JP6665158B2 - X-ray CT system - Google Patents

Description

  The present invention relates to an X-ray CT apparatus equipped with an energy-separating X-ray detector that separates and counts X-ray photons incident on a plurality of energy ranges, and a multi-energy image creating method.

  The X-ray CT apparatus calculates an X-ray absorption coefficient (ray attenuation count) from an X-ray transmission image (hereinafter, referred to as projection data) of a subject taken from a plurality of directions, and obtains a tomographic image of the subject (hereinafter, referred to as a reconstructed image). This is a device for obtaining a structural image, and is widely used in the fields of medical treatment and nondestructive inspection.

  Many current medical X-ray CT systems are equipped with an X-ray detector called an integral type. The integrating X-ray detector converts X-rays into light using a scintillator, converts the light into charges using a photodiode, converts the charges into digital signals using a readout circuit, and outputs the digital signals. The readout circuit integrates the electric charge for each view to obtain a digital signal. Here, a large number of X-ray photons are incident during one view, but the obtained signal amount is a signal amount corresponding to the total energy of each X-ray photon. Therefore, it is not possible to know the energy information of each incident X-ray photon.

  As a conventional technique for overcoming this, there is a dual energy imaging method. In this imaging method, for example, as described in Patent Document 1, an image is acquired with a plurality of X-rays having different energy spectra, and the projection data and the reconstructed image are used to obtain a reference material assumed by the apparatus. And the tissue can be separated to obtain respective identification images. In the dual energy imaging method, images such as a monochromatic X-ray equivalent image, a reference material density image, an effective atomic number image, an electron density image, a photoelectric effect image, a Compton scattering image, and an absorption coefficient image in a spectrum other than the spectrum used in imaging are used. Is obtained. Here, an image created by such a dual energy imaging method is hereinafter referred to as a dual energy image.

  On the other hand, as in Non-Patent Document 1, in recent years, the development of an X-ray CT apparatus equipped with a photon counting type X-ray detector (hereinafter referred to as a photon counting CT apparatus) has been progressing. This X-ray detector includes an X-ray detection element having a semiconductor detection layer such as CdTe (cadmium telluride) and a readout circuit that obtains a digital signal by separating each energy range according to the energy of incident X-ray photons. Have.

  In the X-ray detector, when an X-ray photon is incident on the X-ray detection element, first, charges corresponding to the energy of the X-ray photon are generated in the detection layer. Next, a readout circuit reads out this electric charge at such a high speed that it can be read out for each X-ray photon, separates the electric charge into several energy ranges according to the energy of the incident X-ray photons, Count the number of photons. At this time, the energy of the incident X-ray photon is determined using the generated charge amount.

  Further, this detection is similarly performed for each of a plurality of X-ray photons, the number of X-ray photons is counted in each energy range, and each count is converted into a digital signal. By such measurement, projection data is obtained for each energy range, and by using these, a reconstructed image can be obtained for each energy range. By using this, an image in which a specific substance is extracted can be obtained (hereinafter, referred to as a photon counting image). For example, in Non-Patent Document 1, a K-edge image of gold is acquired using a K-edge of gold in a contrast agent.

JP 2009-17984 A

David P. Cormode, Ewald Roessl, Axel Thran, et al. Analysis with Multicolor CT and Target Gold Nanoparticles. Radiology 2010; 256 (3): 774-782

  An X-ray CT system using dual energy imaging (hereinafter referred to as a dual energy CT system) requires a large effective energy to acquire accurate dual energy images, separate reference materials, and identify materials. Ideally, images acquired using different energy spectra should be used.However, due to the limitations of the X-ray tube and the problem of increasing the exposure dose in order to obtain sufficient transmitted X-rays at low energies, the There is a limit to the energy spectrum that can be obtained, and there is a limit to the creation of a dual energy image and the improvement of the classification accuracy and the identification accuracy. There was also a limit on the number of reference substances to be separated.

  On the other hand, in the photon counting CT apparatus, in order to realize a large number of energy ranges, a large number of comparators are required to compare signal pulses generated by incident X-rays within the energy range. There is a problem that the power consumption of the circuit increases. Alternatively, since it is necessary to compare the energy range of the signal pulse generated by the incident X-ray many times, there is a problem that it takes time to perform classification and identification. Further, an increase in the circuit size leads to an increase in the price of the device, and an increase in the power consumption of the circuit may shorten the life of the device.

  Also, with photon counting CT equipment, the energy range cannot be increased countlessly due to the circuit scale, imaging time, realization of the required SNR of projection data of one energy range, etc., as in the case of using dual energy imaging method In addition, there is a limit in improving the accuracy of creating a photon counting image.

  An object of the present invention is to provide an X-ray CT apparatus capable of acquiring a multi-energy image with high accuracy of classification.

  In order to solve the above problems, the X-ray CT apparatus of the present invention has an X-ray detector that separates and detects energy in a plurality of energy ranges, and changes the energy distribution of X-rays detected by the X-ray detector for each energy range. (A detected energy distribution changing unit), and a multi-energy calculating unit that performs arithmetic processing on data in a plurality of energy ranges detected by the X-ray detector under different energy distribution conditions.

  In one embodiment of the present invention, the detected energy distribution changing unit includes an irradiated X-ray spectrum changing unit, and the irradiated X-ray spectrum changing unit changes the irradiated X-ray spectrum to change the detected X-ray energy distribution for each energy range. I do. In another aspect, the detected energy distribution changing unit includes an energy threshold changing unit that changes an energy threshold that determines an energy range of the X-ray detector, and the energy threshold changing unit changes the energy threshold, Change the detected X-ray energy distribution.

  The multi-energy calculation unit creates multi-energy projection data using projection data obtained from a plurality of detected X-ray energy distributions, and / or generates a multi-energy projection data using a reconstructed image obtained from a plurality of detected X-ray energy distributions. Create an energy image. In this specification, a multi-energy image refers to a broad concept including a dual energy image and a photon counting image.

  According to the present invention, it is possible to obtain more projection data having different energy ranges and different irradiation X-ray spectra as compared with a conventional dual energy CT apparatus, and it is possible to improve the accuracy of dual energy images and classification and identification. . In addition, a photon counting image can be created with higher accuracy than a conventional photon counting CT apparatus. Further, the number of energy ranges can be reduced, and the circuit scale, circuit power consumption, and conversion processing time can be reduced.

FIG. 1 is an overall schematic diagram showing an X-ray CT apparatus according to a first embodiment of the present invention. Explanatory drawing explaining an example of the arrangement of X-ray detection elements in an X-ray detector Explanatory drawing for explaining an example of the structure of an X-ray detector Explanatory drawing explaining the energy classification method performed by the X-ray detector of the first embodiment Functional block diagram showing a configuration example of an arithmetic unit according to the first embodiment FIG. 3 is a diagram illustrating an example of a data processing flow performed by the X-ray CT apparatus of the first embodiment. The figure which shows an example of the data processing flow of another method performed by the X-ray CT apparatus of the first embodiment. The figure which shows an example of the data processing flow of another method performed by the X-ray CT apparatus of the first embodiment. Explanatory drawing for explaining an example of an identification map used for X-rays of the first embodiment The figure which shows an example of the data processing flow of another method performed by the X-ray CT apparatus of the first embodiment. Functional block diagram showing a configuration example of a calculation unit of the second embodiment of the X-ray CT apparatus of the present invention The figure which shows an example of the control flow performed in the X-ray CT apparatus of the second embodiment. The figure which shows another example of the control flow performed in the X-ray CT apparatus of the second embodiment. Overall schematic diagram showing an X-ray CT apparatus according to a third embodiment of the present invention Explanatory drawing explaining the function of the energy threshold changing unit in the third embodiment Functional block diagram showing a configuration example of an arithmetic unit according to the third embodiment The figure which shows an example of the data processing flow performed in the X-ray CT apparatus of 3rd Embodiment Functional block diagram showing a configuration example of an arithmetic unit according to the fourth embodiment Explanatory drawing explaining the function of the energy threshold changing unit in the fourth embodiment The figure which shows an example of the data processing flow of 4th Embodiment Illustration for explaining the principle of K-edge imaging Explanatory drawing explaining the relationship between the energy threshold value and the K edge energy in the fourth embodiment Overall schematic view showing an X-ray CT apparatus according to a fifth embodiment of the present invention Functional block diagram illustrating a configuration example of a control unit according to the fifth embodiment. The figure which shows an example of the timing control of 5th Embodiment The figure which shows an example of the data acquired in 5th Embodiment Functional block diagram illustrating a configuration example of an arithmetic unit according to the sixth embodiment. FIG. 9 is a view for explaining an example of a GUI according to the sixth embodiment. The figure which shows an example of the control flow of 6th Embodiment. Explanatory drawing explaining an example of the subject model of the sixth embodiment The figure which shows the example of a change of the control flow of FIG.

