Cl (or output value C1) can be set to 0 [LSB]. Accordingly, when the radiation dose is 0 [millientgen], that is, the radiation imaging apparatus 1 is not irradiated with radiation at a predetermined dose lower than half the maximum dose that can be detected by the radiation detection element 7. Cases are also included.

  Further, a new offset correction value and gain correction value may be recalculated and corrected from the offset correction value, gain correction value, and A / D conversion correction value q. Furthermore, after obtaining the A / D conversion correction value q first without obtaining the offset correction value and the gain correction value, a new image of the same image or maximum dose is obtained, and the offset for each pixel is obtained. It is also possible to obtain a correction value and a gain correction value.

[Second Embodiment]
In the first embodiment, the case where the radiographic image capturing apparatus 1 itself calculates the correction value q has been described. In this case, the correction value q is stored in, for example, the storage unit 40 and held in the radiographic image capturing apparatus 1 itself. Then, when radiographic imaging is performed, correction processing using the correction value q for the output value Gi output from each A / D conversion circuit 20 is automatically performed in the radiographic imaging apparatus 1.

  However, with this configuration, the radiographic image capturing apparatus 1 performs all the correction processing with the correction value q for the calculation of the correction value q at the time of calibration and the output value Gi from each A / D conversion circuit 20 for each radiographic image capturing. If the radiographic image capturing apparatus 1 is a portable radiographic image capturing apparatus with a built-in battery, the battery 41 may be exhausted.

  Therefore, in the second embodiment of the present invention, the calculation of the correction value q at the time of calibration and the correction processing for each radiographic image capturing are not performed by the radiographic image capturing apparatus 1, but are performed by the console which is an external apparatus. An explanation will be given of the radiographic imaging system.

[Radiation imaging system]
FIG. 17 is a diagram illustrating an overall configuration of the radiographic image capturing system according to the present embodiment. As shown in FIG. 17, the radiographic imaging system 50 includes, for example, an imaging room R <b> 1 that performs imaging of a subject (a patient's imaging target region) that is a part of a patient by irradiating radiation, and an operator such as a radiographer. Are arranged in the anterior chamber R2 for performing various operations such as control of radiation applied to the subject, and the outside thereof.

  In the radiographing room R1, a bucky device 51 that can be loaded with the radiographic imaging device 1 described above, a radiation generating device 52 that includes an X-ray tube (not shown) that generates radiation to irradiate a subject, the radiographic imaging device 1 and a console. A base station 54 provided with a wireless antenna 53 that relays these communications when wirelessly communicating with 58 is provided.

  Note that FIG. 17 shows a case where the portable radiographic image capturing apparatus 1 is used by being loaded into the cassette holding unit 51a of the bucky device 51. However, as described above, the radiographic image capturing device 1 is the bucky device 51. Or may be formed integrally with a support base or the like. In addition, as shown in FIG. 17, the radiographic imaging device 1 and the base station 54 can be connected by a cable so that data can be transmitted by wired communication via the cable.

  In the present embodiment, the cradle 55 that reads the cassette ID from the radiographic image capturing apparatus 1 and notifies the console 58 via the base station 54 when the radiographic image capturing apparatus 1 is inserted into the radiographing room R1. Is provided. Note that the cradle 55 is not necessarily provided, and the cradle 55 can be configured to simply charge the radiographic image capturing apparatus 1 or the like.

  The anterior room R2 is provided with an operation console 57 for controlling radiation irradiation, which includes a switch means 56 for instructing the radiation generator 52 to start radiation irradiation and the like.

  Although the configuration of the radiographic image capturing apparatus 1 is as described above, in the present embodiment, the radiographic image capturing apparatus 1 calculates the correction value q at the time of calibration and from each A / D conversion circuit 20 for each radiographic image capture. The correction process with the correction value q for the output value Gi is not performed. The radiation image capturing apparatus 1 converts image data read from each radiation detection element 7 into digital image data by each A / D conversion circuit 20 when predetermined radiation is irradiated during calibration or radiation image capturing. The converted image data is stored in the storage means 40 and transferred to the console 58 via the base station 54 via the antenna device 39 as a transfer means.

  The radiographic image capturing apparatus 1 may be used by being loaded into the bucky device 51 as described above, but it is not loaded into the bucky device 51 and can be used in a so-called state. .

  That is, the radiation image capturing apparatus 1 is disposed on the upper surface side in a single state, for example, on a bed provided in the imaging room R1 or a bucky apparatus 51B for lying position photographing as shown in FIG. (See FIG. 1) The patient's hand, which is the subject, can be placed on the top, or the patient's waist, legs, etc. lying on the bed can be inserted between the bed and the bed. It has become. In this case, for example, radiation image capturing is performed by irradiating the radiation image capturing apparatus 1 with radiation from a portable radiation generating device 52B or the like via a subject.

