JP2019037550A - X-ray CT apparatus and image generation method - Google Patents

Abstract

To accurately acquire projection data, consequently a CT value, in an X-ray CT apparatus employing a detector of a photon counting system.SOLUTION: An X-ray CT apparatus is provided which includes: a detector detecting X-ray photons emitted from an X-ray source circulating around a subject on a bed and transmitted through the subject; a data collection unit for outputting a counted value every energy range by collecting and processing the X-ray photons detected by the detector, concerning predetermined multiple energy ranges; data addition unit for multiplying weight determined every energy range to the counted value outputted by the data collection unit, and being proportional to the medium value of the energy range, and performing addition processing; and an image generation unit for generating an image by performing reconstruction processing using the data added by the data addition unit.SELECTED DRAWING: Figure 3

Description

  The present invention relates to an X-ray CT apparatus (Photo Counting Computed Tomography) equipped with a photon counting detector that acquires energy information of incident X-rays by counting incident X-ray photons. Hereinafter, it is referred to as a PCCT apparatus) and an image generation method.

  An X-ray CT apparatus obtains X-ray transmission data (projection data) of a subject while rotating a pair of an X-ray source and an X-ray detector arranged opposite to each other with the subject interposed therebetween, and obtains a tomographic image (hereinafter referred to as a CT image). Is used for industrial and security inspection devices, medical image diagnostic devices, and the like.

  In such an X-ray CT apparatus, a semiconductor detector, a scintillator detector, or the like is applied to generate a charge amount or light proportional to the energy of the X-ray incident on the detector, and obtained by detecting this. A tomographic image is acquired by reconstructing the projection data. As a more specific example, in a scintillator detector, when, for example, 20 keV X-rays are incident on the scintillator detector, a photon corresponding to 20 keV, that is, proportional to 20 keV energy is generated in the detector, which is converted into a photoelectron. It is converted into an electrical signal by a multiplier and output. Therefore, in the X-ray CT apparatus, a signal value corresponding to the energy of the incident X-ray, that is, a CT value can be obtained.

  Incidentally, in recent years, PCCT apparatuses, which are X-ray CT apparatuses equipped with a photon counting type detector, are known (for example, Patent Document 1 and Patent Document 2). In the PCCT apparatus, a photon counting type detector counts (counts) X-ray photons (X-ray photons) irradiated from an X-ray source and transmitted through a subject for each detection element, and obtains a count value obtained thereby. Based on this, an X-ray CT image is reconstructed. In the PCCT apparatus, for example, a spectrum capable of estimating the elements constituting the internal tissue of the subject through which X-rays are transmitted can be acquired, and an X-ray CT image in which differences in element levels are depicted in detail can be obtained. Further, in the PCCT apparatus, only X-rays in the energy range near the k-edge existing for each substance can be extracted and imaged, and an image of only the element having the k-edge can be created and used for diagnosis.

JP-A-2015-198833 International Publication 2012/029496

As described above, an X-ray CT apparatus to which a scintillator detector, a semiconductor detector, or the like is applied can obtain a signal value corresponding to the energy of incident X-rays, that is, a CT value.
However, the PCCT apparatus can only obtain the number of counts within a preset energy range, and cannot obtain detailed energy information for each count value. For example, when the energy range is set to 40 to 60 keV and 5 counts are obtained within a predetermined time in the range, information on how much each of those energies is not obtained. Therefore, since the PCCT apparatus cannot obtain information corresponding to energy, the CT value obtained by a general X-ray CT apparatus to which a scintillator detector or the like is applied cannot be obtained.
For this reason, in order to make use of the knowledge in the examination by the conventional X-ray CT apparatus, it is required that the PCCT apparatus also obtain a CT value equivalent to the conventional one.

  Here, although it is conceivable to obtain a CT value from a virtual monochromatic image or the like, the attenuation curve of the corresponding substance needs to be expressed by a combination of the base substances selected at the time of decomposition into base substances. As a result, even if a plurality of base substances are selected, an error occurs if the corresponding substances cannot be expressed by their combination, and the CT value of the substance cannot be obtained with high accuracy. Therefore, PCCT apparatuses are also required to obtain projection data in order to obtain accurate CT values.

  The present invention has been made in view of the above circumstances, and an object of the present invention is to obtain projection data and CT values with high accuracy in an X-ray CT apparatus to which a photon counting detector is applied.

In order to solve the above problems, the present invention provides the following means.
The present invention relates to a detector that detects X-ray photons emitted from an X-ray source that circulates around a subject on a bed and transmitted through the subject, and collects and processes the X-ray photons detected by the detector. The data collection unit that outputs a count value for each energy range for a plurality of predetermined energy ranges, and the energy range that is determined for each energy range with respect to the count value that is output by the data collection unit. A data addition unit that performs addition processing by multiplying a weight proportional to the median of the image data, and an image generation unit that generates an image by performing reconstruction processing using the data added by the data addition unit. An X-ray CT apparatus is provided.

  According to the present invention, in an X-ray CT apparatus to which a photon counting type detector is applied, it is possible to acquire projection data and, in turn, CT values with high accuracy.