<First embodiment>
The X-ray CT apparatus of the present embodiment includes an X-ray source (100) for irradiating X-rays, an X-ray detector (104) for separating and detecting incident X-rays into a plurality of energy ranges, and a plurality of energy ranges. In each of the above, the detected energy distribution changing unit that changes the detected X-ray energy distribution detected by the X-ray detector, and processing the output signals of the X-ray detectors in multiple energy ranges with multiple detected X-ray energy distributions A signal acquisition unit (108) that creates projection data by reconstructing the projection data to create a reconstructed image by reconstructing the projection data (1053), and projecting data acquired by a plurality of the detected X-ray energy distributions. And / or a multi-energy calculation unit (1052) for generating multi-energy projection data using the plurality of detected X-ray energy distributions and / or generating a multi-energy image using the reconstructed images obtained from the plurality of detected X-ray energy distributions.

The detection energy distribution changing unit includes an irradiation X-ray spectrum changing unit (111) that changes an irradiation energy spectrum (irradiation X-ray energy spectrum) of the X-ray irradiated by the X-ray source. The spectrum is changed to change the detected X-ray energy distribution for each energy range.

  In the X-ray CT apparatus of the present embodiment, for example, an X-ray detector (104) detects an incident X-ray photon, and a photon counting type X-ray detection element that performs counting by separating the energy into two or more energy ranges. Reference numeral (400) denotes an X-ray detector in which a plurality of X-ray detectors are arranged. The signal collection unit (108) collects the count number of the X-ray detection elements and acquires projection data. The calculation unit (105) includes a correction processing unit (1051) that performs a correction process on the signal collected by the signal collection unit (108), and a multi-energy calculation that performs a multi-energy calculation to generate projection data of a multi-energy image. An operation unit (1052) and a reconfiguration unit (1053) for performing reconfiguration operation processing are provided.

  The multi-energy calculation performed by the multi-energy calculation unit (1052) uses signals obtained at a plurality of focus positions and a plurality of energy ranges to project projection data of a multi-energy image such as a dual energy image or a photon counting image (hereinafter referred to as a multi-energy image). , Multi-energy projection data).

  Hereinafter, the configuration and operation of the X-ray CT apparatus of the present embodiment will be described with reference to the drawings.

  As shown in FIG. 1, the X-ray CT apparatus of the present embodiment includes an X-ray source 100, an irradiation X-ray spectrum changing unit (hereinafter, abbreviated as a spectrum changing unit) 111, and an X-ray source 100 as an imaging system. An X-ray detector 104 arranged in an irradiation range of X-rays irradiated from the gantry rotating unit 101 which is arranged so that these X-ray source 100 and the X-ray detector 104 face each other, and rotates around a predetermined rotation axis. It has. An opening into which the subject 102 is inserted is provided at the center of the gantry rotating unit 101, and a couchtop 103 on which the subject 102 is laid is arranged in the opening. The couch top 103 and the gantry rotating unit 101 are configured to be relatively movable in a predetermined direction.

  The X-ray source 100 according to the present embodiment causes an electron beam accelerated by a tube voltage to collide with a target metal such as tungsten or molybdenum, and generates X-rays from the collision position (focal point). The spectrum changing unit 111 changes the spectrum of the X-ray emitted from the focal point, for example, by changing the tube voltage or the X-ray filter. The X-ray filter can include metals such as tungsten, molybdenum, copper, tin, and the like.

  The X-ray CT apparatus includes a control unit 107, a signal collection unit 108, an arithmetic operation unit, and a control system that controls the imaging system and a signal processing system that processes a signal acquired by the X-ray detector 104 in accordance with the operation of the imaging system. A unit 105, a display unit 106, an input unit 110, a storage unit 109, and the like are provided.

The control unit 107 includes an X-ray control unit that controls the operation of the generation drive source of the X-ray source 100, a readout control unit that controls the signal readout operation of the X-ray detector 104, the rotation of the gantry rotation unit 101, and the couchtop 103. It is composed of a photographing control unit that controls the movement of the camera, and an overall control unit that controls all of these units.
Further, a display control unit or the like for controlling display on the display unit 106 can be provided.

  The control unit 107 and the arithmetic unit 105 can be partially or entirely configured as a system including a CPU (Central Processing Unit), a memory and a storage unit 109, and the functions of each unit constituting the control unit 107 and the arithmetic unit 105 are as follows. The program can be realized by the CPU loading a program stored in the storage unit in advance into the memory and executing the program. Some of the functions can be configured by hardware such as an ASIC (Application Specific Integrated Circuit) or an FPGA (Field Programmable Gate Array).

  Unless otherwise described, components constituting an imaging system, a control system, and a signal processing system have the same configuration and have the same functions as those included in a known X-ray CT apparatus.

  The X-ray detectors 104 are arranged in a plurality of arcs around the X-ray source 100, and rotate while maintaining the positional relationship with the X-ray source 100 as the gantry rotating unit 101 rotates. Although FIG. 1 shows the case where the number of X-ray detectors 104 is eight for the sake of simplicity, the actual device has, for example, about forty. An X-ray grid (not shown) is provided on the front surface of the X-ray detector 104, and among the X-rays emitted from the X-ray source 100, X-rays scattered by the subject 102 and the like are X-rays. It prevents the light from entering the line detector 104.

The X-ray detector 104 has a structure in which a plurality of photon-counting X-ray detection elements 400 are two-dimensionally arranged in a channel direction and a slice direction, for example, as shown in FIG. Here, FIG. 2 shows a part of the X-ray detection element 400 arranged in the X-ray detector 104, which is cut out for four in the channel direction and three in the slice direction. Further, the X-ray detection element 400 is arranged such that the channel direction and the rotation direction coincide with the slice direction and the rotation axis direction.

  As shown in FIG. 3, each X-ray detection element 400 of the X-ray detector 104 has a structure in which positive and negative electrodes 402 and 403 are provided so as to sandwich the detection layer 401, and a readout circuit 405 is connected to the electrodes. Having. In the present embodiment, the negative electrode 402 has a common structure for each X-ray detection element 400. The X-rays enter the detection layer 401 from the negative electrode 402 side as indicated by an arrow 404. The detection layer 401 is made of a semiconductor material such as, for example, CdTe (cadmium telluride), CdZnTe (cadmium zinc tellurium), and Si (silicon), detects an incident X-ray photon, and generates an amount of charge corresponding to the energy. . The readout circuit 405 reads out the charges generated in the detection layer 401 at a predetermined sampling interval, and separates the energy of the incident X-ray photons into a plurality of energy ranges based on a predetermined threshold value based on an electric signal generated by the charges.

