JP2011188157A - Radiographic image photographing apparatus and radiographic image photographing system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To reduce the size of information relating to data compression processing while reducing influences of a defective pixel. <P>SOLUTION: A radiographic image photographing apparatus includes: a plurality of radiation detecting elements 7 arrayed in a two-dimensional shape by a plurality of scan lines and signal lines; a calculation means 491 for creating differential data from the radiation detecting elements adjacent in a direction of the signal lines; a compression means 492 for performing data compression processing; a transfer means 39 for performing data transfer; and an output failure determination means 491c for determining output failure from the differential data. For a radiation detecting element determined as output failure by differential data with a radiation detecting element at one side in the direction of the signal lines, the compression means excludes differential data with a radiation detecting element adjacent at the opposite side from a target of compression processing and, for the radiation detecting element adjacent at the opposite side, calculates differential data with radiation detecting elements around the radiation detecting element of the output failure. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to a radiographic image capturing apparatus and a radiographic image capturing system, and more particularly to a radiographic image capturing apparatus that compresses and transfers image data and a radiographic image capturing system that receives the image data and restores the original image data.

  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).

  By the way, in such a radiographic imaging apparatus, a plurality of radiation detection elements are arranged in a two-dimensional shape (matrix shape) to form a detection unit. At this time, the number of radiation detection elements (that is, the number of pixels) is Usually, the number of pixels is several million to several tens of millions or more. Therefore, if the image data read from each radiation detection element is transferred to the external device without being compressed, the transfer time becomes longer. Further, in a portable radiographic imaging apparatus with a built-in battery, when the transfer time of image data becomes long, the power consumed at the time of transfer increases, leading to battery consumption.

  Therefore, as described in, for example, Patent Document 4 and Patent Document 5, the read image data is usually reversible compression (also referred to as lossless compression) or lossy compression (also referred to as lossy compression). The data is compressed by a data compression method and transferred to an external device such as a console or a server.

  Then, for example, the radiographic imaging device is used as a medical image imaging device that images a part of a body such as a patient's head, chest, and limbs as a subject and uses the acquired radiographic image as a medical image for diagnosis or the like. In this case, as a data compression method for compressing image data, in general, image data before compression and after restoration are compared with a lossy compression method in which a part of information included in the image data is lost by the compression. It is considered preferable to employ a lossless compression method in which compression is performed so that the image data completely matches.

However, even when the above-described lossless compression method is employed, if there is a variation in each image data, sufficient compression is not performed, and the compression processing time, data transfer time, etc. may not be sufficiently reduced. It was.
Further, some radiation detection elements of the radiographic imaging apparatus may output an abnormal detection value from the beginning of manufacture, and this abnormal detection value hinders efficient compression of image data. It was.
Therefore, the position information of the radiation detection element that shows an abnormal detection value is stored in advance as a defective pixel map, and the radiation detection element that shows an abnormal detection value is not adopted at the time of radiographic imaging, and its surroundings 2. Description of the Related Art Radiation imaging apparatuses that employ detection values obtained by complementation processing from detection values of radiation detection elements have been developed. As a result, the influence of the abnormality of the radiation detection element is eliminated, and the dispersion of the image data is eliminated (see, for example, Patent Document 6).

JP-A-9-73144 JP 2006-058124 A JP-A-6-342099 JP 2000-275350 A JP 2005-287927 A JP 2009-172078 A

However, in the case of a radiographic imaging apparatus that eliminates the influence of a defective pixel by the above-described defective pixel map, it is necessary to identify a radiation detection element that shows an abnormal detection value in advance by inspection or the like and create a defective pixel map. There was a problem. Further, a defective pixel map cannot deal with a radiation detection element that will cause a detection value abnormality later, and the defective pixel map is updated by regular maintenance in order to solve this problem. Measures such as were necessary. Moreover, even if such measures were taken, it was not possible to cope with radiation detecting elements that showed abnormal values due to the influence of noise or the like during use.
Furthermore, the variation in each image data that hinders the efficient compression of the reversible compression method is not caused only by defective pixels, but more fundamentally, efficient compression is achieved by varying the shade of the image forming the captured image itself. In some cases, more stable image data compression is desired.

  Therefore, the present invention has been made in view of the above circumstances, and when performing compression processing using a static dictionary, efficient data compression is achieved by reducing unnecessary transfer time and time required for compression processing. An object of the present invention is to provide a radiographic image capturing apparatus and a radiographic image capturing system capable of performing data transfer.

  In order to solve the above problem, a radiographic imaging apparatus according to the present invention is divided by a plurality of scanning lines and a plurality of signal lines arranged so as to cross each other, and the plurality of scanning lines and a plurality of signal lines. A plurality of radiation detection elements arranged two-dimensionally in each region, and a charge is read from the radiation detection element through the signal line, and the charge is converted into an electrical signal for each radiation detection element. A readout circuit that outputs the data as image data, a calculation unit that calculates a difference between the image data output from adjacent radiation detection elements in a certain arrangement direction in the detection unit, and creates difference data; A compression means for compressing the difference data using a dynamic dictionary, a transfer means for transferring data, and for each of the radiation detection elements, the fixed array direction Output abnormality determination means for determining an output abnormality based on the difference data calculated from the difference between the image data with the radiation detection element adjacent to one side or the transfer data amount based on the difference data, and the compression means, When an output abnormality is determined for any radiation detection element by the output abnormality determination means, it is determined that the output abnormality is the radiation detection element adjacent to the radiation detection element adjacent to the opposite side in the certain arrangement direction. The difference data between the radiation detection element adjacent to the opposite side and the other radiation detection elements around the radiation detection element determined to be abnormal in output is not subjected to the difference data with the radiation detection element. It is characterized by.

  Moreover, the radiographic imaging system which is another side surface of this invention is the radiographic imaging apparatus as described in any one of Claims 1-7, and the compression process transferred from the said radiographic imaging apparatus A console for receiving data including difference data, and the console is configured to restore the compressed difference data to original difference data based on a static dictionary, and from the restored difference data. Image data calculating means for calculating image data, wherein the image data calculating means outputs the output abnormality for the image data of the radiation detecting element adjacent to the opposite side to the radiation detecting element that is abnormal in output. It is calculated based on difference data with the surrounding radiation detection elements and image data of the radiation detection elements.

  The radiographic image capturing apparatus and the radiographic image capturing system according to the present invention obtain difference data from image data of radiation detecting elements adjacent to each other, and compress the difference data using a static dictionary. In a plurality of radiation detection elements arranged two-dimensionally, two adjacent radiation detection elements are close to each other. The difference data between two adjacent radiation detection elements is mostly 0 or a value close to it. On the other hand, when a data compression process, especially a processing method that performs compression by data conversion using a static dictionary, is generally used when data variation is small and the number of individuals with common values increases. It is possible to compress more efficiently. In other words, by converting the difference data between adjacent radiation detection elements using a static dictionary, data compression can be performed with high efficiency, data transmission can be performed more quickly, and power can be saved. Can be achieved.

  Further, the output abnormality determination means detects a radiation detection element corresponding to a so-called defective pixel that does not perform proper output by determining a difference amount of image data from a radiation detection element adjacent thereto or a transfer data amount based on the difference data. The radiation detection element that performs abnormal output from the beginning, the radiation detection element that has started performing abnormal output later, the radiation detection element that occasionally performs abnormal output, and occasionally performs abnormal output when taking a certain radiographic image Thus, it is possible to specify all the radiation detection elements after imaging, and it is not necessary to create a map indicating the location of the defective radiation detection element in advance, and it is possible to eliminate maintenance for updating.

Further, the compression means is configured such that when one of the radiation detection elements is determined to be abnormal in output based on difference data from adjacent radiation detection elements on one side thereof, the radiation detection element determined to be abnormal in output and vice versa. The difference data with the radiation detecting element adjacent to the side is not subject to compression processing.
For example, when calculating difference data between adjacent radiation detection elements in a certain arrangement direction, if an abnormal output is performed for one radiation detection element, the difference value is large for one side and the opposite side of the abnormal radiation detection element. The difference data is likely to take a value that greatly deviates from the value with high appearance frequency, and the presence of the difference data is a factor that deteriorates the compression efficiency of the data. It becomes.
However, the compression means compresses the difference data with the radiation detection element adjacent on the opposite side when an output abnormality is determined based on the difference data with the radiation detection element adjacent to the radiation detection element on one side. Because it is excluded from the target, it is possible to at least halve the difference data that takes a value that deviates greatly from the concentration range of the appearance frequency distribution caused by the radiation detection element of abnormal output, and thereby the efficiency of compression of the entire difference data Can be improved.

It is an external appearance perspective view which shows the radiographic imaging apparatus in this 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. It is an enlarged view which shows the structure of the radiation detection element, TFT, etc. which were formed in the small area | region on the board | substrate of FIG. It is sectional drawing which follows the BB 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 block diagram showing the equivalent circuit of a radiographic imaging apparatus. It is a block diagram showing the equivalent circuit about 1 pixel which comprises a detection part. It is a figure explaining the state which image data read from the radiation detection element all at once by each reading IC is rearranged after being stored in a buffer memory, and is transmitted to a data memory. It is a block diagram showing the functional structure of the control means of a radiographic imaging apparatus. It is a graph which shows the range of the difference data by which Huffman code is matched and tabulated in this embodiment. It is a figure explaining the state in which difference data is calculated for every image data of a signal line direction. It is a figure explaining the state in which difference data is calculated for every image data of the direction of a scanning line. It is a figure explaining the structure of a difference calculation part, and the method of producing the difference data of the image data adjacent to the signal line direction connected to the same signal line. It is a figure explaining the state by which difference data is calculated from each image data read from each radiation detection element connected to the reference | standard data and the line L1 of a scanning line. (A)-(C) is a figure explaining the method of producing the difference data of the image data adjacent to the signal line direction connected to the same signal line using one buffer register. It is a virtual conceptual diagram in case the difference data with each radiation detection element of the both sides of a signal line direction are calculated in the radiation detection element in which abnormal output was performed. FIG. 18A is a conceptual diagram showing the flow of data in two buffer registers, and FIG. 18A shows a state in which each image data is moved from one to the other buffer register except for the radiation detection element with abnormal output. FIG. 18B shows a state in which new image data is accumulated in one buffer register, and FIG. 18C shows a state in which difference data is calculated by skipping a radiation detection element having an abnormal output. Distribution of appearance frequency of difference data of image data output from adjacent radiation detection elements connected to (A) same signal line or (B) same scanning line when radiation image capturing apparatus is irradiated with radiation uniformly It is a graph which shows. When the radiation dose to the radiographic imaging device is increased, (A) the frequency of appearance of difference data of image data output from adjacent radiation detection elements connected to the same signal line or (B) the same scanning line It is a graph which shows distribution. It is a graph showing an example of distribution of the appearance frequency of the image data read from each radiation detection element. It is a figure which shows the structural example of the data for transfer. It is a figure which shows the whole structure of the radiographic imaging system in this embodiment. It is a flowchart which shows the process of the compression in this embodiment. FIG. 25A is a conceptual diagram showing a process for calculating image data from difference data in the console. FIG. 25A adds the difference data of one buffer register and the image data of the other buffer register to update the other buffer register. FIG. 25B shows a state in which the other buffer register is not updated for the image data of the radiation detecting element with abnormal output, and FIG. 25C shows new difference data for one buffer register. The processing in a state where is accumulated.