1 is a block diagram showing a schematic configuration of an X-ray CT apparatus according to a first embodiment of the present invention. It is a block diagram which shows a part of detector of the X-ray CT apparatus which concerns on the 1st Embodiment of this invention. It is a block diagram which shows schematic structure of the calculating part of the X-ray CT apparatus which concerns on the 1st Embodiment of this invention. It is a flowchart which shows the flow of the imaging process in the X-ray CT apparatus which concerns on the 1st Embodiment of this invention. (A) is a graph which shows the attenuation coefficient of water, (b) shows the amount detected when the amount of input X-rays in all energy ranges is set to 1 when 100 mm of water is transmitted. It is a graph. It is a flowchart which shows the flow of the imaging process in the X-ray CT apparatus which concerns on the 2nd Embodiment of this invention. It is a graph which shows the energy distribution of an X-ray tube, and shows the peak which especially depends on a setting tube voltage, and the some peak resulting from the target material of an X-ray tube.

  An X-ray CT apparatus according to an embodiment of the present invention includes a detector that detects X-ray photons emitted from an X-ray source that circulates around a subject on a bed and transmitted through the subject, and an X-ray detected by the detector. By collecting and processing photons, for each of a plurality of predetermined energy ranges, a data collection unit that outputs a count value for each energy range, and a count value output by the data collection unit for each energy range. A data addition unit that performs addition processing by multiplying a weight proportional to the median of the energy range, an image generation unit that generates an image by performing reconstruction processing using the data added by the data addition unit, It has.

Hereinafter, embodiments of the present invention will be described in more detail with reference to the drawings.
(First embodiment)
The X-ray CT apparatus according to this embodiment is a photon counting CT apparatus (PCCT apparatus) provided with a photon counting type detector, and photons (X-ray photons) derived from X-rays transmitted through a subject are detected in the detector. Count. Individual X-ray photons have different energies, and the detector discriminates and counts X-ray photons for each predetermined energy band. Thereby, the number of X-ray photons for each energy band, that is, the X-ray intensity is obtained.

As shown in FIG. 1, the PCCT apparatus 100 according to the present embodiment includes a UI unit 200, a measurement unit 300, and a calculation unit 400.
The UI unit 200 includes an input unit 210 such as a keyboard and a mouse, and an output unit 220 such as a display unit (monitor) or a printer. The UI unit 200 receives input from the user and presents the processing result of the calculation unit 400 to the user. . A liquid crystal display, a CRT (Cathode Ray Tube), or the like can be applied as the display unit, and the display unit can have a touch panel function and function as the input device 210.

  The measurement unit 300 irradiates the subject 101 with X-rays and measures X-ray photons transmitted through the subject 101 according to control by the CPU 401 of the calculation unit 400 described later. The measurement unit 300 includes an X-ray irradiation unit 310, an X-ray detection unit 320, a gantry 330, a control unit 340, a data collection unit 404, and a bed 102 on which the subject 101 is placed.

  In the center of the gantry 330, a circular opening 331 for placing the subject 101 and the bed 102 on which the subject 101 is placed is provided. Inside the gantry 330, a rotating plate 332 on which an X-ray tube 311 and an X-ray detector 321 described later are mounted, and a driving mechanism for rotating the rotating plate 332 are arranged. In the following description, the circumferential direction of the opening 331 is the x direction, the radial direction is the y direction, and the direction orthogonal to them is the z direction. In general, the z direction is the body axis direction of the subject 101.

  During normal imaging, the rotating plate 332 rotates and the bed 102 moves in the z direction so that X-rays are emitted from various directions of the subject 101 and X-rays transmitted through the subject 101 are emitted by the X-ray detector 321. Detect and get data. On the other hand, at the time of scanogram imaging, the rotating plate 332 does not rotate and only the bed 102 is moved once in the z direction to finish, or after the imaging, the X-ray irradiation direction is changed (for example, the 90-degree rotating plate 332 is changed). The image is taken after being rotated a predetermined number of times, and the data acquisition of irradiation data from a certain direction is performed a plurality of times at different angles to obtain data.

  The time required for the rotation of the rotating plate 332 depends on parameters input by the user via the UI unit 200. In this embodiment, for example, the time required for rotation is 1.0 s / time. The number of times of imaging in one rotation by the measurement unit 300 is 900, for example, and one imaging is performed every time the rotating plate 332 rotates 0.4 degrees. Each specification is not limited to these values and can be variously changed according to the configuration of the PCCT apparatus 100.

  The X-ray irradiation unit 310 generates X-rays and irradiates the subject 101 with the generated X-rays. The X-ray irradiation unit 310 includes an X-ray tube 311, an X-ray filter 312, and a bowtie filter 313.

  The X-ray tube 311 irradiates the subject 101 with an X-ray beam by a high voltage supplied in accordance with the control of an irradiation controller 341 described later. The irradiated X-ray beam spreads with a fan angle and a cone angle. The X-ray beam is applied to the subject 101 as the rotating plate 332 of the gantry 330 described later rotates.

  The X-ray filter 312 adjusts the X-ray dose of X-rays emitted from the X-ray tube 311. That is, the X-ray spectrum is changed. The X-ray filter 312 of this embodiment attenuates the X-rays irradiated from the X-ray tube 311 so that the X-rays irradiated from the X-ray tube 311 to the subject 101 have a predetermined energy distribution. The X-ray filter 312 is used to optimize the exposure amount of the patient who is the subject 101. For this reason, it is designed so that the dose in the required energy band is increased.

  The bow tie filter 313 suppresses the exposure amount of the peripheral part. It is used for optimizing the exposure dose by using the elliptical cross section of the human body as the subject 101, increasing the dose near the center, and lowering the ambient dose.