  For example, the two energy ranges are determined based on whether the energy range is less than a predetermined threshold (hereinafter referred to as a low energy range) or an energy range equal to or more than a predetermined threshold (hereinafter referred to as a high energy range). Such discrimination is performed for each sampling, and when an X-ray photon is incident, it is classified into a high energy range and a low energy range, and the number of each X-ray photon is counted for each view.

  The sorting method will be described with reference to FIG. FIG. 4 is a graph showing an output voltage 120 generated by generated charges, wherein the horizontal axis represents time and the vertical axis represents voltage. In the illustrated example, X-rays enter during a sampling time 123 to generate a pulse output 121, and X-rays enter during a sampling time 125 to generate a pulse output 122. Note that FIG. 4 shows a case where sampling is performed not only at the timing of X-ray incidence but also periodically when no X-rays are incident (sampling time 124), but sampling is performed at the timing of X-ray photon incidence. It can be done.

  The readout circuit (405 in FIG. 3) compares the maximum value of the output voltage in that section with the energy threshold value 126 and the energy threshold value 127 by a comparator for each sampling to discriminate. The energy threshold 126 is a range in which the incident X-ray photon has a high energy range or a low energy range. The energy threshold 127 determines whether or not there is an input of an X-ray photon. Here, since the output voltage 120 fluctuates due to the circuit noise of the X-ray detector 104 even when X-rays are not input, the energy threshold 127 is larger than zero so as not to erroneously detect this as a signal due to X-rays. Expected a value.

  By using these energy thresholds, for example, at the sampling time 124 in FIG. 4, the output voltage 120 is equal to or less than the energy threshold 127, so that it is determined that no X-ray photon is input. At the sampling time 125, since the output voltage 120 is larger than the energy threshold 126, it is determined that X-rays in a high energy range have entered. At the sampling time 123, the output voltage 120 is higher than the energy threshold 127 but lower than the energy threshold 126, so that it is determined that X-rays in the low energy range have entered. As described above, the presence or absence of incidence and the energy range are classified.

  Instead of performing the classification using the maximum value in the sampling, for example, an integrated value of the output voltage during the sampling may be used, and the classification method is not limited to the above method.

  Based on the above configuration, a general imaging operation of the X-ray CT apparatus will be described as an example of a case where the energy range is two and the spectrum of the irradiated X-ray is two. However, this is for the sake of simplicity and does not limit the present invention. Three or more energy ranges may be provided, and the irradiation X-ray spectrum may be changed to three or more types.

  First, when the start of actual imaging is input from the input unit 110, the control unit 107 controls irradiation of the X-rays from the X-ray source 100 and the gantry rotating unit 101 to start imaging. At this time, for example, it is assumed that the electron beam is accelerated by a tube voltage of 140 kV and X-rays are emitted from the X-ray source 100, and the spectrum of the irradiated X-rays at this time is hereinafter referred to as a first spectrum.

  The X-rays radiated from the focal point of the X-ray source 100 are radiated toward the subject 102 placed on the couch top 103, and the X-rays transmitted through the subject 102 are detected by the X-ray detector 104. The X-ray detector 104 separates into a high energy range and a low energy range as described above, according to the energy of the incident X-ray. Further, this classification is performed a predetermined number of times during one view, and the number of X-ray photons incident on the high energy range and the low energy range is counted. The signal collecting unit 108 converts a signal corresponding to each X-ray photon number into a digital signal and outputs the digital signal as a count number in each energy range.

  Next, the control unit 107 changes the X-ray irradiation angle on the subject 102 by rotating the gantry rotation unit 101 in the rotation direction in such imaging. In this view, measurement is performed in the same way as in the previous view, and the count is output in each energy range. Here, the X-ray generated from the X-ray source 100 may be a pulse X-ray synchronized with a view or a continuous X-ray. Further, while rotating in this way, the focus position is changed for each view, and the photographing is repeated, thereby obtaining a 360-degree digital signal. The photographing is performed over a plurality of views, for example, with one view being 0.4 degrees.

  Next, the spectrum changing unit 111 changes, for example, the tube voltage for accelerating the electron beam to 80 kV, and changes the spectrum of the X-ray source 100. This spectrum is hereinafter referred to as a second spectrum. In the imaging of the second lap, the number of X-ray photons incident on the X-ray detector 104 is counted in each view as in the first lap while being classified into a high energy range and a low energy range. Counting is performed while changing the irradiation angle of the line to obtain digital data for 360 degrees. The digital data obtained in this manner is hereinafter referred to as projection data. Here, projection data is obtained for each of two rounds of shooting.

  Next, the calculation unit 105 performs a predetermined correction process and a multi-energy calculation process on the projection data collected by the signal collection unit 108 to create multi-energy projection data. Next, the arithmetic unit 105 performs a reconstruction process on the multi-energy projection data to create a multi-energy image of the subject 102. The result is displayed on the display unit 106.

  The X-ray CT apparatus according to the present embodiment is characterized by multi-energy operation processing performed by the operation unit 105. Hereinafter, details of the arithmetic unit 105 will be described. First, a configuration example of the arithmetic unit 105 is shown in FIG. As shown in FIG. 5, the calculation unit 105 is roughly divided into a main control unit 1050, a correction processing unit 1051, a multi-energy calculation unit 1052, and an image reconstruction unit 1053, and the correction processing unit 1051, the multi-energy calculation unit 1052, The image reconstruction unit 1053 operates under the control of the main control unit 1050.

  The correction processing unit 1051 performs a process necessary for a later calculation on the projection data (raw data) acquired by the signal collection unit 108, and includes a defect element correction unit 1054, an air correction unit 1055, and the like. The multi-energy operation unit 1052 mainly has a function of creating multi-energy projection data by multi-energy operation, and FIG. 5 illustrates a case where the multi-energy operation unit includes a density image calculation unit 1056 and a multi-energy projection data calculation unit 1057. The configuration of the multi-energy operation unit 1052 is appropriately changed according to the energy operation method and the additional function.

  The parameters and data used for the calculation of the calculation unit 105 are stored in the storage unit 109, and the calculation unit 105 reads the parameters and the like from the storage unit 109 as necessary, and performs correction processing, multi-energy calculation processing, image reconstruction. Perform calculations such as The parameters and data include, for example, a defective element position map used by the defective element correction unit 1054, an X-ray sensitivity distribution and an X-ray distribution used by the air correction unit 1055, and data used by the multi-energy calculation unit 1052 when performing multi-energy calculation. (X-ray spectrum distribution and mass absorption coefficient data).

  Next, data processing performed by the arithmetic unit 105 will be described with reference to the flowchart of FIG.

  As shown in FIG. 6, the calculation unit 105 first performs defect element correction S701 on the projection data 143 received from the signal collection unit. This correction is performed, for example, by identifying a defective X-ray detection element (defective element) based on the defective element position map 141 measured and created in advance of the main imaging and stored in the storage unit 109, and outputting the X-ray detection element. This is a process of estimating a value. In the method of estimating the output value, for example, an average value is calculated using output values of normal X-ray detection elements 400 around the defective element, and the average value is used as the output value of the defective element.

  Next, air correction S702 is performed. This correction is realized, for example, by dividing the projection data by the sensitivity / X-ray distribution data 142 measured and created in advance of the main imaging and stored in the storage unit 109.

  The sensitivity / X-ray distribution data 142 is created for each energy range. The creation method is, for example, without providing the subject 102, irradiating X-rays from the X-ray tube 100 to obtain projection data 143 for each energy, performing defect element correction S701 on them, and then performing X-ray detection. Addition and averaging is performed in the view direction for each element 400, and normalized by the average value of the output from the X-ray detector 104 to create the element.