  Embodiments of a radiographic image capturing apparatus and a radiographic image capturing system according to the present invention will be described below with reference to the drawings.

[Radiation imaging equipment]
First, the configuration of the radiographic image capturing apparatus according to the present embodiment will be described with reference to FIGS. In the following description, the radiographic imaging device is a so-called indirect radiographic imaging device that includes a scintillator or the like and converts the irradiated radiation into electromagnetic waves of other wavelengths such as visible light to obtain an electrical signal. As will be described, the present invention can also be applied to a direct radiographic imaging apparatus. Although the case where the radiographic image capturing apparatus is portable will be described, the present invention is also applicable to a radiographic image capturing apparatus formed integrally with a support base or the like.

  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 the line AA in FIG. As shown in FIGS. 1 and 2, the radiographic image capturing apparatus 1 according to the present embodiment is configured by housing a scintillator 3, a substrate 4, and the like inside 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 where the housing 2 is a so-called lunch box type formed by the frame plate 2A and the back plate 2B, but the shape of the housing 2 is not limited to this. . It is also possible to use a so-called monocoque type in which the housing 2 is integrally formed in a rectangular tube shape.

  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 wirelessly transferring image data or the like to an external device such as a console 58 (see FIG. 23) described later is embedded in the side surface of the lid member 38. It is. 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.

  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 switching element, 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 an on-voltage is applied to the connected scanning line 5 and applied to the gate electrode 8g by a scanning driving means 15 (see FIG. 7) described later, and the radiation detection element 7 is turned on. The electric charge generated and accumulated therein 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 BB in FIG.

  As shown in FIG. 5, 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 the gate electrode 8g and the surface 4a. The radiation detecting element 7 is disposed above the gate electrode 8g on the gate insulating layer 81 made of silicon nitride (SiNx) or the like laminated on the semiconductor layer 82 made of hydrogenated amorphous silicon (a-Si) or the like. The source electrode 8 s connected to the first electrode 74 and the drain electrode 8 d 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 (SiNx) or the like, and the first passivation layer 83 covers 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 The upper side is covered with a second passivation layer 79 made of silicon nitride (SiNx) or the like.

  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, a correlated double sampling circuit 19, an analog multiplexer 21, and an A / D converter 20. 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.

  The signal line 6 is connected to the inverting input terminal on the input side of the operational amplifier 18a of the amplifier circuit 18, and the reference potential V0 is applied to the non-inverting input terminal on the input side of the amplifier circuit 18. Yes. The reference potential V0 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 via the signal), the electric charge released from the radiation detection element 7 flows into the capacitor 18b and is accumulated, and a voltage value corresponding to the accumulated electric charge is output from the operational amplifier 18a. Output from the side.

  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. When the charge reset switch 18c is turned on, the input side and the output side of the amplifier circuit 18 are short-circuited, and the charge accumulated in the capacitor 18b is discharged to reset the amplifier circuit 18. ing. 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.

  At the time of reading the image data from each radiation detection element 7, the charge is read from each radiation detection element 7, and the voltage value that is output after being subjected to charge-voltage conversion by the amplifier circuit 18 is sampled by the correlated double sampling circuit 19. It is processed and output downstream as image data. The image data of 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 converter 20. Then, the A / D converter 20 sequentially converts the image data into digital values, which are output to the storage means 40 and sequentially stored.

  In the present embodiment, when the image data is read from each 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 the above-described respective A process for reading image data from the radiation detection element 7 is performed.

  In the present embodiment, for example, 128 signal lines 6 are processed by one readout IC 16. That is, one read IC 16 corresponds to each signal line 6 with 128 read circuits 17 (that is, amplifier circuit 18 and correlated double sampling circuit 19 etc.), one analog multiplexer 21, and one A / D. It is formed by a converter 20 or the like.

  If the number of signal lines 6 is 2048, for example, 2048/128 = 16 readout ICs 16 are arranged in parallel to form a readout unit. Hereinafter, the number of readout circuits 17 formed in one readout IC 16, that is, the number of signal lines 6 connected to one readout IC 16 is 128, and the total number of signal lines 6 is 2048. However, the present invention is not limited to this case.

  As shown in FIG. 9, when an on-voltage is applied to the line L <b> 1 of the scanning line 5, for example, when the image data is read out, each radiation detection element (1, 1) connected to the line L <b> 1 of the scanning line 5. ) To (1,2048) are simultaneously read out and sent to the reading ICs 16 in parallel.

  Then, each read circuit 17 (not shown in FIG. 9) of each read IC 16 performs charge-voltage conversion and the like, and each of the 128 image data transmitted in parallel is converted into each analog multiplexer 21 (in each read IC 16). The image data is serially transferred to the A / D converter 20 (not shown) sequentially and digitized image data is temporarily stored in the buffer memory 45 from the A / D converter 20.

  That is, when image data corresponding to each radiation detection element (x, y) is represented as D (x, y), first, from each readout IC 16, D (1, 1), D (1, 129), D ( 1, 257),..., D (1, 1921) are transmitted and stored in the buffer memory 45, and then D (1, 2), D (1, 130), D (1, 258). ,..., D (1, 1922) are transmitted and stored in the buffer memory 45.

  Then, the image data D (1, 1) to D (1, 2048) from the radiation detection elements (1, 1) to (1, 2048) connected to the line L1 of the scanning line 5 are stored in the buffer memory 45. Are stored, the image data D are rearranged in the order of image data D (1,1), D (1,2), D (1,3), D (1,4),. The image data D (difference data ΔD based on the image data D) is sequentially transmitted and stored in the data memory 46 in this order.

  In addition, the reading process of each image data D (1,1) to D (1,2048) from each radiation detection element (1,1) to (1,2048) connected to the line L1 of the scanning line 5 is completed. Then, the line of the scanning line 5 to which the ON voltage is applied is subsequently switched to L2. Similarly, the image data D (2,1) to D (2,2048) are transmitted to the buffer memory 45 for each readout IC 16 and rearranged, and then sequentially transmitted to the data memory 46 and stored. The

  The reading process and the storing process in the data memory 46 are sequentially repeated for each of the lines L1 to Lx of the scanning line 5 so that the reading process of the image data D from all the radiation detection elements 7 is performed. It has become.

  The rearrangement process of the image data D is normally performed by converting the image data D into D (1,1), D (1,1) regardless of the external device (not shown) that transfers the image data D. 2), D (1,3), D (1,4),... Can be handled by transferring them in this order. Therefore, the image data D is generally used at the stage of storing the image data D in the data memory 46. This is a process for rearranging and storing D in the above order.

  Accordingly, when the order of transferring the image data D from the radiation image capturing apparatus 1 to the external apparatus can be determined in advance, the image data D can be rearranged according to the determined order. .

  In addition, each image data D is preliminarily output from the radiation imaging apparatus 1 to the external device, for example, in the order of output from each readout IC 16, D (1, 1), D (1, 129),..., D (1, 1921), If it is determined that the data is transferred in the order of D (1,2), D (1,130),..., D (1,1922),. It is also possible to directly transmit and store the data directly in the data memory 46 without going through 45.

  When rearranging the image data D as described above, the image data D is rearranged when the image data D is read from the data memory 46, not when the image data D is stored in the data memory 46. It is also possible to configure so that

  In the present embodiment, the image data D read from each radiation detection element 7 as described above is stored in the data memory 46, and in parallel with this, difference data ΔD, which will be described later, is calculated, compressed, and emitted. The image data may be transferred from the image capturing apparatus 1 to an external device (not shown), or each image data D read from each radiation detection element 7 is stored in the data memory 46 and then the difference from each stored image data D is stored. It is also possible to perform processing so that the data ΔD is calculated and transferred to the outside.

[Control means]
Here, the structure of the control means 22 in this embodiment is demonstrated, referring FIG.7 and FIG.10. In FIG. 10, the outline of the processing flow of the image data read from the reading circuit 17 is indicated by a broken line.
As shown in FIG. 10, in this embodiment, the control means 22 includes a main body control unit 23, an FPGA (Field Programmable Gate Array) 24, a data memory 46, and the like.
The control means 22 may be constituted by a dedicated control circuit.

In the present embodiment, the control means 22 is connected to a storage means 40 constituted by a DRAM (Dynamic RAM) or the like and the antenna device 39 described above.
Further, as shown in FIG. 7, a battery 41 for supplying power to each member such as the detection unit P, the scanning drive unit 15, the readout circuit 17, the storage unit 40, and the bias power supply 14 is connected to the control unit 22. Has been. 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 unit 22 (main body control unit 23) 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 to the amplification circuit 18 of the readout circuit 17. Various processes such as on / off control of the charge reset switch 18c and transmission of a pulse signal to the correlated double sampling circuit 19 to control on / off of the sample hold function are executed. ing.