Each time an X-ray photon enters, the X-ray detection unit 320 outputs a signal capable of measuring the energy value of the X-ray photon. The X-ray detection unit 320 includes an X-ray detector 321.
FIG. 2 illustrates a part of the X-ray detector 321. The X-ray detector 321 according to the present embodiment includes a plurality of detection elements 322, a counting circuit 324, and a collimator 323 that limits the incident direction to the X-ray detector 321.
In order to facilitate the production, a plurality of planar detectors (detector modules) are created, arranged so that the central portion of the plane is an arc, and are arranged in a pseudo arc shape to obtain an X-ray detector. It is good also as 321.

  The X-rays incident on each detection element 322 are converted into one pulse electrical signal (analog signal) by the counting circuit 324 every time one X-ray photon enters. The converted electrical signal is input to the data collection unit 404.

  As the detection element 322, for example, a CdTe cadmium telluride semiconductor element that directly converts incident X-ray photons into an electric signal is used. The detection element 322 may be a scintillator that emits fluorescence upon receiving X-rays and a photodiode that converts fluorescence into electricity.

The number of detection elements 322 (number of channels) of the X-ray detector 321 is, for example, 1000. The size of each detection element in the x direction is, for example, 1 mm.
For example, the distance between the X-ray generation point of the X-ray tube 311 and the X-ray incident surface of the X-ray detector 321 is 1000 mm. The diameter of the opening 331 of the gantry 330 is 700 mm.
Like the gantry, each specification is not limited to these values and can be variously changed according to the configuration of the PCCT apparatus 100.

  The control unit 340 includes an irradiation controller 341 that controls irradiation of X-rays from the X-ray tube 311, a gantry controller 342 that controls rotation driving of the rotating plate 332, a bed controller 343 that controls driving of the bed 102, and X A detection controller 344 that controls X-ray detection in the line detector 321 is provided. Each of these units operates according to control by the measurement control unit 420 of the calculation unit 400 described later.

  The calculation unit 400 controls the overall operation of the PCCT apparatus 100, collects data acquired by the measurement unit 300, and processes the image to perform imaging. As shown in FIGS. 1 and 3, the calculation unit 400 includes a central processing unit (hereinafter referred to as “CPU”) 401 that performs various types of processing of data obtained by causing the measurement unit 300 to perform imaging, and at the time of processing. And an HDD (Hard disk drive) device 403 that stores in advance necessary programs and data used for processing, data generated during processing, data obtained as a result of processing, and the like. Note that the processing result by the calculation unit 400 is also output to the output device 220 of the UI unit 200.

  As shown in FIG. 3, the CPU 401 includes an imaging condition setting unit 410, a measurement control unit 420, a data addition unit 430, a correction unit 440, and an image generation unit 450. Automatic exposure control is performed to acquire a positioning image and determine an X-ray irradiation amount based on the positioning image.

  The functions of each unit included in the CPU 401 can be realized as software by the CPU 401 reading and executing a program stored in advance in a storage device such as the HDD device 403. In addition, part or all of the operations executed by each unit included in the CPU 401 can be realized by hardware such as an application specific integrated circuit (ASIC) or a field-programmable gate array (FPGA).

The imaging condition setting unit 410 sets imaging conditions input from the user via the UI 200. For example, the imaging condition setting unit 410 displays a reception screen for receiving imaging conditions on the display device 220, and receives and sets an input of imaging conditions by the user via the reception screen.
The imaging conditions include, for example, the tube current and tube voltage of the X-ray tube 311, the imaging range of the subject 101, the shape of the X-ray filter 312, the shape of the bow tie filter 313, and the resolution. The imaging conditions do not necessarily have to be input by the user every time. For example, typical imaging conditions may be stored in advance in the HDD device 403 and the CPU 401 may read out and use them.

The measurement control unit 420 controls the control unit 340 according to the imaging conditions set by the imaging condition setting unit 410 to perform measurement.
Hereinafter, a CT image capturing method performed in accordance with the control of the measurement control unit 420 will be specifically described.
The measurement control unit 420 causes the bed controller 343 to move the bed 102 in a direction perpendicular to the rotating plate 332 and stops moving when the imaging position of the rotating plate 332 matches the designated imaging position. To instruct. Thereby, the arrangement of the subject 101 is completed.

  The measurement control unit 420 instructs the gantry controller 342 to operate the drive motor and instructs the rotating plate 332 to start rotation at the same timing as the instruction to the bed controller 343.

  In response to this instruction, the rotation plate 332 starts rotating, the rotation of the rotation plate 332 becomes a constant speed state, and when the placement of the subject 101 is completed, the measurement control unit 420 instructs the X-ray tube 311 to the irradiation controller 341. In addition to instructing the irradiation timing, the imaging controller 344 is instructed of the imaging timing of the X-ray detector 321. Thereby, the measurement control unit 420 performs X-ray irradiation and X-ray photon detection by the detector, and starts measurement. The measurement control unit 420 repeats these instructions a predetermined number of times to measure the entire imaging range by the measurement unit 300. In addition, you may control to image, moving the bed 102 like the well-known helical scan (Helical Scan).

  Next, a scanogram imaging method will be described. The measurement control unit 420 moves the bed 102 in a direction perpendicular to the rotating plate 332 with respect to the bed controller 343 and stops the movement when the imaging position of the rotating plate 332 matches the designated imaging position. To instruct. Thereby, the arrangement of the subject 101 is completed.