  This correction process is performed on each of the plurality of projection data 143 acquired in each energy range and irradiation X-ray spectrum.

  Next, in multi-energy operation processing S703, multi-energy projection data 144 is created. Details will be described later.

  After performing the above processing to obtain the multi-energy projection data 144, a reconstruction process S704 is performed to create a reconstructed image 145. Finally, the reconstructed image 145 is displayed on the display unit 106 (S705).

  Next, the principle and processing method in the multi-energy operation processing S703 will be described in detail.

First, the principle will be described in comparison with the conventional dual energy imaging method. As an example, a case will be described where a subject assumes two reference substances and creates a multi-energy image such as a reference substance density image or a monochromatic X-ray equivalent image. In the following description, two reference substances are referred to as reference substance 1 and reference substance 2, and their mass absorption coefficient (mass attenuation coefficient) is μm _n (ε) (n is an integer of 1 or 2; The density at the position r is _denoted by ρ _n (r).

  Here, the mass absorption coefficient of the reference substance can be stored as calculation data 140 in the storage unit 109 before imaging the subject, and is used in the calculation. Therefore, it is desirable that the reference material be selected from those that are determined in advance. For example, the input imaging method, imaging conditions, the characteristics of the subject and the imaging site, is automatically determined from the purpose of dual energy imaging or the like, or the operator can input the input unit 110 from the reference material registered in the apparatus. Alternatively, it may be configured to select and determine via the Internet.

  Here, the imaging method means simple imaging, contrast imaging, perfusion imaging, and the like, the imaging conditions include tube current, tube voltage, scan time, dose, and the like, and the purpose of dual energy imaging is bone and This means separation of contrast agents, plaque discrimination, reduction of metal artifacts, and the like.

First, the case of the conventional dual energy imaging method will be described. In this imaging, X-rays of two types of irradiation X-ray spectra are irradiated and detected without energy separation. These two types of X-ray spectra are referred to as a first spectrum and a second spectrum, the energy is ε, the number of irradiated X-ray photons at the energy ε of the first spectrum is S ₁ (ε), and the energy of the second spectrum is The number of irradiated X-ray photons at the energy ε is denoted as S ₂ (ε), respectively. At this time, the number of detected photons detected by the X-ray detection element 400 having the irradiated X-ray photons having the energy ε is calculated by using the first and second X-ray spectra, the mass absorption coefficient and the density of the reference substance, and the equation (1 ). Here, a is an integer of 1 or 2, and represents the type of X-ray spectrum.

Where [delta] _n are those obtained by multiplying the thickness of the density [rho _n (r) of the reference substance n, as the mark the path from the focal point to the X-ray detecting elements 400 of the subject by s equation (2) I can write.

Next, the energy ε is considered in each energy range and each X-ray spectrum. At this time, the output is obtained by integrating the term of the equation (1) over the entire energy range, and thus the output value P _a of the projection data in each X-ray spectrum obtained by a certain X-ray detection element 400 (a is the X-ray spectrum It can be understood from Equation (1) that Equation (1) can be written as Equations (3-1) and (3-2) (hereinafter, collectively referred to as Equation (3) as appropriate). Where [delta] _n is a parameter defined in equation (2). Also, the output value of the projection data may have a gain unique to the apparatus, but the output value here represents the number of X-ray photons without gain.

Here, since the reference substance is a known substance, generally known data can be used for the mass absorption coefficients μm ₁ (ε) and μm ₂ (ε). Further, the X-ray photon numbers S ₁ (ε) and S ₂ (ε) of the irradiation X-ray spectrum irradiated from the X-ray tube can be determined in advance by a database, a simulation, an actual measurement, or the like, if the irradiation conditions are determined. Therefore, in equation (3), Δ ₁ and Δ ₂ are variables. This variable can be determined using output values P ₁ and P ₂ determined from projection data of the subject by performing imaging.

When this δ _n (n = 1, 2) is determined, for example, projection data P (ε ₀ ) of a specific energy ε ₀ can be calculated from Expression (4), and by reconstructing this, data of the subject can be obtained. A monochromatic X-ray equivalent image can be obtained.

Here Equation (4), the irradiated X-ray spectrum as existing only energy epsilon _0, are those derived from equation (3), S (epsilon ₀₎ is the X-ray photons number emitted by the energy epsilon ₀ is there. This S (ε ₀ ) is a value determined by obtaining an X-ray spectrum under a certain imaging condition by a database, simulation, actual measurement, and the like, and using the number of X-ray photons of the first spectrum and the second spectrum. Alternatively, the number of X-ray photons in the spectrum under other conditions may be used.

On the other hand, in imaging using the apparatus of the present embodiment, X-rays of two types of irradiation X-ray spectra are irradiated, and X-ray photons incident on the X-ray detection element 400 are classified into a high energy range and a low energy range. I do. This energy range is described as b (b is H when the energy is high, and L when the energy is low), and the value of the projection data of the X-ray detection element 400 in each energy range b is P _ab for each spectrum type a. Thus, it can be understood that P _ab can be written from Expression (1) as Expressions (5-1) to (5-4) (hereinafter, collectively referred to as Expression (5) as necessary).

Here, ∫ _L means integration in the low energy range, and ∫ _H means integration in the high energy range. The [delta] _n (n = 1, 2) is a parameter defined in equation (2).

Equation (5) is, similarly to equation (3), the mass absorption coefficients μm ₁ (ε) and μm ₂ (ε) of reference substance 1 and reference substance 2, and the number of photons S _{1 in} the first and second spectra. Since (ε) and S ₂ (ε) can be determined in advance, only δ ₁ and δ ₂ are variables. Then, this variable can be determined when P ₁ and P ₂ are determined by photographing, similarly to the case where δ is obtained from Expression (3). However, in the method of the present embodiment, since there are four simultaneous equations for two variables, δ _n (n = 1, 2) is more accurately compared to the conventional dual energy imaging method. You can ask.

As an example of a method of solving the equation (5), for example, a conventional method of least squares can be used. That is, the residual Δ is defined as in equation (6), and Δ _n (n = 1, 2) is determined so as to minimize the residual. Here, w is a weight, for example, a value such as 0.5, 1, or 2.

  As a modification example of the above-described solution, for example, the residual Δ is defined by changing the weight according to the energy range and the X-ray spectrum as shown in Equation (7), and is calculated by minimizing this. May be. Here, c, d, and f are constants, and may be determined in accordance with, for example, the number of X-ray photons irradiated to each energy range in each X-ray spectrum, its noise, or SNR.

  Further, a residual of another definition may be used. Further, it may be determined using a method other than the least squares method, for example, singularity decomposition.

In general, the measurement value has noise due to fluctuations in the X-ray quantum number and noise due to equipment and extrinsic noise.Therefore, there is an error in the coefficient.However, in the method of the present embodiment, when the equation (5) is solved, In consideration of the measurement error, output value, energy, and the like of each projection data, a weight can be provided for each equation when solving the equation, so that δ _n can be obtained with high accuracy. On the other hand, in the conventional dual energy imaging method, since the output value is obtained as the sum of the low energy range and the high energy range, only two equations are used. This means that the data of each energy range has the same weight and is fixed, and there may be a case where accuracy is high in one energy range but bad in the other.

Note that, in Expressions (5) to (7), the case where the parameters δ ₁ and δ ₂ are determined using all four projection data is described, but this is an example and does not limit the present invention. . It goes without saying that the calculation may be performed using some projection data. However, in order to take advantage of changing the irradiation X-ray spectrum in a plurality of energy ranges, it is desirable to use projection data of a number larger than the number of energy ranges and larger than the number of types of irradiation X-ray spectra.