  In addition, the control means 22 (main body control unit 23) applies to the scanning drive means 15 at the time of reset processing of each radiation detection element 7 or at the time of reading the image data D from each radiation detection element 7 after radiographic imaging. 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 is transmitted from the scanning drive means 15 via each scanning line 5.

The main body control unit 23 is a computer including a CPU (Central Processing Unit) 23a, a ROM (Read Only Memory) 23b, a RAM (Random Access Memory) 23c, and the like, and the operation of each functional unit of the radiation image capturing apparatus 1 and the like. Is to control.
In the ROM 23b, for example, a program for performing data compression processing, information relating to data compression processing, that is, a table of Huffman codes Hc used for Huffman coding in this embodiment, threshold values for determining whether or not to perform compression processing are stored. (Details will be described later) are stored.
Various programs, information, and the like are not limited to being stored in the ROM 23b of the main body control unit 23, and a separate program memory or the like may be provided and stored therein.

Here, a table of the Huffman code Hc will be described.
The Huffman code Hc table is a static dictionary used for image data compression processing in this embodiment. The static dictionary is not limited to the Huffman code Hc table, and a static dictionary corresponding to this is prepared when another method is used as the compression processing method.
In this embodiment, a Huffman code Hc table is created in advance and stored in a memory such as the ROM 23b.

In Huffman compression, it is more efficient to prepare an appropriate Huffman code Hc table according to imaging conditions such as an imaging region such as a chest or abdomen of a patient's body as a subject and a radiation dose. It is possible.
Therefore, when creating a table of Huffman codes Hc for each of various imaging conditions, radiographic imaging is performed for each imaging condition, image data D is read from each radiation detection element 7, and the same signal line The difference between the image data D of the adjacent radiation detection elements 7 connected to 6 is calculated as difference data ΔD, and the appearance frequency F of each difference data ΔD is plotted as shown in FIG. Then, for example, a table is created in advance by associating the difference data ΔD with the Huffman code Hc for the difference data ΔD in the range of X to Y in the figure that appears at the appearance frequency F equal to or higher than the threshold.

  That is, the Huffman code Hc is assigned by a known method from the value of the difference data ΔD within the range of X to Y where the appearance frequency F is equal to or greater than the threshold and the individual appearance frequencies F, and the difference between the individual difference data ΔD and the Huffman code Hc. The correspondence relationship is tabulated and stored in the ROM 23b. The ROM 23b may be non-rewritable or may be rewritable. In the former case, an appropriate Huffman code Hc stored in advance may be mounted in advance, and in the latter case, after obtaining a table of the appropriate Huffman code Hc, Just write. As the Huffman code Hc table, it is desirable to assign a table having a shorter code length as the appearance frequency F of the difference data ΔD increases.

  In the present embodiment, a case where one table of Huffman codes Hc is provided will be described. Since the table of the Huffman code Hc has a relatively large amount of data, when one table is used as described above, the memory capacity necessary for storing the table can be reduced accordingly. It is. An imaging region and imaging condition setting means is provided in the radiographic image capturing apparatus 1 or the console 58 described later, and a plurality of types of tables are prepared according to the imaging region and the like. An optimal table may be selected as appropriate and the compression process may be performed.

The range in which the table is created for the appearance frequency F or the values X and Y of the difference data ΔD (that is, how many threshold values are set) may be set in advance by default, It may be set arbitrarily.
If a table is created for a wide range of difference data ΔD, the amount of data that can be encoded by compression processing increases accordingly. However, if a table is created by importing data with a low appearance frequency F, the code length of the Huffman code Hc is increased. Even if the compression process is performed and the difference data ΔD is replaced with the Huffman code Hc, the compression effect is not so much improved, or conversely, the data amount may be larger than the original image data. For this reason, it is preferable that the setting of the threshold value is determined in consideration of the compression efficiency. Then, as will be described later, a special code indicating that compression processing is not performed on the original image data is added to a portion where the appearance frequency F is below this threshold value and the Huffman code Hc is not assigned. This can be done by performing special code processing.

  The FPGA 24 constituting the control means 22 is provided with a buffer memory 45, a transfer data generation unit 49, a register unit 49d, and the like.

  In the present embodiment, the buffer memory 45 temporarily stores the original image data (that is, the raw signal value and the pixel value) read from the readout circuit 17 and stores the original image data (that is, the raw signal value). , Pixel values) are sent to the data memory 46, and the original image data is sequentially stored in the data memory 46. Further, the image data in the data memory 46 is sent to the compression processing unit 49a of the transfer data generation unit 49 to be processed later, but the data is also sent directly from the buffer memory 45 to the compression processing unit 49a. May be transferred.

The transfer data generation unit 49 is a functional unit that generates transfer data to be transferred to an external device based on the difference data ΔD.
The transfer data generation unit 49 adds compression section 49a as compression means in the present invention that performs compression processing on the difference data ΔD, and adds information on the delimiter position of the data to the compressed data for each line. An information adding unit 49b and a data generation unit 49c that generates transfer data for each line are provided.

The compression processing unit 49a is a compression unit that performs compression processing of image data using a static dictionary.
In the present embodiment, the compression processing unit 49a performs a difference calculation unit 491 that calculates the difference data ΔD for the image data D, and a compression process by Huffman coding described later on the difference data ΔD generated by the difference calculation unit 491. And an encoding processing unit 492 that performs the processing.

  The difference calculation unit 491 includes two buffer registers 491a and 491b (see FIG. 14), reads the original image data D stored in the data memory 46, and calculates difference data ΔD for the image data D. It is a calculation means for doing.

In the present embodiment, a case where the difference calculation unit 491 is provided in the FPGA 24 will be described as an example. However, an existing register in the main body control unit 23 that is a computer may be used as the difference calculation unit 491. Is possible. Furthermore, in this embodiment, the case where the difference calculation unit 491 is provided with two buffer registers 491a and 491b is illustrated, but the number of buffer registers provided in the difference calculation unit 491 is not particularly limited. For example, one buffer register may be provided, or three or more buffer registers may be provided.
Further, each buffer register 491a, 491b is exemplified as having a capacity capable of storing image data of the radiation detection element 7 for one line along the direction of the scanning line 5, but more or less than this. It may be a capacity.

[Difference calculation processing]
Here, a method for calculating the difference value for the image data D by the difference calculation unit 491 and generating the difference data ΔD will be described.
In the present embodiment, as described below, the difference data ΔD generated by the difference calculation unit 491 is output to the encoding processing unit 492, and the encoding processing unit 492 performs compression processing on the difference data ΔD. After that, a case will be described in which the difference data ΔD sent to the main body control unit 23 and stored in the storage means 40 of the radiographic image capturing apparatus 1 is stored (see the broken line shown in FIG. 10), but the compression processing has been performed. The difference data ΔD can also be configured to be transferred from the antenna device 39 to the external device wirelessly without being stored in the storage unit 40.

  Note that the order in which the image data D is processed when creating the difference data ΔD is not particularly limited. However, as shown in FIG. When each line Ln is stored in the data memory 46 or the like in the order of D (n, 1), D (n, 2), D (n, 3), D (n, 4),. Each image data can be easily read by configuring the storage and reading to 46 and the like in the same order.

  Therefore, in the present embodiment, as shown in FIG. 12, the difference between the image data D output from the adjacent radiation detection elements 7 connected to the same signal line 6 while reading each image data D in the above order. Is calculated as difference data ΔD. That is, the difference between the image data D from each radiation detection element 7 adjacent to the extending direction of the signal line (hereinafter referred to as the signal line direction) represented by the vertical arrow in the figure is calculated as difference data ΔD. It is supposed to be.

  For example, when the difference between the image data D output from the adjacent radiation detection elements 7 connected to the same scanning line 5 is calculated as the difference data ΔD, as shown in FIG. Are calculated in the above order, the difference between the image data D from each radiation detection element 7 adjacent in the direction of the scanning line 5 (the horizontal arrow direction in the figure) is calculated, and the difference data ΔD is calculated. be able to. In that case, for example, D (1,1) -D (1,2), D (1,2) -D (1,3), D (1,3) -D (1,4),. The calculation process is performed as described above.

Specifically, in the present embodiment, the difference calculation unit 491 is provided with two buffer registers 491a and 491b as shown in FIG.
When the difference value is calculated and the difference data ΔD is generated, the difference calculation unit 491 scans the data memory 46 from each radiation detection element 7 connected to the line Ln of the scanning line 5. The image data D (n, 1), D (n, 2), D (n, 3),... Arranged in the line direction are read out and temporarily stored in the buffer register 491a. Further, the image data D (n + 1, 1), D arranged in the scanning line direction read from the radiation detecting elements 7 connected to the line Ln + 1 adjacent to the line Ln of the scanning line 5 from the data memory 46. (N + 1, 2) and D (n + 1, 3) are read out and temporarily stored in the buffer register 491b.

  Then, by calculating the difference ΔD (ie, ΔD (n + 1,1), ΔD (n + 1,2), ΔD (n + 1,3),...) Between the image data D at the same address in the buffer registers 491a and 491b, The difference data ΔD between the image data D of the radiation detecting elements 7 adjacent to each other in the signal line direction connected to the same signal line 6 is calculated.

  At that time, if the image data D arranged in the scanning line direction for two adjacent lines is read from the data memory 46 every time in order to calculate the difference data ΔD, the reading control becomes troublesome.

  Therefore, in this embodiment, the difference calculation unit 491 performs buffer register 491a, 491b for each image data D read from each radiation detection element 7 connected to each line Ln, Ln + 1 of the adjacent scanning line 5. When the difference ΔD (ie, ΔD (n + 1, 1), ΔD (n + 1, 2), ΔD (n + 1, 3),...) Between the image data D at the same address is calculated in the order of the addresses, each image data D (n + 1, 1 ), D (n + 1, 2), D (n + 1, 3),... Are moved from the buffer register 491b to the buffer register 491a, and the scanning line direction of the line Ln + 2 of the scanning line 5 next to the empty buffer register 491b Are stored in the image data D (n + 2, 1), D (n + 2, 2), D (n + 2, 3),.