  Next, the measurement control unit 420 issues only a movement instruction to the bed controller 343 and does not instruct the gantry controller 342. Therefore, the rotating plate 332 does not rotate and only the bed 102 moves. In this state, the measurement control unit 420 instructs the X-ray irradiation timing of the X-ray tube 311 to the irradiation controller 341 and instructs the imaging timing of the X-ray detector 321 to the detection controller 344. Thereby, the measurement control unit 420 performs X-ray irradiation and X-ray photon detection by the detector, and starts measurement. Then, when the moving distance of the bed 102 reaches the distance designated by the user or the moving limit distance of the bed 102, the imaging is finished.

  By repeating this operation while changing the irradiation-measurement direction of the X-ray tube 311 and the X-ray detector 321 by changing the angle of the rotating plate 332 between imaging, transmission data in a plurality of directions is acquired. be able to. Transmission data relating to X-ray photons detected by the detector 321 in this way is collected by the data collection unit 404.

  For example, in the case of a scanogram, data collection by the data collection unit 404 is performed as follows. As described above, the rotating plate 332 moves the bed 102 while irradiating X-rays in a stationary state. In the meantime, the count number of each energy width detected by the X-ray detector 321 every 0.1 ms, for example, and discriminated into a plurality of predetermined energy ranges is obtained.

  Here, when the X-ray irradiation range is wide, multiple data collection is performed while passing through the X-ray irradiation range. In that case, the accuracy of the data is increased by using the average value for the data at the same position. Data collected by such a method is saved, or saved after the calculation by the correction unit 440 and the image generation unit 450.

  In addition, you may limit the irradiation range of the z direction of an X-ray as needed. As a method for limiting the irradiation range in the z direction, for example, a method for turning on / off the voltage application to the X-ray tube, or a shutter for attenuating and absorbing X-rays in front of the X-ray tube and turning the shutter on / off The line may be controlled. Further, data obtained by irradiating X-rays from another angle by rotating the rotating plate 332 by, for example, 90 degrees may be created.

As described above, the data collection unit 404 counts the number of X-ray photons for each energy range. The result obtained by counting indicates the distribution of energy values (unit: keV) of X-ray photons. Accordingly, the data collection unit 404 thereby obtains the energy distribution (spectrum) of the X-rays detected by the X-ray detector 321. The data collection unit 404 outputs the obtained result to the CPU 401 as count information.
Details of the energy range of the data collection unit 404 will be described later.

The data adding unit 430 adds the weighted data for each energy range collected by the data collecting unit 404. Details of the data adder 430 will be described later.
The correction unit 440 performs correction processing on the data collected by the data collection unit 404 according to predetermined correction data. The correction processing performed here is, for example, linearity correction of the reference correction circuit, logarithmic conversion processing, offset processing, sensitivity correction, beam hardening correction, water phantom calibration, CT value correction, and the like.

  The correction data is, for example, an X-ray attenuation coefficient in each energy range of the reference material that is required when discriminating the reference material from the data in each energy range. In addition, correction data in each energy range, CT value correction data, and the like may be required for each set energy range when performing phantom calibration that performs correction according to the size of the phantom. The correction data is created by executing a correction data acquisition command from the UI unit 200, for example, and using the correction unit 440 to calculate data obtained by imaging the correction phantom such as an acrylic container containing water. Ask for it.

  By setting the energy range of these correction data in advance before imaging from a scanogram or the like, calculation before imaging can be performed, and the image creation time can be shortened. These correction data can be generated for each imaging, and correction data can also be collected in advance when the data does not change even when the subject 101 changes.

  The image generation unit 450 reconstructs an X-ray CT image from the data added by the data addition unit 430. For example, the image is reconstructed by performing Log transformation on the number of X-ray photons. Various known methods such as the FeldKamp method and the successive approximation method can be used for the reconstruction.

  The data collection unit 404 counts photons derived from the X-rays detected by the X-ray detector 321 (X-ray photons) for each energy range for a plurality of predetermined energy ranges, and provides count information for each energy range. obtain.

  The data collection unit 404 acquires the energy value of each X-ray photon detected by the X-ray detector 321 and adds it to the counting result of the energy bin (Bin) provided for each energy range according to the energy value. To do. The energy bin is a storage area set for each energy range. The energy range is obtained by dividing an energy range from a certain minimum energy to a maximum energy of the X-ray tube 311 by a predetermined energy width ΔB.

  Here, the setting of the energy width is important for obtaining the CT value with high accuracy. The greater the energy range, the greater the possibility of errors. When the CT value error in water contained in a 165 cm 5 mm thick polyethylene container was obtained by ray tracing simulation, when the energy range was set to one, a maximum CT value shift of about 8 occurred. When the energy range was set to 5, the maximum was 1 or less, and when 4 was set, it was 2 or less.

  Since it is considered that there is no clinical problem if the CT value does not deviate by 2 or more, it is considered that there is no problem if the energy range is set to 4 or more. On the other hand, even if the energy range is set finely, if it is set higher than the energy resolution of the detector, the correctness of energy discrimination becomes a problem, so it is not necessary to make it too fine.

  When the CdTe semiconductor detector is used at room temperature, the energy resolution is several keV. Although the energy resolution is not unique because it depends on temperature, when the tube voltage is 120 kV with an energy resolution of 2 keV, the energy range is 60 from 120/2. However, even if an energy range of 60 or more is used, an improvement in image quality cannot be expected, and a data transfer amount increases, resulting in a complicated system. Therefore, in this embodiment, it is considered that the energy range is preferably 4 or more and 60 or less. Moreover, in each energy range, it is desirable that the difference between the upper limit value and the lower limit value is larger than the half-value width of the energy range detectable by the detector.