Next, returning to FIG. 6, the processing method of the multi-energy operation processing S703 will be described. This processing includes density image calculation processing S7031 by the density image calculation unit 1056 and multi-energy projection data calculation processing S7032 by the multi-energy projection data calculation unit 1057. The storage unit 109 stores numerical values for use in these processes, specifically, data of mass absorption coefficient values (μm ₁ (ε), μm ₂ (ε)) per energy of the reference material, and first and second data. The number of photons (S ₁ (ε), S ₂ (ε)) of the spectrum 2 is stored as calculation data 140.

  The spectrum data is, for example, the energy distribution of the number of X-ray photons irradiated at a tube voltage of 140 kV and 80 kV, and the calculation data 140 is prepared, for example, in a range from 20 keV to 140 keV for each 1 keV. However, these values are only examples and do not limit the present invention.

First, in the density image calculation processing S7031, the calculation data 140 stored in the storage unit 109 and the projection data obtained by performing the defect element correction S701 and the air correction S702 are used to calculate the equations (5) and (6). Te to calculate the [delta] ₁ and [delta] _2. In multi-energy projection data calculation process S7032 Next, using the calculated [delta] ₁ and [delta] _2, for example, to calculate the multi-energy projection data 144 of the object from the equation (4). At this time, the calculation data 140 stored in the storage unit 109 is used. The multi-energy projection data 144 is created, for example, every 1 keV. The image reconstruction unit 1053 performs reconstruction processing S704 on the multi-energy projection data 144, and creates a monochromatic X-ray equivalent image as the reconstructed image 145.

  By obtaining the dual energy image as described above, the accuracy can be improved as compared with the conventional dual energy CT device, and the accuracy of the dual energy image and the classification can be improved.

  The X-ray CT apparatus according to the present embodiment is not only superior to the conventional dual energy CT apparatus but also has an advantage compared to the conventional photon counting CT apparatus.

  As an example, by irradiating a suitable X-ray at the time of imaging with the second spectrum, it is possible to prevent the accuracy of the dual energy image from lowering. This is particularly effective when the subject is relatively large. In such a case, when photographing with the first spectrum, the attenuation at low energy becomes large, and the number of X-ray photons in the low energy range becomes significantly smaller than that in the high energy range. Dual-energy images created with only one have reduced accuracy in the low-energy range.

  However, in the X-ray CT apparatus of the present embodiment, when imaging in the second spectrum, particularly in the low energy range, or can be irradiated with X-rays only in the low energy range, the X-ray CT apparatus in the low energy range The projection data can be supplemented, and a decrease in the accuracy of the dual energy image in the low energy range can be prevented.

  The configuration and operation of the X-ray CT apparatus according to the present embodiment have been described above. However, the X-ray CT apparatus according to the present embodiment can be variously limited to the above-described configuration and operation. Hereinafter, modified examples will be exemplified.

  In the above embodiment, the defect correction S701 and the air correction S702 are performed as the correction processing, but these can be omitted as appropriate. For example, when there is no defective element, the defective element correction S701 does not need to be performed, and when the variation in the sensitivity of each X-ray detection element 400 is small, the air correction processing S702 may not be performed. That is, one or both of these correction processes need not be performed.

  Further, for example, other characteristics may be corrected. As such correction, for example, processing such as correction of the count number by pile-up or polarization is conceivable. Further, the correction processing of this embodiment was performed before the multi-energy calculation processing S703, but part or all of the correction processing was performed during or after the multi-energy calculation processing S703, or during or after the reconstruction processing S704. For example, the correction order may be different.

In the above embodiment, the case where two reference substances are used is described, but the number of reference substances is not particularly limited. For example, three or more reference substances may be provided. Here, it is desirable that the number of reference substances is larger than the number of types of spectra and larger than the number of energy ranges. At this time, δ _n (parameter obtained by integrating the density ρ _n (r) of the reference substance n by the path) described in the equation (2) needs to be determined only for the reference substance, and many parameters are determined. According to the embodiment, as shown in the equation (5), an equation can be set up by the product of the number of spectra and the number of energy ranges, so that there is an advantage that determination can be made with high accuracy even when there are many parameters.

  However, there may be a case where the equation is not necessary for the product of the number of spectra and the number of energy ranges. This can occur when the SNR of some projection data is low. In this case, it is desirable to use projection data of a number greater than the number of spectrum types and greater than the number of energy ranges, and to calculate using an equation equal to or greater than the number of reference substances. This is because in the calculation, it is necessary to determine indefinite parameters by the number of reference substances, but the equation can be determined more stably when there are more than this indefinite number of parameters.

  In the above embodiment, the case where two types of irradiation X-ray spectra are used is described, but this is merely an example, and it goes without saying that three or more X-ray spectra may be used. By using various types of X-ray spectra, the number of equations (5) can be increased, and the parameter determination accuracy can be improved. Further, as the X-ray spectrum, the case of the tube voltages of the X-ray tubes of 80 kV and 140 kV is described, but this is an example and does not limit the present invention. For example, one or both may use another tube voltage. Further, although a method of changing the tube voltage of the X-ray tube was used as a method of changing the X-ray spectrum, the present invention is not limited to this method.For example, an X-ray filter may be used, or the X-ray filter and the X-ray filter may be used. Both of the tubes may be used.

  Instead of separately irradiating X-rays of various types of spectra, X-rays that have been simultaneously irradiated may be separated into a plurality of different spectra by an X-ray detector to acquire data. As such an X-ray detector, for example, there may be a structure composed of multiple X-ray detection layers. In an example of such a structure, for example, there are two layers, a first layer close to the X-ray tube, and a separate second layer, and each layer can independently detect X-rays. A thin detection layer is provided. At this time, some X-rays can pass through the first layer and reach the second layer, but low-energy X-ray photons are more likely to be attenuated than high-energy photons, and therefore X-rays are higher than high-energy photons. Many are detected in the first layer near the tube.

  Therefore, X-rays having different X-ray spectra are detected in the first layer and the second layer. The data obtained in each layer as described above may be subjected to the processing of the present embodiment to create multi-energy projection data or a multi-energy image.

  Further, in the present embodiment, a case where a plurality of X-ray spectra are irradiated from one X-ray tube is described, but this is an example and does not limit the present invention. For example, an X-ray CT apparatus may have a plurality of X-ray tubes, and may irradiate X-rays individually or simultaneously to realize a plurality of X-ray spectra. Furthermore, even if the X-ray CT apparatus has a plurality of pairs of an energy-separation type X-ray detector and an X-ray tube, and individually or simultaneously, irradiates X-rays to realize a plurality of X-ray spectra. good.

In the above-described embodiment, the case where two rounds of imaging are performed and X-rays having different spectra are irradiated in each round is described, but the timing of changing the irradiation X-ray spectrum is not limited to each round of imaging.
For example, the spectrum of the irradiated X-ray may be changed at various timings, such as every half circle, every plural views, every view, and the like. Further, the number of acquired views and the number of revolutions of the two shootings may be different.

In the above embodiment, a monochromatic X-ray equivalent image of the reference substance is created as a multi-energy image, but the multi-energy image is not limited to a monochromatic X-ray equivalent image. For example, create and defined by the equation (2), the [delta] _n is a parameter density [rho _n the (r) integrated over the path of the reference substance n performs reconstruction processing as projection data, the reference material density image You may.