  When the difference data ΔD (n + 2, 1), ΔD (n + 2, 2),... Is calculated, the image data D (n + 2, 1), D (n + 2, 2),... Are transferred from the buffer register 491b to the buffer register 491a. The image data D (n + 3, 1), D (n + 3, 2),... Are accumulated in the buffer register 491b. In this way, the process of calculating the difference ΔD between the image data D at the same address in the buffer registers 491a and 491b while transferring each image data D from the buffer register 491b to the buffer register 491a is repeated to sequentially calculate the difference data ΔD. It has come to do.

  When configured in this way, the image data D (1,1), D (1,2), D arranged in the scanning line direction read from the radiation detecting elements 7 connected to the line L1 of the scanning line 5 .. Are required as data for calculating the difference data ΔD (1,1), ΔD (1,2), ΔD (1,3),. Therefore, in this embodiment, preset reference data Dc (1), Dc (2), Dc (3),... Are stored in advance in a memory such as a ROM.

  Then, when calculating the difference data ΔD (1,1), ΔD (1,2), ΔD (1,3),..., The difference calculation unit 491 reads from the memory as shown in FIG. The reference data Dc (1), Dc (2), Dc (3),... Are stored in the buffer register 491a and read from each radiation detection element 7 connected to the line L1 of the scanning line 5 read from the data memory 46. The image data D (1,1), D (1,2),... Arranged in the scanning line direction is accumulated in the buffer register 491b, and the difference ΔD is obtained as the difference data ΔD (1,1), ΔD ( 1, 2),...

  At this time, each value of the reference data Dc (1), Dc (2),... May be set to the same value, or may be set to different values. Alternatively, a plurality of types may be prepared according to the imaging region and imaging conditions.

As described above, even when only one buffer register 491a is provided in the difference calculation unit 491, the difference between the image data D adjacent in the signal line direction similar to the above is calculated to calculate the difference data ΔD. It can be configured as follows.
That is, as shown in FIG. 16A, the image data D (n, 1), D (n, 2), D (n, 3),... Of the line Ln of the scanning line 5 are accumulated in the buffer register 491a. In this case, the control unit 22 sequentially outputs the image data D (n + 1, 1), D (n + 1, 2), D (n + 1, 3),... Of the line Ln + 1 of the adjacent scanning line 5 from the data memory 46. As shown in FIGS. 16B and 16C, the difference data from the corresponding image data D (n, 1), D (n, 2), D (n, 3),. While calculating ΔD, processing for sequentially replacing the data in the buffer register 491a with each image data D of the line Ln + 1 is performed.
With this configuration, even when there is only one buffer register 491a, it is output from the adjacent radiation detection elements 7 connected to the same signal line 6 in the same manner as shown in FIGS. It is possible to calculate difference data ΔD between the image data D.

[Determination of output abnormality of radiation detector]
Although the value of the image data itself output from each radiation detection element 7 is expected to vary for each element 7 according to the incident dose of radiation, since the radiation detection element 7 is very small, adjacent elements 7 are In many cases, radiation is incident with substantially the same intensity, and when the difference data ΔD is obtained from the difference between the image data between the adjacent radiation detection elements 7, for example, as shown in FIG. 11, the value of each difference data ΔD It can be seen that the distribution concentrates on values around 0 centered on 0. When a compression method using a static dictionary is used, data compression with higher efficiency can be realized as the values of the difference data ΔD are concentrated in a narrow numerical range.

However, some of the radiation detection elements 7 perform abnormal output later in the manufacturing stage or after manufacturing, and some cause abnormal output due to the influence of noise or the like. Since an output different from the value indicating the dose is performed, only one specific image data is output as compared with the surrounding radiation detection element 7. When such image data is output, the difference calculation unit 491 has the following problem when obtaining a difference between adjacent radiation detection elements 7 in a certain direction (for example, a direction along the signal line 6). Arise.
FIG. 17 is a conceptual diagram when abnormal output as described above is performed in the radiation detection element 7 that outputs the image data of D (2, 2), and the difference data ΔD is calculated in each radiation detection element 7. As shown in this figure, when the image data of D (2, 2) has an abnormal value, two image data D (1, 2), D (3, 2) positioned before and after in the direction of the signal line 6 are used. Difference data ΔD (2, 2) and ΔD (3, 2) are abnormal values, and deviate from the range where the distribution of the appearance frequency F is concentrated. That is, an abnormality occurs in two differential data ΔD that is twice that due to an output abnormality of one radiation detection element 7. When performing Huffman compression or the like using a static dictionary, the compression efficiency decreases as the number of values greatly deviating from the range in which the distribution of the appearance frequency F is concentrated is as described above. .

  Therefore, the difference calculation unit 491 is provided with an output abnormality determination unit 491c serving as an output abnormality determination unit that determines an output abnormality of each radiation detection element 7 from the calculated difference data ΔD. That is, when the difference calculation unit 491 calculates the difference data ΔD, the output abnormality determination unit 491c determines the output abnormality of the radiation detection element 7 from the value of the difference data ΔD. A process of excluding the radiation detection element 7 that has performed an abnormal output in the line Ln from the calculation target of the difference data ΔD with the adjacent radiation detection element 7 in the next line Ln + 1 is executed.

  The output abnormality determination unit 491c sets a numerical range that is regarded as a normal output for the calculated difference data ΔD, and determines that the output is normal if it is within the range, and that the output is abnormal if it is outside the range. The numerical value range regarded as normal is desirably matched with the range X to Y that is equal to or higher than the threshold value for the appearance frequency F when determining whether or not to perform the compression processing of the difference data ΔD described above. The range may be set wider or narrower than the range of X to Y. In this case, the user may arbitrarily set the numerical value range that regards normal output.

  When the output abnormality determination unit 491c determines that the output is abnormal, the difference calculation unit 491 is read from each radiation detection element 7 connected to the line of the scanning line 5 stored in the buffer register 491b. When the image data D arranged in the scanning line direction is moved to the buffer register 491a, only the image data of the radiation detection element 7 determined to have the output abnormality is excluded.

For example, in the specific example shown in FIG. 18, when the output abnormality determination unit 491c determines that the image data D (n + 1, 2) is an output abnormality (FIG. 18A), D (n + 1, 2) is When moving each image data D read from each radiation detecting element 7 connected to the scanning line 5 including the buffer register 491b to 491a, only the image data D (n + 1, 2) determined to be output abnormal is transferred. I won't let you. As a result, the values of the image data D (n + 1, 1), D (n, 2), D (n + 1, 3), D (n + 1, 4),... Are arranged in the buffer register 491a (FIG. 18). (B)).
In the buffer register 491b, each image data D (n + 2, 1), D (n + 2, 2), D (n + 2, 2) read from each radiation detecting element 7 connected to the next line Ln + 2 of the scanning line 5 is stored. 3), D (n + 2, 4),... Are accumulated.
Then, difference data ΔD (n + 2,1), ΔD (n + 2,2), ΔD (n + 2,3), ΔD (n + 2,4),... Are calculated (FIG. 18C). At this time, ΔD (n + 2, 2) is not the difference between the radiation detecting elements 7 connected to the same signal line 6 and adjacent to each other, but the direction of the signal line 6 across the radiation detecting element 7 determined to be abnormal in output. Is calculated from the difference between the image data D (n, 2) and D (n + 2, 2) of the radiation detecting elements 7 and 7 on both sides.

Thereby, when the radiation detection element 7 with abnormal output is included, the difference data ΔD is not calculated for one side of the arrangement direction connected to the same signal line 6 with respect to the radiation detection element 7. For this reason, it is possible to effectively avoid the calculation of the two difference data ΔD having an abnormal value for one radiation detection element 7 that performs an abnormal output, thereby reducing the variation of the difference data ΔD and improving the efficiency. It makes it possible to perform good compression.
Note that the radiation detection element 7 adjacent on the line Ln + 2 next to the radiation detection element 7 from which the image data D (n + 1, 2) has been output may also be continuously output abnormally. The image data D (n + 2, 2) output from the element 7 is not shifted from the buffer register 491b to 491a, and as a result, the value of D (n, 2) is maintained as it is in the buffer register 491a. Therefore, when the image data on the next line Ln + 3 is accumulated in the buffer register 491b, the value of the difference data ΔD (n + 3, 2) is calculated by the difference between D (n, 2) and D (n + 3, 2). The Rukoto.

Further, as in the example of FIG. 16 described above, when there is only one buffer register 491a in the difference calculation unit 491, each image data D (n, 1), D (n, 2), D (n, 3),... Are stored in the buffer register 491a, the difference calculation unit 491 outputs the image data D (n + 1,1) and D (n + 1) of the line Ln + 1 of the adjacent scanning line 5. , 2), D (n + 1, 3),... Are sequentially read out from the data memory 46, and as shown in FIGS. 16B and 16C, the corresponding image data D (n, 1), D, respectively. (N, 2), D (n, 3),... Are calculated one by one, and when the obtained difference data ΔD is determined as an output abnormality by the output abnormality determination unit 491c, The image data on the line Ln + 1 is stored in the buffer register 4 Only when the difference data ΔD is determined to be normal output by the output abnormality determination unit 491c without being replaced with 91a, the data in the buffer register 491a is replaced with each image data D of the line Ln + 1.
As a result, it is possible to perform the same processing as when there are two buffer registers 491a and 491b.

The output abnormality determination unit 491c determines the output abnormality of the radiation detection element 7 based on whether or not the difference data ΔD is within a predetermined numerical range X to Y, but is not limited thereto. It is not something. For example, when the difference data ΔD is obtained for one radiation detection element 7, the data amount (number of bits) of the data after processing by the encoding processing unit 492 is calculated, and the value is set in a predetermined data amount threshold range. When it exceeds, you may determine with the output abnormality of the radiation detection element 7. FIG.
Here, the encoding processing unit 492 performs a process of compressing and transferring only the difference data ΔD that falls within the predetermined numerical value range X to Y with respect to the difference data ΔD. Processing for transferring image data with a special code added is performed (details will be described later).
Therefore, as described above, when the output abnormality of the radiation detection element 7 is determined based on the data amount (number of bits) of the data after processing by the encoding processing unit 492, the difference data ΔD is the target of compression. In such a case, the determination is made based on the data amount of the compressed data, and if it is not the object of compression, the determination is made based on the data amount of the image data to which the special code is added.