  Therefore, in this embodiment, the data collection unit 404 sets the energy width ΔB to 20 keV, for example, sets the minimum energy to 40 keV, and sets the maximum energy to 140 keV. That is, the energy range of 40 keV to 140 keV is divided into five energy ranges of B1 (40 to 60 keV), B2 (60 to 80 keV), B3 (80 to 100 keV), B4 (100 to 120 keV), and B5 (120 to 140 keV). To do.

  The energy width ΔB may not be the same value for B1 to B5, or may be set so that a part of the energy data is overlapped, or not set to detect specific energy data. Is possible. The DAS adds to the count result of the energy bin provided in association with the corresponding energy range according to the detected energy value of the X-ray photon.

(Details of data adding unit 430)
As described above, when obtaining a CT value, the PCCT apparatus can only obtain a count number of a specific energy width and cannot obtain detailed energy information, so it is difficult to obtain a highly accurate CT value. It is. On the other hand, in an X-ray CT apparatus to which a detector that is not a photon counting method is applied, a CT value proportional to the magnitude of energy can be obtained. This can be said to be a value obtained by applying a weight substantially corresponding to the energy of the X-ray and adding it. Therefore, also in the PCCT apparatus, the data adding unit 430 multiplies the counting result of each energy range obtained from the data collecting unit 404 with an appropriate weight in consideration of the X-ray attenuation rate and energy distribution. A CT value corresponding to the magnitude of energy can be obtained.

  The attenuation rate of X-rays (the ratio of X-rays scattered and absorbed while traveling a unit distance) increases as the energy decreases, except for special situations such as K-edge. In addition, the effect of the K edge is limited, and as a whole, the smaller the energy, the higher the effect. For this reason, when X-rays that are uniformly distributed in the energy direction are irradiated, the X-rays on the high energy side increase after passing through the subject. This means that when a certain energy range is cut, the average energy is higher after the subject is transmitted than when the subject is transmitted. In view of this effect, a more accurate CT value can be obtained by applying a weight.

  Furthermore, the X-rays emitted from the X-ray tube are not uniformly distributed in the energy direction as described above. For example, in the case of a tube voltage of 120 kV, the X-ray energy has a distribution having a peak around 60 keV. Due to this distribution, the high energy side shifts to a lower energy side than the median value of the energy range before the subject is transmitted, and the low energy side shifts to a higher energy side than the median value. It becomes more accurate by applying weights including these effects.

  In a photon counting type detector, if all signals are attenuated in one detector, the total X-ray energy is converted into an electrical signal of the corresponding detector, but some signals are attenuated by another detector. When the energy of a certain X-ray signal is obtained, X-rays may enter and be measured as a signal in another energy range. If the weight is calculated including these effects, a more accurate CT value than the above can be obtained.

  The data adding unit 430 performs addition processing multiplied by the weight determined in consideration of the above-described characteristics, and the image generation unit 450 performs processing such as reconstruction, so that the conventional X-ray CT apparatus Almost equivalent projection data can be obtained. Then, by using the projection data obtained here and performing data processing equivalent to the conventional one, it becomes possible to obtain a CT value equivalent to that of the conventional X-ray CT apparatus.

  Therefore, the data adding unit 430 first calculates the weight. For example, as the weight, the median value in each energy range is used in this embodiment. Specifically, in the case of B1 (40 to 60 keV), it is 50 keV, and in the case of B2 (60 to 80 keV), it is 70 keV. In addition to the weight set in advance as described above, the weight may be calculated each time according to the subject.

Next, imaging processing in the X-ray CT apparatus configured as described above will be described with reference to the flowchart of FIG.
First, the imaging condition setting unit 410 receives imaging conditions from the user via the UI unit 200 (step S101). The imaging conditions for accepting input include tube voltage, tube current, the thickness and shape of the X-ray filter 312, and the shape of the bow tie filter 313.
Next, a scanogram is imaged (step S102). Imaging is performed without rotating the rotating plate as shown in the description of the data collecting unit 404.

Next, the data collection unit 404 sets or changes the energy range (step S103). In the present embodiment, the ranges of B1 (40 to 60 keV), B2 (60 to 80 keV), B3 (80 to 100 keV), B4 (100 to 120 keV), and B5 (120 to 140 keV) are set.
Next, according to this setting, correction data used in the correction unit 440 and the image generation unit 450 is created (step S104). In this embodiment, since five energy ranges are set, correction data necessary for these five settings is created.

  Next, the measurement control unit 420 performs imaging in accordance with the imaging conditions set in step S102 (step S105), and the data collection unit 404 collects data according to the CT imaging method in which the rotating plate is rotated.

  Next, the collected data is added (step S106). Multiply and add the weights set in each energy range. In the present embodiment, for example, an intermediate value of each energy range is used as a weight. Only the data after the addition may be stored, or the data before the addition may be left as necessary.

  Thereafter, the correction unit 440 corrects the data collected by the data collection unit 404 (step S107). Here, it is possible to shorten the image generation time by performing correction using the correction data created in advance in step S104. In the correction unit 440, the processing may be performed using the collected count information as it is, or the processing may be performed after the information is once converted into the base material distance information.

  The image generation unit 450 generates an image using the data corrected in step S107 (step S108). Again, image generation is performed using the correction data generated in step S104 as necessary. The image generated in step S108 is stored in the HDD device 403, and the imaging sequence is completed by displaying the image on the monitor of the input device 210, for example (step S109). Thereafter, the user makes a diagnosis using the image, and performs image analysis as necessary.