  In addition, a photoelectric effect image, a Compton scattering image, and an electron density image may be created by separating the photoelectric effect component and the Compton effect component by using a calculation technique in a conventional dual energy imaging method. Further, an absorption coefficient image under different imaging conditions may be created from an X-ray spectrum other than the one used in imaging and the calculated monochromatic X-ray equivalent image. Here, the imaging conditions may include a tube voltage, an X-ray filter, and the like. Further, various other multi-energy images may be used.

In the above embodiment, by using the projection data, the reference (the n 1 or 2) material n describing the method of obtaining the multi-energy images by determining the [delta] _n is a parameter density integrated over the path of, this Is an example and does not limit the present invention. There may be a method of obtaining other parameters using projection data. Further, there may be a method using a reconstructed image or a method not using a reference material. Further, there may be a case where both a method using projection data and a method using a reconstructed image are provided. Hereinafter, a description will be given of a modified example of the multi-energy arithmetic processing using data obtained by the X-ray CT apparatus of the present embodiment (data obtained in a plurality of energy ranges with a plurality of energy spectra).

<< Modification Example 1 of First Embodiment >>
In the first modification, a multi-energy image (here, a dual-energy image) is created by expressing the absorption coefficient of the subject using the absorption coefficient (linear attenuation coefficient) of the reconstructed image of the reference substance.
In the first modification, although not shown, the multi-energy calculating unit 1052 in FIG. 5 includes a reference substance density determining unit and a multi-energy image calculating unit.

  An example of the flow of the processing method according to the first modification will be described with reference to FIG. In FIG. 7, the same processes as those in FIG. 6 are denoted by the same reference numerals, and detailed description will be omitted. In the processing method illustrated in FIG. 7, first, for example, in the same manner as the processing in FIG. 6, the defect data correction S701 and the air correction S702 are performed on the projection data 143, and then the reconstruction processing S704 is performed first. For each of the second spectra, a reconstructed image (referred to as a conventional reconstructed image 146, distinct from the multi-energy reconstructed image) is obtained.

  Next, in a multi-energy calculation process S703, a reference material density determination process S7034 and a multi-energy image calculation process S7035 are performed.

  In the reference material density determination processing S7034, the density of the reference material is determined using a set of the absorption coefficient values of the conventional reconstruction image 146 acquired for each of the first and second spectra. Hereinafter, a processing method in the conventional dual energy imaging method (conventional method) will be described first to make it easier to understand the difference between the method of the present modified example and the conventional method.

The energy of the X-ray is E, the absorption coefficient (linear attenuation count) of the object obtained when imaging with the X-ray energy is μ (E), the reference substances are reference substances 1 and 2, and the X-ray energy is The mass absorption coefficient when photographed with E is μm _n (E) (n is an integer of 1 or 2 and indicates which reference substance), and if their density is c _n , the absorption coefficient μ ( E) can be written as equation (8) using the absorption coefficient and density of the reference material.

Equation (8) holds for each position of the subject. Therefore, it is necessary to determine the density c _n at each position of the subject, and in the conventional method, the density c _n is calculated using the absorption coefficient value of the reconstructed image measured by two X-ray spectra. Specifically, it is calculated using the simultaneous equations of equations (9-1) and (9-2) (collectively referred to as equation (9)) that can be derived by integrating equation (8) with each spectrum. . In equation (9), E ₁ means the X-ray energy when photographed at one energy of the two X-ray spectra, E ₂ means the X-ray energy when photographed at the other energy, and the left side Are the absorption coefficient values of the subject obtained at the respective tube voltages.

In Equation (9), μm _n (E _p ), which is the mass absorption coefficient of the reference substance n obtained at each energy _Ep (p is an integer of 1 or 2 and represents the type of X-ray energy), is based on literature values. and simulation by use of the X-ray spectrum and the energy ε mass absorption coefficient in (monochrome) when the X-ray energy E _p obtained by, can be determined by such simulation. Therefore, equation (9) can be solved, and the densities c ₁ and c ₂ of the reference material can be calculated.

  On the other hand, in the method of the present modified example, even when data is obtained with one spectrum, a measurement value obtained by separating the low energy range and the high energy range is obtained. Specifically, using the simultaneous equations of four equations (10-1) to (10-4) (collectively referred to as equation (10)) derived by integrating equation (8) with each spectrum. Can be calculated. In Expression (10), the left side is the absorption coefficient value of the subject obtained at each tube voltage.

Where l in E _pl is H or L, representing the energy range. The absorption coefficient μm _n (E _pl ) can be determined in the same manner as in the case of μm _n (E _p ) by performing a simulation with a limited energy range.

As described above, since the measurement value has noise, the density c _n (n is an integer of 1 or 2 and represents the type of the reference material) also has an error. However, in the method according to the present embodiment in which measurement is performed for each energy range, many simultaneous equations can be obtained as compared with the conventional calculation method using dual energy imaging. When solving the equations in consideration of the output value, the energy, and the like, it is possible to provide a weight for each equation, and the density c _n can be determined with high accuracy.

Here, the density c _n was determined using all the reconstructed images obtained by counting and reconstructing in the entire energy range of the entire X-ray spectrum, but this is an example, and is not intended to limit the present invention. Absent. It goes without saying that the calculation may be performed using some reconstructed images. However, in order to take advantage of changing the irradiation X-ray spectrum in a plurality of energy ranges, it is desirable to use a larger number of reconstructed images than the number of energy ranges and the number of types of irradiation X-ray spectra.

By performing c _n on all voxels in the reconstructed image as described above, values of all the voxels of the reference substance 1 and the reference substance 2 can be obtained.

Next, multi-energy image calculation processing S7035 is performed. Here, for example, a monochromatic X-ray equivalent image as a reconstructed image 145 of the reference material 1 is obtained by multiplying the density image c ₁ of the reference material by a mass absorption coefficient μm ₁ (ε) at an energy ε obtained from literature values and simulations. Create Similarly, a monochromatic X-ray equivalent image is created for the reference material 2 as well.

  The reconstructed image 145 created in this way is displayed on the display unit 106 (S705).

By performing the processing described above, in this processing, the density c _n (n is an integer of 1 or 2) can be determined more accurately than in the conventional calculation method in dual energy imaging, and the dual energy An image can be obtained.

In this modified example, the case where the product of the density and the absorption coefficient of the reference material is created as a reconstructed image 145 of the reference material as a dual energy image is described, but this is an example, and various images and data of There may be cases. For example, the dual energy image may be the density c _n (n is an integer of 1 or 2) of the reference substance 1 or the reference substance 2 in all voxels obtained by calculation, that is, a reference substance density image. Further, there may be a case of an effective atomic number image, an electron density image, a photoelectric effect image, a Compton scattering image, an absorption coefficient image in a spectrum other than the above-mentioned spectrum used in photographing, and the like, which are created using this density.

  In this modification, the case of two reference substances is described, but this is merely an example, and it goes without saying that three or more reference substances may be used.

<< Modification Example 2 of First Embodiment >>
In this modified example, a substance is identified using the value of the absorption coefficient of the reconstructed image obtained from the first and second spectra.

  In the known dual energy imaging, imaging is performed using the first and second spectra, and a substance is identified by comparing the absorption coefficient value of the obtained reconstructed image of the pair with the identification map. In the method according to the present modified example, in imaging in each spectrum, a reconstructed image is obtained separately for the low energy range and the high energy range, and identification is performed. In the second modification, although not shown, the multi-energy calculating unit 1052 in FIG. 5 includes a substance identifying unit and a multi-energy image calculating unit.

  Hereinafter, an example of a processing method according to the present modification will be described with reference to FIG. 8, the same processes as those in FIGS. 6 and 7 are denoted by the same reference numerals, and a detailed description thereof will be omitted.