  In this embodiment, the console 58 that controls the entire radiographic imaging system 50 is provided outside the imaging room R1 and the front room R2. For example, the console 58 is provided in the front room R2 and the like. It is also possible.

  The console 58 is constituted by a computer or the like in which a CPU, a ROM, a RAM, an input / output interface and the like (not shown) are connected to a bus. A predetermined program is stored in the ROM, and the console 58 reads out the necessary program, expands it in the work area of the RAM, executes various processes according to the program, and controls the entire radiographic imaging system 50 as described above. Is supposed to do.

  The console 58 is connected to the base station 54, the console 57, the storage means 59 configured by a hard disk or the like, and the cradle 55 is connected to the console 58 via the base station 54. The radiation generating device 52 and the like are connected via this. The console 58 is provided with a display screen 58a composed of a CRT (Cathode Ray Tube), an LCD (Liquid Crystal Display), or the like, and other input means such as a keyboard and a mouse are connected thereto.

  When the console 58 receives the image data converted into the digital value by each A / D conversion circuit 20 of the radiation image capturing apparatus 1 through the base station 54 as described above, the console 58 stores the image data. 59 is memorized.

  Further, in the present embodiment, at the time of calibration of the radiographic image capturing apparatus 1, as described above, two predetermined predetermined (or preset) doses lower than the half of the maximum dose that can be detected by the radiation detecting element 7 (or When image data of two (or a plurality of) digital values is transferred for each radiation detection element 7 from the radiation image capturing apparatus 1 irradiated with a plurality of doses of radiation, the console 58 displays the image data. Based on the above, the correction value q is calculated for each A / D conversion circuit 20 of the radiographic imaging apparatus 1.

  Therefore, in this embodiment, among the steps performed in accordance with the principle of the calibration method of the radiographic image capturing apparatus 1 shown in FIG. 16, the radiation irradiation step (step S1) is performed in the radiographing room R1, and the read conversion step ( Step S2) is performed in the radiation image capturing apparatus 1, and a correction value calculation step (step S3) is performed on the console 58.

  Hereinafter, at the time of calibration of the radiation image capturing apparatus 1, the radiation image capturing apparatus 1 is irradiated with two predetermined doses of radiation on the lower dose side than the half of the maximum dose that can be detected by the radiation detecting element 7. As described above, the radiation imaging apparatus 1 is irradiated with predetermined three or more doses of radiation on the lower dose side than the half of the maximum dose that can be detected by the radiation detection element 7. It is also possible to configure so that calibration is performed.

  There are two consoles 58 for each radiation detection element 7 corresponding to the high-dose side and the low-dose side irradiated to the radiographic image capturing apparatus 1 from the radiographic image capturing apparatus 1 when the radiographic image capturing apparatus 1 is calibrated. When the image data is transferred, correction processing such as offset correction and gain correction is performed on the image data, and then each image on the high dose side and the low dose side for each A / D conversion circuit 20. The average value of the data is calculated, and the output values Gi (high) and Gi (low) on the high dose side and the low dose side from each A / D conversion circuit 20 are obtained.

  Then, the correction value q for each A / D conversion circuit 20 of the radiographic imaging apparatus 1 is calculated and stored in the storage unit 59 in accordance with the above equation (7). In this manner, the storage unit 59 of the console 58 includes a table in which the identification information (for example, cassette ID) of the radiation image capturing apparatus 1 and the correction value q for each A / D conversion circuit 20 are associated with each radiation. The values are stored for each image capturing apparatus 1 and are updated each time the radiographic image capturing apparatus 1 is calibrated.

  Note that the console 58 transfers digital image data of each radiation detection element 7 from the radiographic imaging device 1 used for radiographic imaging after radiographic imaging during normal radiographic imaging. Then, correction processing such as offset correction and gain correction is performed on the image data.

  And it corresponds to the said radiation detection element 7 with reference to the table memorize | stored in the memory | storage means 59 based on the identification information of the said radiographic imaging apparatus 1, and the information of the A / D conversion circuit 20 to which those radiation detection elements 7 belong. The correction value q to be calculated is determined, and each image data is corrected by multiplying the image data from the radiation detection element 7 by the correction value q.

  The console 58 generates final image data by performing logarithmic conversion processing as shown in FIG. 9 on each corrected image data.

  As described above, according to the radiographic image capturing system 50 and the calibration method for the radiographic image capturing apparatus according to the present embodiment, it is possible to obtain the same effect as that of the first embodiment.

  In actual radiographic imaging, logarithmic conversion processing is performed on the image data from each radiation detection element 7 to obtain final image data. A final radiographic image generated based on the logarithmic conversion processing is obtained. Thus, it is possible to reduce variations in image data caused by component variations in each A / D conversion circuit 20.