  As described above, according to the present embodiment, the data adding unit 430 multiplies the weight in the addition process for the count value of each energy range, and then performs the process such as the reconstruction, so that the photon counting method detector is provided. It is possible to obtain projection data equivalent to projection data obtained in a conventional X-ray CT apparatus that is not used. Therefore, a desired CT value with high accuracy can be obtained by performing predetermined processing according to this projection data.

  In the above-described example, the data adding unit 430 uses the median value of the energy range as a weight, but another value may be used as long as it is between the energy ranges. For example, the maximum value of the energy range can be used, or the minimum value can be used. In addition, the weight condition used for each energy range may be changed.

  For example, in the energy range of B4 (100-120 keV), B5 (120-140 keV), the minimum values of 100 keV and 120 keV are used, respectively, and in B3 (80-100 keV), the median value of 90 keV is used, and B1 (40-60 keV) ) And B2 (60 to 80 keV) may use the maximum values of 60 keV and 80 keV, respectively.

(Second Embodiment)
In the X-ray CT apparatus according to the first embodiment described above, the median value is applied as a weight to each energy range. For example, the weight is set according to the transmission distance of the subject for each energy range. It may be optimized.

  The X-ray CT apparatus according to the present embodiment has the same configuration as the X-ray CT apparatus according to the first embodiment described above, and only the processes performed by each unit included in the calculation unit 400 are different. Therefore, each part constituting the X-ray CT apparatus according to the present embodiment is denoted by the same reference numerals as those of the X-ray CT apparatus according to the first embodiment, and detailed description thereof is omitted.

  In the present embodiment described below, the weight set for each energy range is changed according to the transmission distance of the subject. That is, in the present embodiment, the data adding unit 430 corrects the weight by adding a value that takes into account a change due to the transmission distance of the subject to a predetermined weight (for example, the median value in each energy range). The addition process is performed by applying the corrected weight.

  For this reason, the X-ray CT apparatus according to the present embodiment holds a correction coefficient for correcting the weight in the HDD 403 in advance, and performs an addition process using the weight corrected based on the correction coefficient. When the predetermined weight is the median value of each energy range, for example, a deviation from the median value of the energy range is stored in the HDD 403 as a correction coefficient for correcting the weight.

In the HDD 403, for example, “deviation” with respect to the median value of the energy range calculated by the following procedure is stored as a correction coefficient.
Prior to imaging the subject 101, several water phantoms are prepared and imaged. A water phantom is, for example, a phantom filled with water in an acrylic container, which simulates a human body composed mostly of water or a substance with a density close to that of water. For example, four water phantoms having a diameter of 100 mm, 200 mm, 300 mm, and 400 mm are prepared.

  When these phantoms are imaged, the energy ranges are set with finer energy ranges. For example, in B1 (40 to 60 keV), 40-42 keV, 42-44 keV,... Are set in increments of 2 keV. In this embodiment, this is imaged in increments of 2 keV in four phantoms and five energy ranges.

  Next, based on the imaging data obtained in this way, a deviation from the median of each energy range in each diameter water phantom is calculated. More specifically, the deviation of the energy range with respect to the median value is calculated based on the respective count values of the energy range set in increments of 2 keV.

For example, the deviation with respect to the energy after phantom transmission in the energy range of 40-60 keV is calculated as follows. In this embodiment, 40-42 keV, 42-44 keV,..., And 2 keV increments, respectively, ((40-42 keV count number) × 41 keV + (42-44 keV count number) × 43 keV + ... + ( The average energy is calculated as (number of counts of 58-60 keV) × 59 keV) / (count number of 40-60 keV), and the difference between the value and the median is calculated.
When passing through the phantom, X-rays on the higher energy side are transmitted, so the energy distribution is usually shifted to the higher energy side.

  Here, the attenuation count of water is shown in FIG. When 100 mm of water is transmitted, the amount detected when the amount of input X-rays in all energy ranges is 1, as shown in FIG. In this way, when the input energy spectrum is constant regardless of the energy, it shifts to the high energy side. The average energy is calculated for each phantom.

  For example, when the count shown in FIG. 5B is obtained, the average energy is 50.8 keV, and 0.8 keV is the shift amount. In the present embodiment, the method for obtaining the correction coefficient by actual measurement is shown. However, for example, the calculation can be performed by using a Monte Carlo simulation or using a theoretical attenuation amount.

  Hereinafter, imaging processing in the X-ray CT apparatus according to the present embodiment will be described with reference to the flowchart of FIG.

  First, in step S201, the imaging condition setting unit 410 receives imaging conditions from the user via the UI unit 200, and captures a scanogram in step S202. In step S203, the energy range is set or changed, and correction data is created in step S204 according to the set energy range.

  In the next step S205, the data adder 430 corrects the weight used when performing the addition process. That is, the data addition unit 430 calculates the size of the subject from the scanogram captured in step S202, reads the correction coefficient for the imaging data close to the size of the subject from the HDD 430, and corrects the weight using this. Here, the data adding unit 430 may linearly interpolate the data values of the two phantoms without using the correction coefficient for the imaging data close to the size of the subject, or function-fit several correction coefficients. The weight may be corrected accordingly.

Here, when the energy range applied when calculating the correction coefficient stored in the HDD 403 when the data adding unit 430 corrects the weight is different from the energy range obtained by imaging the scanogram, the following is performed. Determine the correction factor.
For example, when the range of 40-50 keV is set, the energy after phantom transmission in the corresponding range is once again calculated from the values of the five imaging data of 40-42, 42-44,. The deviation of the range is calculated and used as a correction coefficient.