  As shown in FIG. 8, for example, similarly to the first modification, first, the defect data correction S701, the air correction S702, and the reconstruction processing S704 are performed on the projection data 143, so that the first spectrum and the second spectrum are obtained. To create a conventional reconstructed image 146 acquired in the low energy range and the high energy range. Next, material identification S7036 is performed using the absorption coefficient value of the conventional reconstruction image 146. Here, the substance is identified by comparing both values of the absorption coefficient obtained by imaging with the first spectrum and the second spectrum with the identification map. The identification map maps a range of absorption coefficient values that a substance can take in a predetermined energy range, and can be created from data obtained based on experiments and simulations in advance.

  FIG. 9 shows an example of the identification map used in this modification. In these identification maps 166 and 167, the horizontal axis represents the absorption coefficient value in the first spectrum, and the vertical axis represents the absorption coefficient value in the second spectrum. The identification map 166 shows the low energy range of the X-ray detector. , And the identification map 167 is an identification map for values in the high energy range. The regions 160, 161 and 162 indicate regions of a combination of absorption coefficient values that can be taken by different substances in the low energy range. For example, the region 160 represents fat, the region 161 represents water, and the region 162 represents a contrast agent. However, these substances are only examples and do not limit the present invention. In the present embodiment, the case where the number of regions is three is described, but this is also an example, and it goes without saying that there may be two or four or more regions. Similarly, the regions 163, 164, and 165 indicate regions of combinations of absorption coefficient values that can be taken by different substances in the high energy range.

  By using such two kinds of identification maps, when the absorption coefficient values in the low energy range and the high energy range obtained by imaging with the first spectrum and the second spectrum in actual imaging are obtained, The substance can be identified by comparing which of the regions 160, 161, 162 and the regions 163, 164, 165 belong. Such identification is performed at each position of the reconstructed image.

  Here, different substances may be identified as a result of the low energy range and the high energy range. This is because the signal to be measured has noise due to fluctuations in the quantum number of X-rays, readout circuits, and external factors. In this case, for example, a result far from the boundary between the substances is adopted. For example, when the absorption coefficient value obtained for a certain voxel belongs to the region 160 in one identification map and is determined to be fat and belongs to the region 164 in another identification map and is determined to be water, the identification coefficient of the absorption coefficient value is determined. The distance from the boundary between the region 160 and the region 161 in the map 166 and the distance from the boundary between the region 163 and the region 164 in the identification map 167 are compared, and the longer distance is determined as the region to which the absorption coefficient value belongs.

  As described above, in the method of the present modified example, the substance is identified using the absorption coefficient of the product of the number of spectra and the number of energy ranges (four), so in the conventional method using two spectra, the measurement result is identified. Even in the case where the identification accuracy of a substance is reduced at the boundary of the map, the identification accuracy can be improved by adopting a result farther from the boundary of the identification map. The identification accuracy can be improved by using more absorption coefficients.

  Next, in the multi-energy image calculation processing S7035, a reconstructed image 145 is created by adding the result of the identified substance to the conventional reconstructed image 146. The result of the substance identified here is displayed in the conventional reconstructed image 146, for example, by color coding. Finally, display S705 of the reconstructed image 145 created in this way is performed.

  As described above, according to the present modification, a reconstructed image 145 in which a substance is identified with higher accuracy than in a conventional dual energy CT apparatus can be created.

  Note that, in this modified example, a case is described in which, when different identification results are obtained between the low energy range and the high energy range, a result apart from the boundary between substances is adopted, but this is an example, and the present invention is not limited to this. It is not limited. As another method, for example, when using the one with smaller data noise, or when using the low energy range and the high energy range obtained from one spectrum, four methods of the low energy range and the high energy range obtained in each spectrum are used. There may be various cases such as identification using only three of the measurement results. However, in order to take advantage of changing the irradiation X-ray spectrum in a plurality of energy ranges, it is desirable to use a larger number of reconstructed images than the number of energy ranges and the number of types of irradiation X-ray spectra.

  In the present modified example, the case where the low energy range, the high energy range, and the identification map are provided is described, but this is an example and does not limit the present invention. There may be a case where the identification is provided for each spectrum. Although the case where the identification map is composed of two axes is described, an identification map having three or more axes may be used, and the identification map having four indices as axes has a low energy range obtained for each spectrum. And four measurement results in the high energy range may be identified at once. If the number of energy ranges or the number of spectra further increases, the substance may be identified using an identification map having more axes, instead of being limited to four.

  In this modified example, the case where two X-ray spectra and two energy ranges are used is described, but this is merely an example, and it goes without saying that one or both of them may be three or more. At this time, it is possible to perform discrimination by using at most the number of products of the number of spectra and the number of energy ranges as indices.

<< Modification Example 3 of First Embodiment >>
In this modified example, photoelectric effect, Compton scattering image, coherent scattering, Create a scatter image.

  In conventional dual energy imaging, imaging is performed with the first and second spectra, fitting is performed on the absorption coefficient values of the obtained reconstructed image of the pair, and the physical factors of absorption are separated, and the photoelectric effect image is obtained. Create Compton scatter images and coherent scatter images. In this modified example, in imaging in each spectrum, a reconstructed image is obtained separately in a low energy range and a high energy range, and physical factors are classified. The configuration of the device and the configuration of the calculation unit of this modification are the same as those shown in FIGS. 1 and 5, but a physical factor classification unit (not shown) is added to the calculation unit 105.

  Hereinafter, an example of the process of the multi-energy calculating unit 1052 of the present modified example will be described with reference to FIG. In FIG. 10, the same processes as those in FIG. 8 are denoted by the same reference numerals, and detailed description will be omitted.

  As shown in FIG. 10, in the processing method of the present modification, similarly to the first modification, first, the defect data correction S701, the air correction S702, and the reconstruction processing S704 are performed on the projection data 143, and the first spectrum is obtained. Then, a conventional reconstructed image 146 obtained by photographing with the second spectrum and acquiring in the low energy range and the high energy range is created. Next, by performing fitting on the absorption coefficient value of the conventional reconstructed image 146, physical factor classification S7037 is performed.

  The fitting function used in the physical factor classification S7037 can be represented by, for example, Expression (11). Equation (11) represents the absorption coefficient μ (ε) at the energy ε as a function of physical factors, where f, g, and h are the physical factors of absorption, the photoelectric effect, Compton scattering, It represents the cross-sectional area of coherent scattering, respectively.

  Here, Z represents the effective atomic number, and ρ represents the electron density. Further, f, g, and h are functions of the energy ε but have different dependencies, and f and h are functions of the effective atomic number Z, but have different dependencies.

  By performing imaging with irradiation X-rays of the first and second two energy spectra, as shown in equations (12-1) and (12-2) (hereinafter collectively referred to as equation (12)), two An equation for the absorption coefficient is obtained.

Here, since the electron density ρ does not depend on the energy, μ ₁ (ε) / μ ₂ (ε) can be written as Expression (13).

  In the physical factor classification S7037, Expression (13) is used as a fitting function. Thereby, the effective atomic weight Z can be determined. Further, the electron density ρ can be determined by using the equation (12).

At the time of fitting, in the case of conventional dual energy imaging, one equation (13) (μ ₁ (ε) / μ ₂ (ε)) will be used.In this modification, for example, a low energy range and a high energy Equation (13) can be established for each of the ranges, and the fitting accuracy can be improved by using _two μ ₁ (ε) / μ ₂ (ε). Furthermore, even when one μ ₁ (ε) / μ ₂ (ε) is used, it is possible to use a value obtained in an energy range where the variation between μ ₁ (ε) and μ ₂ (ε) is small, and to use low energy and high energy. By using a value having a large difference between the absorption coefficient values of the energy μ ₁ (ε) and μ ₂ (ε), it is also possible to improve the fitting accuracy.