[Compression processing]
The encoding processing unit 492 is a functional unit that performs compression processing on the difference data ΔD calculated by the difference calculation unit 491.
As described above, when the radiographic image capturing apparatus 1 is used as a medical image capturing apparatus for capturing a part of a patient's body as a subject and using the acquired radiographic image as a medical image for diagnosis or the like, compression is performed. As a method, it is preferable to employ a lossless compression method in which compression is performed so that the difference data ΔD before compression and the difference data ΔD after restoration completely match.

  In the present embodiment, as described above, the Huffman coding method is employed as the lossless compression method. In the following, a case where the differential data ΔD is compressed by the Huffman coding method will be described. However, the compression method is not necessarily required by the Huffman coding, and the loss data of the differential data ΔD can be obtained using another lossless compression method. It is also possible to perform a compression process.

  As described above, a table of the Huffman code Hc as a static dictionary is stored in the memory such as the ROM 23b, and the encoding processing unit 492 refers to the table of the Huffman code Hc and adds the Huffman code to the difference data ΔD. A compression process is performed by assigning and encoding the code Hc. In data compression by Huffman coding, as is well known, data having a higher appearance frequency F (see FIG. 11 and the like) is assigned a shorter Huffman code Hc.

Here, the Huffman coding (Huffman coding) process will be described in detail with reference to FIGS.
When the luminance gradation of the image data D read from each radiation detection element 7 is, for example, 2 ¹⁶ (= 65536) gradations, for example, in the normal Huffman coding method, 2 ¹⁶ for each luminance gradation. Various types of Huffman codes Hc must be prepared, and the table for associating the image data D (difference data ΔD based thereon) with the Huffman codes Hc becomes very large.

  Therefore, as described above, in the radiographic imaging device 1 according to the present embodiment, not all the difference data ΔD is subjected to Huffman coding, but the image data D output from the adjacent radiation detection elements 7 is exchanged. Only a part of the difference data ΔD calculated as the difference is compressed.

  As described above, the ROM 23b of the main body control unit 23 stores a table associating a part of the difference data ΔD, which is information related to data compression processing for performing compression processing, with the Huffman code Hc.

  Then, when the difference data ΔD does not belong to the part of the difference data ΔD, the encoding processing unit 492 does not perform compression processing on the data, and the difference data ΔD becomes the part of the difference data ΔD. The compression process is performed with reference to the table only when belonging to the group. Then, the encoding processing unit 492 controls the main body by combining the compressed difference data ΔD when compression is performed, and the original image data (that is, the raw signal value) of the difference data ΔD when compression is not performed. The data is sent to the unit 23 and transferred from the main body control unit 23 to the external device.

  Of the difference data ΔD, which range is compressed and which range is uncompressed, that is, which difference data ΔD is associated with the Huffman code Hc and tabulated, and which difference data ΔD Is selected based on the distribution of the appearance frequency F (see FIG. 11 and the like) of the difference data ΔD calculated from the image data D read from each radiation detection element 7 of the radiographic imaging device 1. It is determined.

  That is, for example, in the distribution of the appearance frequency F of the difference data ΔD, the difference data ΔD having a high appearance frequency F, that is, the difference data ΔD having an appearance frequency F higher than a set threshold (that is, X˜ in FIG. 11 described above). A table is created by associating the Huffman code Hc with the difference data ΔD only for the range Y). No table is created for the difference data ΔD other than the difference data ΔD.

  Specifically, when the compression or non-compression target is the difference data ΔD, the distribution of the appearance frequency F of the difference data ΔD that is the difference between the image data D read from the adjacent radiation detection elements 7 is, for example, FIG. The distribution is as shown in A) and (B). In this case, for example, a table is created only for the difference data ΔD that appears at the appearance frequency F equal to or higher than the threshold.

  19A shows the difference data ΔD of the image data D output from the adjacent radiation detection elements 7 connected to the same signal line 6 when the radiation image capturing apparatus 1 is uniformly irradiated with radiation. FIG. 19B is a graph showing the distribution of the appearance frequency F. FIG. 19B is output from the adjacent radiation detection elements 7 connected to the same scanning line 5 when the radiation image capturing apparatus 1 is uniformly irradiated with radiation. 6 is a graph showing a distribution of appearance frequency F of difference data ΔD of image data D. Even when the radiation image capturing apparatus 1 is irradiated with radiation through a subject, the distribution of the appearance frequency F of each difference data ΔD is substantially symmetrical about ΔD = 0 as in FIGS. 19A and 19B. I know that

  In addition, when the compression or non-compression target is the difference data ΔD, when the radiation dose irradiated to the radiographic imaging apparatus 1 is increased, the distribution of the appearance frequency F of the difference data ΔD is shown in FIGS. As shown in FIG. 19B, the width of the distribution is slightly widened as can be seen from FIGS. 19A and 19B. Although not shown, the distribution width is widened or narrowed even when the photographing conditions including the photographing part of the patient's body as the subject are different.

  For this reason, a plurality of tables of Huffman codes Hc, which are information relating to data compression processing, are stored in advance in the memory of the control means 22 for each shooting condition, and set when the control means 22 performs data compression processing of the difference data ΔD. It is possible to select a table corresponding to the photographing condition and perform data compression processing on the difference data ΔD with reference to the selected table.

  However, according to the study by the present inventors, as shown in FIGS. 19A and 19B and FIGS. 20A and 20B, they are connected to the same signal line 6 and the same scanning line 5. It is known that when the difference data ΔD that is the difference between the image data D output from the adjacent radiation detection elements 7 is calculated, the difference data ΔD is distributed in a relatively narrow range.

  Further, in the study by the inventors of the present application, even if the imaging condition including the imaging region is variously changed and the width of the distribution of the appearance frequency F of the difference data ΔD is widened, for example, the difference data that appears at the appearance frequency F equal to or higher than the threshold value. It has been found that the range of ΔD falls within a part of the entire range of values that can be taken by the difference data ΔD (−65535 to 65535 in this embodiment).

  Therefore, one of the distributions of the appearance frequency F of the difference data ΔD that is created by setting in advance the range of the difference data ΔD that associates the Huffman code Hc, that is, the range of the difference data ΔD that creates a table (information on data compression processing). It is possible to use only a table, and this embodiment will be described in this embodiment.

  As shown in FIG. 19A and FIG. 20A, the appearance frequency F of the difference data ΔD between the image data D output from the adjacent radiation detection elements 7 connected to the same signal line 6 is , ΔD = 0 becomes a distribution with a normal distribution. This is considered to be due to manufacturing variation of each radiation detection element 7 that occurs when each radiation detection element 7 is formed by stacking each layer as shown in FIG.

  On the other hand, as shown in FIG. 19B and FIG. 20B, the appearance frequency F of the difference data ΔD between the image data D output from the adjacent radiation detection elements 7 connected to the same scanning line 5. The distribution of is not a normal distribution but a distribution that can also be called a trapezoid.

  According to the study by the present inventors, the main cause of the distribution of the appearance frequency F of the difference data ΔD as shown in FIGS. 19B and 20B is shown in FIG. As described above, in the above case, the image data D is read from the radiation detection elements 7 by the different readout circuits 17, but the output characteristics of the readout circuits 17 are different for each readout circuit 17, and the output of each readout circuit 17 is different. This is thought to be due to variations in characteristics.

  Then, the distribution of the output characteristics of each readout circuit 17 and the distribution of the normal distribution of the difference data ΔD itself are superimposed, so that the distribution of the appearance frequency F of the difference data ΔD has a substantially trapezoidal shape, and FIG. If the width of the distribution is wider than the case of the difference data ΔD between the image data D output from the adjacent radiation detection elements 7 connected to the same signal line 6 shown in FIG. Conceivable.

The distribution of the appearance frequency F of the image data D read from each radiation detection element 7 is, for example, as shown in FIG. 21, and the width of the distribution of the appearance frequency F is wider than the difference data ΔD. In the case of performing compression processing, it is known that the compression efficiency is improved when the distribution width of the appearance frequency F is narrow.
Therefore, in the present embodiment, the difference data ΔD in which the distribution width of the appearance frequency F is narrower than that in the case of the image data D is used, and further connected to the same signal line 6 in which the distribution width becomes narrower as described above. It is assumed that difference data ΔD between image data D output from adjacent radiation detection elements 7 is used. Note that the distribution width of the appearance frequency F is narrower when the difference data ΔD between the image data D output from the adjacent radiation detection elements 7 connected to the same scanning line 5 is used than when the image data D is used. Since the compression efficiency is improved, the present invention can be applied.

In this embodiment, only one Huffman code Hc table is prepared in the memory of the ROM 23b, and the compression processing unit 49a refers to this table and performs the compression process regardless of the type of shooting. As will be described, a plurality of tables may be provided for each imaging condition including the imaging region of the patient's body, which is the subject, as the Huffman code Hc table. In this case, for example, when information such as imaging conditions is input to the radiographic imaging apparatus 1 by an operator such as a radiographer, the compression processing unit 49a is input from the table stored in the memory. A table to be used is selected based on the shooting conditions, and the difference data ΔD is compressed with reference to the selected table.
It is also possible to transmit the information of the table to be used from the external device to the radiographic image capturing device 1, and the compression processing unit 49a may select the table accordingly. Further, the Huffman code Hc table suitable for the radiographic imaging conditions is transmitted from the external device to the radiographic imaging device 1 for storage or rewriting every time radiographic imaging is performed, and the compression processing unit 49a is transmitted. It is also possible to perform a compression process on the difference data ΔD with reference to the table of the Huffman code Hc.

On the other hand, since the Huffman code Hc is not prepared for the appearance frequency F below the predetermined threshold as described above, the compression processing unit 49a calculates the difference in which the calculated difference data ΔD is associated with the Huffman code Hc. If it is not within the range of the data ΔD, that is, within the target range not less than the lower limit value X and not more than the upper limit value Y (see FIG. 11), the compression process is not performed (that is, uncompressed). Special code processing for adding a special code to image data (that is, raw signal value, pixel value) is performed.
As described above, the numerical range of the difference data ΔD that the output abnormality determination unit 491c determines that the radiation detection element 7 is abnormal in output matches the range in which the compression processing unit 49a does not perform compression. For the radiation detection element 7 determined to have an output abnormality in 491c, the compression data 49D is not compressed in the compression processing unit 49a, and a special code is added to the image data.