  Furthermore, in order to obtain the subject size more accurately, the main image may be taken prior to the weight correction without using the scanogram data, and the size of the subject 101 may be obtained from the data obtained by the main imaging. .

  In the description of the present embodiment, since the predetermined weight is the median value of the energy range, the deviation with respect to the median value is used as the correction coefficient. However, the correction value is not limited to the deviation amount, and the calculated average energy correction coefficient may be used. .

  In step S206, according to the imaging condition set in step S202, imaging is executed by the measurement control unit 420, and data is collected by the data collection unit 404. The collected data is output to the data adding unit 430. In step S207, the weights set and corrected in the respective energy ranges are multiplied and added in step S205.

  Thereafter, the correction unit 440 corrects the data collected by the data collection unit 404 according to the correction data (step S208). Note that the correction unit 440 may perform processing using the information of the count collected by the data collection unit 404 as it is, or may perform processing after converting the information into distance information of the base material.

  The image generation unit 450 generates an image using the data corrected in step S208 (step S209), stores the generated image in the HDD device 403, and displays the image on the monitor of the input device 210, for example. The imaging sequence is completed (step S210).

  As described above, according to the present embodiment, the data adding unit 430 multiplies the weight in the addition process for the count value of each energy range, and then performs the process such as the reconstruction, so that the photon counting method detector is provided. It is possible to obtain projection data equivalent to projection data obtained in a conventional X-ray CT apparatus that is not used. Therefore, a desired CT value with high accuracy can be obtained by performing predetermined processing according to this projection data. A CT value with higher accuracy can be calculated by correcting the weight used in the addition processing in the data adding unit 430 according to the shift of the energy range according to the size of the subject.

(Third embodiment)
In the X-ray CT apparatus according to the first embodiment described above, the median value of each energy range is applied as a weight. However, for example, the weight can be optimized according to the energy distribution of incident X-rays. .

  The X-ray CT apparatus according to the present embodiment has the same configuration as the X-ray CT apparatus according to the first embodiment and the second embodiment described above, and only the processes performed by each unit included in the calculation unit 400 are different. It is. Therefore, each part constituting the X-ray CT apparatus according to the present embodiment is denoted by the same reference numerals as those of the X-ray CT apparatus according to the first embodiment, and detailed description thereof is omitted.

  In the present embodiment described below, the weight set for each energy range is changed in accordance with the energy distribution of incident X-rays. That is, in this embodiment, the data adding unit 430 adds a weight to a predetermined weight (for example, a median value in each energy range) by adding a value that takes into account a change due to the energy distribution of incident X-rays. Correction is performed and the corrected weight is applied to perform addition processing.

  For this reason, the X-ray CT apparatus according to the present embodiment holds a correction coefficient for correcting the weight in the HDD 403 in advance, and performs an addition process using the weight corrected based on the correction coefficient. When the predetermined weight is the median value of each energy range, for example, a deviation from the median value of the energy range is stored in the HDD 403 as a correction coefficient for correcting the weight.

In the HDD 403, for example, “deviation” with respect to the median value of the energy range calculated by the following procedure is stored as a correction coefficient.
Prior to imaging the subject 101, imaging is performed without a phantom. At the time of imaging, the energy range is set by finely setting each energy range. For example, in B1 (40 to 60 keV), 40-42 keV, 42-44 keV,... Are set in increments of 2 keV, and imaging is performed. And the shift | offset | difference of the median value of each energy range is calculated for every detector using the acquired imaging data.

  There are mainly two reasons for calculating for each detector. The first is that there is a change in the energy spectrum in the body axis direction of the subject due to the heel effect well known as a physical phenomenon of X-rays. Second, in order to reduce exposure to the subject 101 with respect to the X-ray tube rotation direction, the imaging center is thin, and a metal (for example, aluminum) that is thicker toward the outside of the imaging field of view is used. This is because the X-ray distribution changes in both the body axis direction and the X-ray tube rotation direction because a filter that suppresses X-ray irradiation is used for the portion that does not pass through.

  Since imaging is performed without a phantom in order to calculate the correction coefficient, it is possible to obtain a deviation from the median value in each energy range of X-rays before entering the subject 101 based on data obtained by imaging. it can.

In general, the energy distribution of the X-ray tube has a peak that depends on the set tube voltage and a plurality of peaks (may be present depending on the material) due to the target material of the X-ray tube (see FIG. 7). The range higher than the peak depending on the set tube voltage is shifted to the low energy side when viewed from the center, and conversely, the portion lower than the peak is shifted to the high energy side when viewed from the center.
In addition, the energy distribution varies depending on the filter that suppresses X-ray irradiation. For example, as in the above-described imaging with a water phantom, the high energy side is transmitted and the low energy side is more scattered and absorbed, so that the energy distribution changes.

  Therefore, in addition to the above-mentioned complex factors such as the deviation depending on the set tube voltage and the deviation due to the filter, etc., correction according to the change in the X-ray energy spectrum accompanying the structure of other X-ray tubes and X-ray irradiated portions A coefficient can be obtained. In the present embodiment, a method for obtaining a correction coefficient by actual measurement by performing imaging in the absence of a phantom, that is, a method for obtaining a correction coefficient by actual measurement, for example, using a Monte Carlo simulation or using a theoretical attenuation amount, A correction coefficient can also be calculated by calculation.