  Therefore, it is possible to accurately prevent the occurrence of the striped pattern as shown in FIG. 19 on the radiographic image, and the striped pattern is used when the doctor diagnoses the lesion using the radiographic image. It is possible to accurately prevent a doctor's misdiagnosis due to the occurrence of this.

  The present invention is not limited to the above embodiments, and it goes without saying that various modifications can be made without departing from the essence of the invention.

DESCRIPTION OF SYMBOLS 1 Radiographic imaging apparatus 5 Scanning line 6 Signal line 7 Radiation detection element 17 Reading circuit 20 A / D conversion circuit 22 Control means (correction value calculation means)
39 Antenna device (transfer means)
50 Radiation Imaging System 58 Output value from console Gi A / D conversion circuit (digital value image data)
P detector q correction value r area Rhigh, Rlow predetermined multiple doses

Claims (4)

A plurality of scanning lines and a plurality of signal lines arranged so as to intersect with each other; a plurality of radiation detecting elements arranged in a two-dimensional manner in each region partitioned by the plurality of scanning lines and the plurality of signal lines; A detector comprising:
A plurality of readout circuits for reading out charges from the radiation detection elements through the signal lines and outputting the charges as image data for each of the radiation detection elements;
With
Furthermore, in the calibration method of a radiographic imaging apparatus including a plurality of A / D conversion circuits that convert analog value image data output from the plurality of readout circuits into digital value image data, respectively,
A radiation irradiation step of irradiating the radiation imaging apparatus with a plurality of predetermined doses of radiation on a lower dose side than a half of the maximum dose of radiation that can be detected by the radiation detection element;
Each charge accumulated in the radiation detection element by irradiation with the plurality of doses of radiation is read as a plurality of analog value image data, and the read image data of the plurality of analog values is the A / D conversion circuit. A read conversion step of converting the image data into a plurality of digital values for each,
A correction value calculating step for calculating a correction value for the digital value image data for each of the A / D conversion circuits, based on the plurality of digital value image data converted for each of the A / D conversion circuits;
A calibration method for a radiographic imaging apparatus, comprising:   In the correction value calculating step, the change rate of the difference between the image data of the plurality of digital values for each of the A / D conversion circuits with respect to the difference between the predetermined plurality of doses is set to a preset reference change rate. 2. The calibration method for a radiographic image capturing apparatus according to claim 1, wherein a correction value for the image data of the digital value for each A / D conversion circuit is calculated. A plurality of scanning lines and a plurality of signal lines arranged so as to intersect with each other; a plurality of radiation detecting elements arranged in a two-dimensional manner in each region partitioned by the plurality of scanning lines and the plurality of signal lines; A detector comprising:
A plurality of readout circuits for reading out charges from the radiation detection elements through the signal lines and outputting the charges as image data for each of the radiation detection elements;
With
A plurality of A / D conversion circuits that convert analog value image data output from the plurality of readout circuits into digital value image data, respectively,
Furthermore, based on the digital value image data converted for each A / D conversion circuit, a correction value calculation means for calculating a correction value for the digital value image data for each A / D conversion circuit,
The correction value calculating means is configured to irradiate a plurality of predetermined doses of radiation on the dose side lower than half of the maximum dose of radiation that can be detected by the radiation detection element during calibration. When the accumulated charges are read out as a plurality of analog value image data and converted into a plurality of digital value image data for each of the A / D conversion circuits, each of the A / D conversion circuits A radiographic imaging apparatus, wherein a correction value for the digital image data for each of the A / D conversion circuits is calculated based on the converted digital image data. A plurality of scanning lines and a plurality of signal lines arranged so as to intersect with each other; a plurality of radiation detecting elements arranged in a two-dimensional manner in each region partitioned by the plurality of scanning lines and the plurality of signal lines; A detector comprising:
A plurality of readout circuits for reading out charges from the radiation detection elements through the signal lines and outputting the charges as image data for each of the radiation detection elements;
Transfer means for transferring the image data,
Furthermore, a radiographic imaging apparatus comprising a plurality of A / D conversion circuits for converting analog value image data output from the plurality of readout circuits into digital value image data, and
A console that performs logarithmic conversion processing on the image data transferred from the radiation image capturing apparatus to create final image data;
With
When the radiation image capturing apparatus is irradiated with a plurality of predetermined doses of radiation on a dose side lower than half of the maximum radiation dose that can be detected by the radiation detection element during calibration, the radiation detection element Each of the charges stored therein is read out as a plurality of analog value image data, converted into a plurality of digital value image data for each A / D conversion circuit, and transferred to the A / D conversion circuit Transferring image data of the plurality of digital values for each to the console;
When the console receives the image data of the plurality of digital values for each of the A / D conversion circuits from the radiographic imaging device, the image data of the digital values for each of the A / D conversion circuits based on the image data A radiographic imaging system characterized by calculating a correction value for.

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