  When compression processing is not performed, a special code is added to differential data ΔD that is not compressed (that is, uncompressed differential data ΔD, so-called raw differential data ΔD) instead of image data, You may make it transfer to the control part 23 grade | etc.,.

  This special code is a code indicating that 16-bit data below the special code is raw image data D that has not been subjected to compression processing, and the 16-bit data is erroneously interpreted as Huffman code Hc. This is a code to prevent this. The special code has a longer code length than the normal Huffman code Hc, and in this embodiment, the special code is a code having a data amount of 24 bits. For this reason, the pixel subjected to the special code processing is 40-bit data in which a 24-bit special code is added to the 16-bit original image data D which is the main body.

[Information addition processing]
The compression processing unit 49a performs uncompressed image data D to which each Huffman code Hc assigned to each differential data ΔD is added when compression processing is performed, and special code is added when special code processing is performed (that is, The 40-bit data) is sequentially sent to the information adding unit 49b.

  The information adding unit 49b secures an area to be a header portion at the head of the data for each line and also sends data sent from the compression processing unit 49a (that is, uncompressed data to which a Huffman code Hc or a special code is added). The original image data D) is sequentially arranged and functions as delimiter information adding means for adding information on the delimiter position of the data to the end of the data for one line.

  In the present embodiment, the information adding unit 49b adds a delimiter code to the end of the data for one line as delimiter information. This is because when compression processing is performed, the compression ratio varies depending on the type of image, and therefore, the delimiter code cannot be specified unless the delimiter code is added. The delimiter information added to the data by the information adding unit 49b is not limited as long as the delimiter position of the data for one line can be determined, and the type, mode, and the like are not limited.

The information added by the information adding unit 49b is not limited to the delimiter information.
For example, in the present embodiment, as will be described later, the header data added to the head header portion of the data is described as being added by the CPU 23a of the main body control unit 23, but the header data is added in the information addition unit 49b. May be added. The header data includes, for example, data size information of data for one line, but is not limited thereto.

Data to which information such as a delimiter code is added by the information adding unit 49b is sent to the data generating unit 49c, and is formatted as transfer data for each line and then sent to the main body control unit 23. .
The data size information of each transfer data for one line is also sent to the register unit 49d and held in the register unit 49d.

[Header part addition processing]
When the Huffman code Hc to which the delimiter code is added is sent from the data generation unit 49c, the CPU 23a of the main body control unit 23 refers to the register unit 49d of the FPGA 24 and acquires the data size of the data for one line. A process of writing (adding) this to the header portion of the data is performed.

FIG. 22 shows a configuration example of transfer data. In the transfer data, header data is added to the head (header portion) of data for one line, and a delimiter code is added to the end of the data. In addition, as a result of encoding with the Huffman code Hc, the original image data (difference data ΔD before the compression process) has a data of 16 bits per pixel converted to a Huffman code Hc of 6 bits, 7 bits, etc. (FIG. 22). It is simply replaced with “Hc”. Further, a special code indicating that is added to the data of the pixel (radiation detection element 7) that could not be converted into the Huffman code.
In FIG. 22, as described above, the case where the data size of the original image data D per pixel (that is, the raw signal value and the raw pixel value) is 16 bits and the special code is 24 bits is exemplified. However, the size of the original image data D and the size of the special code are not limited to those exemplified here.
The data subjected to the special code processing has a data size that exceeds the original image data D, such as 40 bits. However, the radiation detection element 7 to be subjected to the special code processing has a low appearance frequency and is converted into another Huffman code. By canceling out with the radiation detecting element 7, the data size of one line as a whole, and hence one image as a whole, is generally smaller than the original image data D (see FIGS. 21 and 22).

Then, the CPU 23a temporarily stores and saves these data in the storage means 40 for each line in a state where the data can be identified on which line, and appropriately stores the external device via the antenna device 39. (Sequentially transferred to the console 58, for example).
In this embodiment, since data for each line is stored in the storage unit 40 in this way, even if an abnormality occurs in the data transfer process, or a problem occurs in the data at the transfer destination, one line is stored. It becomes possible to respond to a retransmission request every time.

[Radiation imaging system]
Next, the configuration of the radiographic image capturing system in the present embodiment will be described. The radiographic imaging system is a system that assumes radiographic imaging performed in, for example, a hospital or a clinic, and can be employed as a system that captures a medical diagnostic image as a radiographic image, but is not necessarily limited thereto. .

  FIG. 23 is a diagram showing an overall configuration of the radiographic image capturing system in the present embodiment. As shown in FIG. 23, the radiographic imaging system 50 includes, for example, an imaging room R1 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 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.

  FIG. 23 shows a case where the portable radiographic image capturing device 1 is used by being loaded into the cassette holding portion 51a of the bucky device 51. However, the radiographic image capturing device 1 includes a bucky device 51, a support base, and the like. It may be integrally formed. Further, as shown in FIG. 23, the radiographic imaging apparatus 1 and the base station 54 may 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. The cradle 55 may be configured to 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.

  The configuration of the radiographic image capturing apparatus 1 is as described above. 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. It can also be used in a single state.

  That is, the radiation image capturing apparatus 1 is arranged in a single state, for example, on the upper surface side of 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 configured to be provided in the front room R2. 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 above-described base station 54, console 57, storage means 59 composed of a hard disk or the like, and a cradle 55 or the like is connected 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 is notified of the cassette ID of the radiographic imaging apparatus 1 from the cradle 55 via the base station 54, the console 58 saves it in the storage means 59, and the radiographic imaging apparatus 1 existing in the imaging room R1. To manage. In the console 58, each image data D is restored from the Huffman code by the method described in the present embodiment, and image processing such as offset correction and gain correction is performed on the restored image data D to obtain a final result. Image data is generated.

Further, in the memory such as a ROM of the computer constituting the console 58 and the storage means 59, the same information as the data compression processing information such as a Huffman code table stored in the memory of the control means 22 of the radiographic image capturing apparatus 1 is stored. It is remembered.
When the image data subjected to the compression process is transferred, the CPU of the console 58 converts the image data into the original image based on information relating to the data compression process such as a Huffman code table stored in the ROM or the like. It functions as a restoration means for decompressing and restoring data.

[Image data compression processing of radiation imaging equipment]
Next, image data compression processing by the control means 22 of the radiation image capturing apparatus 1 will be described with reference to the flowchart shown in FIG. Here, the processing for the image data of the radiation detection elements 7 for one line connected to the same scanning line 5 is illustrated.
First, the image data of the radiation detecting elements 7 of the line Ln and the line Ln + 1 in which the scanning lines 5 are adjacent to each other are accumulated in the buffer registers 491a and 491b, respectively, and the difference ΔD is obtained as difference data ΔD (n + 1, 1), ΔD. (N + 1, 2), ΔD (n + 1, 3)... (Step S1). When the image data compression process is started from the first line L1, the difference calculation unit 491 reads the reference data Dc (1), Dc (2), Dc (3),... From a memory such as a ROM. Are stored in the buffer register 491a, and the image data D (1,1), D (1,2), D (1,3)... Of the line L1 are stored in the buffer register 491b.

Based on the calculated difference data ΔD, the output abnormality determination unit 491c determines the presence or absence of output abnormality for each radiation detection element 7.
Specifically, the lower limit value and the upper limit value (X and Y values in FIG. 11) of the difference data prepared in the data memory 46 are read, and if the difference data ΔD is not less than the lower limit value and not more than the upper limit value, If the output is normal and the output is out of the range, it is determined that the output is abnormal (step S2).

When the output abnormality determination unit 491c determines that the element output is within the normal range, the image data D (n + 1, 1), D (n + 1, 2), D (n + 1), which are data in the buffer register 491b. , 3)... Are transferred to the buffer register 491a, and the image data D (n + 2, 1), D (n + 2, 2) arranged in the scanning line direction of the line Ln + 2 of the scanning line 5 next adjacent to the buffer register 491b which has become empty. ), D (n + 2, 3),... Are accumulated (step S3).
Each image data newly accumulated in the buffer registers 491a and 491b is used for calculating the difference data ΔD of each radiation detecting element 7 on the line along the next scanning line 5.

On the other hand, when there is a radiation detection element 7 whose element output is determined to be outside the normal range in the output abnormality determination unit 491c, the image data D of the radiation detection element 7 is not transferred from the buffer register 491b to the buffer register 491a. (For example, refer to FIG. 18A), all the image data D of the other radiation detection elements 7 are moved (step S4).
As a result, the image data of the previous line Ln remains at the address of the buffer register 491a where the image data of the radiation detection element 7 determined to be abnormal in output should be stored. Therefore, when the difference data ΔD for the next line Ln + 2 is calculated, the image data of the radiation detection element 7 determined to be abnormal in output is not reflected and is excluded from the difference calculation processing target.

Next, the encoding processing unit 492 determines whether or not the difference data ΔD generated by the difference calculation unit 491 is a compression target (step S5).
Specifically, the encoding processing unit 492 determines whether or not the difference data ΔD is in the Huffman code Hc table.
However, since it takes time to determine whether the difference data ΔD is in the Huffman code Hc table with reference to the Huffman code Hc table, the encoding processing unit 492 is prepared in the data memory 46, for example. Whether or not the compression processing is performed depending on whether or not each difference data ΔD falls within the range determined by the lower and upper limit values of the X and Y values in FIG. 11, that is, the data for which the Huffman code Hc table is prepared. It is preferable to determine whether or not.
Further, in this radiographic imaging device 1, the output abnormality determination unit 491 c in the process of step S <b> 2 also determines the output abnormality of each radiation detection element 7, and the difference data ΔD of each radiation detection element 7 has the same X and Y numerical values. Since the determination is based on whether or not the difference data is within the range, the determination whether or not the difference data ΔD is within the range of X to Y is omitted in step S4, and the compression process is performed using the determination result in step S2. It may be determined whether or not. Needless to say, when the determination of output abnormality and the determination of whether or not to perform compression processing are performed in different numerical ranges, the determination must be made individually as described above.