  When there is no need to recalculate the weight for each subject, the weight corrected using the correction coefficient obtained in consideration of the energy distribution is stored in advance, and this is stored in the data adder 430 in each data adder 430. It can also be applied during the addition process for the energy range count value.

  As described above, according to the present embodiment, the data adding unit 430 multiplies the weight in the addition process for the count value of each energy range, and then performs the process such as the reconstruction, so that the photon counting method detector is provided. It is possible to obtain projection data equivalent to projection data obtained in a conventional X-ray CT apparatus that is not used. Therefore, a desired CT value with high accuracy can be obtained by performing predetermined processing according to this projection data. Then, by correcting the weight used in the addition processing in the data adding unit 430 according to the deviation caused by the energy distribution, a CT value with higher accuracy can be calculated.

  Note that, in addition to the correction count in consideration of the deviation caused by the energy distribution, the HDD 403 can store a correction coefficient corresponding to the transmission distance of the subject in the second embodiment. In this case, the weight can be corrected in consideration of the energy distribution and in accordance with the transmission distance of the subject, so that the weight can be further optimized and a highly accurate CT value can be calculated.

(Fourth embodiment)
As described above, the weight applied during the addition process can be optimized by taking into account the deviation due to the energy distribution, taking into account the transmission distance of the subject, and taking into account fluctuations in detection at the detector. Can be optimized.

  Examples of fluctuations in detection by a detector include pileup in a photon counting detector. That is, in the photon counting type detector, if the amount of incident X-rays increases, countdown occurs due to pileup, and it is preferable to correct this.

For this reason, a correction coefficient that takes into account pileup is stored in the HDD 403 in advance. The correction coefficient is calculated as follows.
Prior to imaging of the subject, imaging is performed without a phantom. As in the second embodiment described above, the energy range at the time of imaging is set by finely setting each energy range. For example, in B1 (40 to 60 keV), 40-42 keV, 42-44 keV,... Are set in increments of 2 keV, and imaging is performed by changing the tube current.

  Next, the deviation of the median value of each energy range is calculated for each detector using the imaging data obtained by this imaging, and the calculated deviation is stored in the HDD 403 as a correction coefficient. The reason for calculating the deviation for each detector is that there is a concern that the detector characteristics may vary. As the tube current increases and the amount of incident X-rays increases, the count number increases in direct proportion. However, when a pile-up occurs, it deviates from a directly proportional relationship. With the deviation from this direct proportion, the amount of signal in each energy range decreases. More accurate correction can be performed by measuring the change in the distribution and correcting the decreased signal amount.

  When correcting fluctuation due to pileup, that is, counting down, it is necessary to calculate the number of counts per unit time from the imaged data and to perform correction using the result. For this reason, weight correction is performed after imaging.

DESCRIPTION OF SYMBOLS 100 ... PCCT apparatus, 101 ... Subject, 102 ... Bed, 200 ... UI unit, 210 ... Input device, 220 ... Output device, 300 ... Measurement unit, 310 ... · X-ray irradiation unit, 311 ··· X-ray tube, 312 ··· X-ray filter, 313 ··· Bowtie filter, 320 ··· X-ray detector, 321 ··· X-ray detector, 322 ··· Detecting element, 323 ... collimator, 324 ... counting circuit, 325 ... integrating circuit, 330 ... gantry, 331 ... opening, 332 ... rotating plate, 340 ... controller 341 ... irradiation controller 342 ... gantry controller 343 ... table controller 344 ... detection controller 400 ... calculation unit 401 ... central processing unit 402 ..Memory, 403 HDD Location, 404 ... data acquisition unit, 410 ... imaging condition setting unit, 420 ... measurement control unit, 430 ... data adding unit, 440 ... correction unit, 450 ... image generating unit

Claims (6)

A detector for detecting X-ray photons irradiated from an X-ray source that circulates around the subject on the bed and transmitted through the subject;
A data collection unit that outputs a count value for each energy range for a plurality of predetermined energy ranges by collecting and processing X-ray photons detected by the detector;
A data addition unit that performs an addition process by multiplying the count value output by the data collection unit by a weight that is determined for each energy range and is proportional to the median value of the energy range;
An X-ray CT apparatus comprising: an image generation unit that generates an image by performing reconstruction processing using data added by the data addition unit. A storage unit storing a correction coefficient for correcting the weight according to the size of the subject;
The X-ray CT apparatus according to claim 1, wherein the data adding unit corrects the weight using the correction coefficient in the addition process. A storage unit storing a correction coefficient for correcting the weight according to the energy distribution of incident X-rays;
The X-ray CT apparatus according to claim 1, wherein the data adding unit corrects the weight using the correction coefficient in the addition process. A storage unit storing a correction coefficient for correcting the weight according to the characteristics of the detector;
The X-ray CT apparatus according to claim 1, wherein the data adding unit corrects the weight using the correction coefficient in the addition process. The data collection unit sets the energy range to 4 or more,
The X-ray CT apparatus according to any one of claims 1 to 4, wherein a difference between an upper limit value and a lower limit value is larger than a half-value width of the detector in each energy range. Detecting X-ray photons irradiated from an X-ray source that circulates around the subject on the bed and transmitted through the subject;
By collecting and processing detected X-ray photons, a count value is output for each energy range for a plurality of predetermined energy ranges,
The count value is determined for each energy range, and an addition process is performed by multiplying a weight proportional to the median value of the energy range,
An image generation method for generating an image by performing reconstruction processing using data obtained by the addition processing.

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