If the difference data ΔD is to be subjected to compression processing, it is replaced with the Huffman code Hc corresponding to the value of the difference data ΔD (step S6).
On the other hand, if the difference data ΔD is not subject to compression processing, special code processing is performed for adding the special code to the original image data of the difference data ΔD without compressing the difference data ΔD (step S7). .

When the encoding processing unit 492 performs compression processing or special code processing based on each difference data ΔD, the information adding unit 49b adds a delimiter code to the end of the data for one line (step S8).
Then, the data generation unit 49c generates compressed data transfer data (step S9), and the transfer data is transferred to the main body control unit 23 (step S10).
Information such as the data size of the transfer data is stored in the register unit 49d, and the main body control unit 23 obtains the data size information of the transferred transfer data from the register unit 49d and adds a header to the header portion of the transfer data. It is added as data (step S11).
The main body control unit 23 stores the transfer data in the storage unit 40 and transfers the data to an external device such as the console 58 via the antenna device 39 as appropriate.
Note that the order of the process of adding a delimiter code (step S9), the process of generating transfer data (step S10), and the process of adding header data to the beginning of the data (step S11) is not limited to that described here. . For example, a delimiter code and header data may be added to the data by the information adding unit 49b and then transferred to the main body control unit 23.

Next, a data restoration process performed by the console 58 that acquired the transfer data transferred from the radiographic imaging apparatus 1 will be described.
In the following, it is also assumed that the difference data ΔD between the image data D output from the adjacent radiation detection elements 7 connected to the same signal line 6 is compressed and transferred from the radiation image capturing apparatus 1. As will be described, the compressed difference data ΔD transferred from the radiation image capturing apparatus 1 is the difference data ΔD between the image data D output from the adjacent radiation detection elements 7 connected to the same scanning line 5. The case is explained in the same way.

  Further, as described above, the memory such as the ROM of the computer constituting the console 58 and the storage means 59 include the Huffman code Hc table and the reference data Dc (1) stored in the control means 22 of the radiographic apparatus 1. , Dc (2), Dc (3),... Are stored in the same information. Note that the lossless compression method does not necessarily need to be based on Huffman coding, and can be configured to compress the difference data ΔD (or image data D) using another lossless compression method.

  The console 58 (see FIG. 23) once stores the compressed transfer data from the radiographic image capturing apparatus 1 via the antenna device 39 or the cable via the base station 54 or the like, in the storage means 59. .

  Then, the console 58 reads the transfer data once stored in the storage means 59, extracts the transfer data for one line from the header data and the delimiter code, reads the data for the one line in order, It is determined whether each code matches any code in the Huffman code Hc table, and if it matches, the original difference data ΔD is restored according to the Huffman code Hc table. In addition, when it does not match any of the Huffman codes Hc and matches the special code, the restoration processing of the image data D is performed.

Further, the console 58 also includes two buffer registers 581 and 582. The difference data ΔD or the image data restored according to the line arrangement order is accumulated in each buffer register 581 and 582, and the difference data is obtained by integration processing. Image data D is calculated from ΔD. That is, the console 58 also functions as a restoration unit and an image data calculation unit.
Here, a process of calculating the image data D from the difference data ΔD using the two buffer registers 581 and 582 will be described with reference to FIG. When the difference data ΔD of the first line L1 is calculated, the reference data Dc (1), Dc (2), Dc (3),..., And the other buffer register 582 are stored in one buffer register 581. The difference data ΔD of the line L1 is accumulated. In the example of FIG. 25, the calculation is already performed up to the line Ln, and the image data D is calculated from the difference data ΔD of the line Ln + 1. In this example, calculation processing in a case where uncompressed image data D (n + 1, 2) is included in the difference data ΔD of the line Ln + 1 will be described.

As shown in FIG. 25A, the buffer register 581 has image data D (n, 1), D (n, 2), D (n, 3), D (n, 4),. The buffer register 582 stores restored data ΔD (n + 1,1), D (n + 1,2), ΔD (n + 1,3), ΔD (n + 1,4),.
The restoration data accumulated in the buffer register 582 is discriminated whether the data accumulated in order from the head address is difference data or image data. In the case of difference data ΔD, the data is accumulated at the same address in the buffer register 581. In addition to performing addition processing with data, the data is output to a storage area for image data development prepared in the storage means 59 of the console 58 and stored in order.
For example, as shown in the figure, when the difference data ΔD (n + 1, 1) is stored in the buffer register 582 and the image data D (n, 1) is stored in the buffer register 581, these are added to form an image. Data D (n + 1, 1) is calculated and output to the memory.
Further, the image data D (n, 1) in the buffer register 581 is simultaneously updated to the image data D (n + 1, 1) calculated by addition. As a result, one buffer register 581 is rewritten by the calculation process with the other buffer register 582, so that it is necessary to read the restored data for two lines simultaneously for the two buffer registers 581 and 582. As a result, processing can be speeded up.

  On the other hand, as shown in FIG. 25B, when the data stored in the buffer register 582 is image data (for example, D (n + 1, 2) in the figure), addition processing with the data on the buffer register 581 side is performed. Is not performed and is output to the storage means 59 as it is. In that case, the data at the corresponding address on the buffer register 581 side (for example, D (n, 2) in the figure) is not updated. As a result, the image data of the line Ln remains only in the data of the corresponding address of the buffer register 581 and the other addresses are updated to the image data of the line Ln + 1.

When the calculation process by addition is completed for all the restored data of the line Ln + 1 stored in the buffer register 582, the restored data ΔD (n + 2, 1), ΔD (n + 2, 2) ΔD (n + 2, 3), ΔD (n + 2, 4),... Are accumulated.
As described above, image data calculation processing is performed by adding data having the same address. As shown in FIG. 25C, the difference data ΔD (n + 2, 2) in the buffer register 582 is processed. An addition process is performed on the image data D (n, 2) of the line Ln that has not been updated in the previous calculation process.
This is because the radiation detection device 7 excludes the image data of the radiation detection element 7 from the radiation detection element 7 determined to be abnormal in output when the radiation image capturing apparatus 1 calculates the difference data, and detects the radiation in the direction of the signal line 6. Since the difference data is calculated between the radiation detecting elements 7 on both sides of the element 7 (see FIG. 18), it is necessary to calculate the image data D from the difference data ΔD in consideration of this. It is. That is, the radiation detection element 7 in which the output abnormality is recognized in the radiographic image capturing apparatus 1 is transmitted to the console 58 side with data obtained by adding a special code to the image data instead of the compressed data of the difference data ΔD. When image data is included in the difference data, it can be identified as a mark of the image data of the radiation detection element 7 that has performed an abnormal output. Accordingly, when image data is included in the difference data for one line, an abnormal output is not generated in the next addition process by not updating the address corresponding to the image data in the buffer register 581. The addition of the data of the radiation detecting elements 7 on both sides of the performed radiation detecting element 7 is realized to restore the appropriate image data.

  In some cases, the radiation detection elements 7 that have performed abnormal output are continuously present along the direction of the signal line 6. For example, if the radiation detection element 7 that outputs D (n + 2, 2) of the line Ln + 2 adjacent to D (n + 1, 2) of the line Ln + 1 in the above example also performs abnormal output, the line is input to the buffer register 582. The restoration data ΔD (n + 2, 1), D (n + 2, 2), ΔD (n + 2, 3), ΔD (n + 2, 4),... Of Ln + 2 are accumulated. Then, the same processing as in FIG. 25B is performed, and D (n + 2, 2) accumulated in the buffer register 582 is output to the storage means 59 without being added and accumulated in the buffer register 581. The value of D (n, 2) is maintained without being updated. Therefore, the restoration data ΔD (n + 3, 1), ΔD (n + 3, 2), ΔD (n + 3, 3), ΔD (n + 3, 4),... Is performed, the difference data ΔD (n + 3, 2) in the buffer register 582 is subjected to addition processing with the image data D (n, 2) of the line Ln that has not been updated in the previous calculation process and the previous calculation process. Become.

  In the console 58, the image data for one screen can be developed in the development area of the storage unit 59 by executing the above processing for each line along all the scanning lines 5.

Also, the console 58 is not limited to two buffer registers 581, and may have only one or three or more. An example of processing when only one buffer register 581 is provided will be described below. Further, as in the case of FIG. 25, a case where data is transferred from the radiation image capturing apparatus 1 in a state in which uncompressed image data D (n + 1, 2) is included is illustrated.
The buffer register 581 stores image data D (n, 1), D (n, 2), D (n, 3), D (n, 4),... Of the line Ln, and the adjacent line Ln + 1. Restored data ΔD (n + 1, 1), D (n + 1, 2), ΔD (n + 1, 3), ΔD (n + 1, 4),.
Then, the restoration data of the line Ln + 1 sequentially read from the storage means 59 is identified as difference data or image data in order, and in the case of difference data ΔD, addition processing with data accumulated at the corresponding address of the buffer register 581 is performed. The image data of the line Ln + 1 is calculated and output to the image data development storage area prepared in the storage means 59 of the console 58. At the same time, the image data of the line Ln in the buffer register 581 is converted into the image data of the line Ln + 1. Replace.
On the other hand, when the restoration data of the line Ln + 1 sequentially read from the storage unit 59 is image data (for example, D (n + 1, 2)), the addition process with the data on the buffer register 581 side is not performed, and D (n + 1) is performed. , 2) are output to the storage means 59 as they are. In that case, the data at the corresponding address on the buffer register 581 side (for example, D (n, 2) in the figure) is not updated. As a result, the image data of the line Ln remains only in the data of the corresponding address of the buffer register 581 and the other addresses are updated to the image data of the line Ln + 1.
From then on, the console 58 executes the same processing for each line along all the scanning lines 5, thereby developing one screen of image data in the development area of the storage means 59 by one buffer register 581. Is possible.

[Effect of the embodiment of the invention]
As described above, in the radiographic image capturing apparatus 1 shown in the present embodiment, the difference data ΔD is obtained from the image data D of the radiation detection elements 7 adjacent to each other, and the difference is obtained using the Huffman code Hc table that is a static dictionary. Data ΔD is compressed. In the plurality of radiation detection elements 7 arranged in a matrix, since the two adjacent radiation detection elements 7 are close to each other, it is unlikely that the doser incident state incident on them is greatly different. Therefore, the difference data between two adjacent radiation detection elements 7 occupies most of the data indicating 0 or a value close thereto (see FIG. 11). On the other hand, when the Huffman compression process is performed, it is possible to compress more efficiently when the variation in data is small and the appearance frequency F of the difference data ΔD having a common value increases. That is, by performing the Huffman compression process on the difference data ΔD between the adjacent radiation detection elements 7, the compression can be performed with high efficiency, the data transmission can be performed more quickly, and the radiation image capturing apparatus 1. In addition, power saving can be achieved.

  Further, since the output abnormality determination unit 491c can detect the radiation detection element 7 that is abnormal in output based on the difference data ΔD, the radiation detection element that performs abnormal output from the beginning of manufacture and the abnormal output later. A radiation detection element, a radiation detection element that occasionally outputs abnormal radiation, a radiation detection element that occasionally happens to output abnormal radiation images, etc. can all be identified after imaging, and indicate the location of the defective radiation detection element 7 in advance. There is no need to create a map, and maintenance for updating it can be eliminated.

Further, the encoding processing unit 492 includes any one of the radiation detection elements 7 (for example, an element of the line Ln + 1) and another radiation detection element 7 (for example, an element of the line Ln) adjacent on one side in the signal line direction. When the output abnormality is determined based on the difference data ΔD, the radiation detection element 7 (for example, the element of the line Ln + 1) determined to be the output abnormality and the radiation detection element 7 adjacent to the opposite side (for example, the element of the line Ln + 2) The difference data with) is not subject to compression processing. The difference calculation unit 491 also calculates difference data between the radiation detection element 7 (for example, the element of the line Ln + 1) determined to be abnormal in output and the radiation detection element 7 (for example, the element of the line Ln + 2) adjacent to the opposite side. Do not do.
When the image data of a certain radiation detection element 7 has a value that causes an output abnormality, the two difference data ΔD between the respective elements before and after that have a large absolute value, and the range in which the distribution of the appearance frequency F is concentrated However, the number of such differential data ΔD can be at least halved by the above processing, and the efficiency of compression of the entire transferred data can be improved.

In the above embodiment, the abnormality output determination unit 491c is provided in the difference calculation unit 491, and the output abnormality determination of the radiation detection element 7 is determined depending on whether or not the difference value indicated by the difference data ΔD is within a predetermined range (X to Y). However, the present invention is not limited to this. For example, an abnormal output determination unit is provided in the encoding processing unit 492, and not the difference value indicated by the difference data ΔD but the number of bits at the time of transfer determined depending on whether or not the difference data ΔD is compressed (when compressed) The number of bits at the time of compression, the number of bits of the image data + the number of bits of the special code when not compressed), the output abnormality of the radiation detection element 7 is determined depending on whether or not it falls within a predetermined threshold. Also good.
In this case, when the difference data ΔD between the radiation detection elements 7 between a certain first line and the second line along the scanning line 5 direction is obtained, the difference between the second line and the third line is obtained. Before calculating the difference between the radiation detection elements 7, it is necessary to quickly determine the number of bits of the transfer data.

Further, in the radiographic imaging device 1, when the output abnormality determination unit 491c determines that the radiation detection element 7 is abnormal in output, the radiation detection element 7 located in the next line is detected as an adjacent output abnormality. Difference data ΔD with respect to the radiation detecting element 7 adjacent to the other side instead of the element 7 is obtained, and this is set as a compression processing target. As a result, the difference value of the difference data ΔD can be reduced, and the number of individuals of the difference data ΔD that has a value greatly deviating from the range in which the distribution of the appearance frequency F by the radiation detection element 7 with abnormal output can be concentrated can be reduced. It becomes.
In the present embodiment, for the radiation detection element 7 on the line next to the radiation detection element 7 determined to be abnormal in output, the radiation detection element 7 on the line immediately before the radiation detection element 7 determined to be abnormal in output. The difference value of the difference data ΔD can be brought close to 0, and this is the most desirable, but other radiation detection around the radiation detection element 7 determined to be abnormal in output. You may obtain | require difference data (DELTA) D with the element 7. FIG. In that case, when the console 58 calculates the image data D from the difference data ΔD, it is necessary to perform processing so as to add the difference data ΔD to the image data of the radiation detection element 7 having an appropriate positional relationship.

  Moreover, in the radiographic imaging device 1, since the difference calculation unit 491 calculates the difference data ΔD between the radiation detection elements 7 adjacent along the extending direction of the signal line 6, more difference data ΔD is included within the narrower difference value range. The appearance frequency F can be concentrated, and the compression efficiency can be further increased.

  Further, in the radiographic imaging device 1, the output abnormality determining unit 491c determines that the output abnormality of the radiation detecting element is abnormal when the difference data ΔD exceeds the range determined by the predetermined threshold values X and Y. By determining, it becomes possible to more appropriately determine whether there is an abnormality.

  In addition, when the encoding processing unit 492 determines that the difference data ΔD is not within the range determined by the predetermined threshold value (X, Y), the encoding processing unit 492 does not perform the compression processing and performs the radiation detection element 7 corresponding to the difference data ΔD. A special code indicating that the image data is not compressed is added. For this reason, the difference data ΔD deviating from the value near the difference value 0 where the difference data ΔD is concentrated can be excluded from the compression processing targets. For this reason, the number of codes used in the Huffman code can be reduced, and highly efficient compression can be performed.

Since the difference data ΔD indicating the tendency of appearance frequency to concentrate within a narrow difference value range as in the present embodiment is the target of compression processing, a shorter code is assigned to each possible difference value to perform compression processing. By performing compression processing using the Huffman compression method, more effective data compression can be performed.
A table of Huffman codes Hc, which is a static dictionary, is prepared in advance by storing a plurality of tables for each imaging condition including an imaging region of a patient's body, which is a subject, and according to a specified imaging region or imaging condition. By making this selectable, data compression can be realized more effectively.

  Needless to say, the present invention is not limited to the above-described embodiment, and can be appropriately changed without departing from the gist of the present invention.

DESCRIPTION OF SYMBOLS 1 Radiation imaging device 5 Scanning line 6 Signal line 7 Radiation detection element (pixel)
17 Reading circuit 22 Control means 23 Main body control unit 24 FPGA
39 Antenna device (transfer means)
40 storage means 44 difference calculation unit (calculation means)
45 Buffer memory 46 Data memory 49 Transfer data generation unit 491 Difference calculation unit (calculation means)
491c Output abnormality determination unit (output abnormality determination means)
492 Encoding processing unit 49a Compression processing unit (compression means)
49b Information adding section 49c Data generating section 49d Register section 50 Radiation imaging system 58 Console (restoring means, image data calculating means)
59 Storage means D Image data Hc Huffman code P Detection part r Region D Image data ΔD Difference ΔD Difference data

Claims (8)

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 readout circuit that reads out charges from the radiation detection elements through the signal lines, converts the charges into electrical signals for each of the radiation detection elements, and outputs them as image data;
Calculating means for calculating difference between the image data output from the radiation detection elements adjacent to each other in a certain arrangement direction in the detection unit, and creating difference data;
Compression means for compressing the difference data using a static dictionary;
A transfer means for transferring data;
For each of the radiation detection elements, an output for determining an output abnormality based on the difference data calculated from the difference in image data with the radiation detection element adjacent to one side in the fixed arrangement direction or a transfer data amount based on the difference data An abnormality determination means,
The compression means includes
When an output abnormality is determined for any radiation detection element by the output abnormality determination means, it is determined that the output abnormality is the radiation detection element adjacent to the radiation detection element adjacent to the opposite side in the certain arrangement direction. The difference data between the radiation detection element adjacent to the opposite side and the other radiation detection elements around the radiation detection element determined to be abnormal in output is not subjected to the difference data with the radiation detection element. A radiographic imaging device characterized by the above.   The calculation means, for the radiation detection element adjacent to the opposite side, performs a process of calculating difference data from the radiation detection element adjacent to the one side of the radiation detection element determined to be abnormal in output. The radiographic imaging apparatus according to claim 1.   The radiographic image capturing apparatus according to claim 1, wherein the certain arrangement direction is an extending direction of the signal lines.   The output abnormality determining means determines that the output abnormality of the radiation detecting element is abnormal when the difference amount of the difference data or the data amount of the compression processing data of the difference data exceeds a range determined by a predetermined threshold value. The radiographic imaging apparatus as described in any one of Claim 1 to 3. The compression means does not perform compression processing on differential data between the radiation detection element determined to be abnormal in output and the radiation detection element adjacent on the one side, and the radiation detection element determined to be the differential data or the output abnormality In addition to adding a special code indicating that compression processing has not been performed on the image data of
5. The transfer unit according to claim 4, wherein the transfer unit transfers the data to which the special code is added instead of the differential data subjected to the compression process for the radiation detection element determined to have the output abnormality. Radiation imaging device.   6. The compression unit according to claim 1, wherein the compression unit performs compression processing of each differential data by performing Huffman coding of the differential data using a Huffman code table as a static dictionary. A radiographic imaging apparatus according to claim 1. A plurality of the static dictionaries are stored in advance for each imaging condition including an imaging region of a patient's body as a subject,
The compression unit selects the static dictionary according to the designated shooting condition when performing the compression process, and performs the compression process of the difference data based on the selected static dictionary. The radiographic imaging device according to any one of claims 1 to 6. The radiographic imaging device according to any one of claims 1 to 7,
A console for receiving data including compressed difference data transferred from the radiographic imaging device;
With
The console is
Restoring means for restoring the compressed difference data to the original difference data based on a static dictionary;
Image data calculation means for calculating image data from the restored difference data,
For the image data of the radiation detection element adjacent to the opposite side to the radiation detection element that is abnormal in output, the image data calculation means includes a radiation detection element around the radiation detection element that is abnormal in output. The radiation image capturing system is calculated based on the difference data and the image data of the radiation detection element.

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