  In the next multi-energy image calculation process S7035, the photoelectric effect, Compton scattering, and coherent scattering components are separated by using Expression (11) and Expression (12), and the photoelectric effect at energy ε is obtained as a reconstructed image 145. An image, a Compton scattered image, a coherent scattered image, and a photoelectric effect image, a Compton scattered image, and a coherent scattered image when irradiating the first and second spectra and irradiating other spectra are created. Furthermore, an effective atomic weight image and an electron density image can be created using the calculated effective atomic weight Z and electron density ρ. Finally, display S705 of the reconstructed image 145 created in this way is performed.

In the present modification, the physical factor discrimination S7037 is performed using the fitting function of Expression (13), but this is only an example and does not limit the present invention. For example, as in the first embodiment, a monochromatic X-ray equivalent image may be created using a reference material, and physical factor classification S7037 may be performed on the absorption coefficient using Expression (11).

<Second embodiment>
The X-ray CT apparatus of the present embodiment, in addition to the configuration of the first embodiment, of a plurality of imaging performed respectively with a plurality of irradiation energy spectra, an imaging condition determination unit (1060) that determines imaging conditions of at least one imaging. It is a feature that it is further provided. The imaging condition determination unit uses the projection data of the first imaging performed using the first irradiation X-ray energy spectrum, and performs the processing using a second irradiation energy spectrum different from the first irradiation energy spectrum. The shooting condition (second shooting condition) of the second shooting to be performed is determined.

  The imaging conditions are conditions for determining the spectrum and dose of the irradiated X-ray, and include, for example, a tube voltage value, the presence or absence and thickness of an X-ray filter, a tube current, a view time, and an imaging time. Thereby, the second imaging condition can be optimized according to the subject, and the accuracy of the dual energy image can be improved.

  Here, the irradiation energy spectrum is different between the first and second imagings, but it is preferable that the irradiation energy spectrum of the second imaging be lower than the irradiation energy spectrum of the first imaging. For example, when the tube voltages of the two imagings are different, it is desirable that the second imaging in which the imaging conditions are changed is imaging with a lower tube voltage.

  This is because X-ray photons with low energy are more likely to be absorbed depending on the subject, and the number of X-ray photons tends to be excessive or deficient. Is large. For the same reason, it is highly likely that the shortage in the first imaging is caused by low-energy X-ray photons, so it is more efficient to irradiate with a low tube voltage and add low-energy X-ray photons. Because it is a target.

  The overall configuration of the X-ray CT apparatus according to the present embodiment is the same as that in FIG. 1, but an imaging condition determination unit 1060 is added to the calculation unit 105 as shown in FIG. FIG. 11 shows a case where the spectrum is changed as the second imaging condition, and the imaging condition determination unit 1060 calculates the ratio of the number of X-ray photons detected in the low energy range and the X-ray photon number detected in the high energy range, respectively. An X-ray photon number ratio calculation unit 1061 and a spectrum condition determination unit 1062 for determining an irradiation X-ray spectrum in the second imaging based on the X-ray photon number ratio calculated by the X-ray photon ratio calculation unit 1061 are provided.

  The main control unit 1050 controls the X-ray photon number ratio calculation unit 1061 and the spectrum condition determination unit 1062. Note that, in FIG. 11, the imaging condition determination unit 1060 is shown as a part of the multi-energy operation unit 1052, but may be a functional unit different from the multi-energy operation unit 1052. The imaging conditions for changing the spectrum include, for example, the value of the tube voltage, the presence or absence of an X-ray filter, the thickness, and various other means for changing the X-ray spectrum.

  Hereinafter, the processing procedure and control flow of the imaging condition determination unit 1060 will be described with reference to FIG.

  First, the photographer inputs, to the input unit 110, the photographing conditions for the first round for performing the first photographing and the photographing conditions for the second round for performing the second photographing. At this time, the conditions of the tube voltage and the X-ray filter when imaging with the second spectrum in the second round are not input or are provisional even if they are input, and are determined by the imaging condition determination unit 1060 Use something.

  When the photographer starts photographing with the input unit 110, the control unit 107 controls the first photographing S711 in the first round of photographing. The photographing conditions at this time are the conditions previously input by the photographer from the input unit 110. Next, the X-ray photon number ratio calculation unit 1061 calculates the ratio of the average number of X-ray photons in the low energy range and the high energy range using the projection data obtained in the first round (S712). Next, the second spectrum condition is determined by the spectrum condition determination unit 1062 using this ratio (S713).

  Here, the second spectrum condition is determined based on, for example, the imaging condition of the tube voltage. As a determination method, for example, the relationship between the ratio of the number of X-ray photons and the imaging condition is determined in advance as the parameter 147, stored in the storage unit 109, and used. The relationship between the ratio of the number of X-ray photons and the imaging conditions is, for example, such that when the ratio of the number of X-ray photons is small, the tube voltage is reduced, and when the ratio of the number of X-ray photons is large, the tube voltage is increased. .

Next, the spectrum condition determination unit 1062 passes the photographing conditions to the control unit 107. The control unit 107 controls the spectrum changing unit 111 to change the spectrum S714, and then performs the second round of photographing S715. Using the measurement data obtained in the imaging of the first lap (first imaging) and the imaging of the second lap (second imaging), the same multi-energy as the processing described in the first embodiment or its modified example Operation processing S703 is performed.

  By performing such control, X-ray photons in the energy range insufficient for X-rays of the first spectrum can be supplemented by X-rays of the second spectrum. This makes it possible to improve the SNR (Signal-to-Noise Ratio) and CNR (Contrast-to-Noise Ratio) of the data, thereby improving the accuracy of the multi-energy image. Further, by controlling the second spectrum, it is possible to reduce the amount of exposure while maintaining the image quality of the multi-energy image.

  Further, the present embodiment is also useful for improving the accuracy of a photon counting image. For example, one of the photon counting imaging methods is K-edge imaging. In K-edge imaging, a low-energy range and a high-energy range are set using the energy of the K-edge of a certain substance as a threshold, and a K-edge image is obtained by subtracting reconstructed images obtained in each range. In this case, for example, the SNR or CNR of the K-edge image may be reduced such that the number of X-ray photons in the low energy range is significantly smaller than the number of photons in the high energy range due to attenuation in the subject.

  According to the present embodiment, the X-ray photons in the energy range deficient in the X-rays of the first spectrum, the X-rays of the first spectrum are set by setting the second imaging conditions so as to supplement the X-rays in the second spectrum. SNR and CNR can be improved as compared with the case where K-edge imaging is performed using only lines.

  In the processing illustrated in FIG. 12, the case where the irradiation X-ray spectrum is changed is described as the second imaging condition. However, as described above, the irradiation X-ray dose may also occur. FIG. 13 shows an example of a control flow for realizing this. In order to realize the processing in FIG. 13, the spectrum condition determining unit 1062 shown in FIG. 11 is replaced with a dose condition determining unit (not shown) that determines the dose (second dose) in the second imaging. Similarly, the determination S713 of the second spectrum condition in FIG. 12 is replaced with the determination S716 of the second dose condition that determines the imaging condition for realizing the second dose.

  In this process, the ratio of the average number of X-ray photons in the low energy range and the high energy range is calculated (S712), and the dose for the second imaging is determined (S716). The determination method is the same as in the case of the spectrum change, for example, the relationship between the X-ray photon number ratio and the dose determined in advance as a parameter 147, is determined in advance and stored in the storage unit 109, This can be used.

  For example, if the ratio of the average number of X-ray photons in the low-energy range and the high-energy range is very small, that is, if the number of photons in the low-energy range is relatively small, in the second imaging performed at low energy, the dose To increase. On the other hand, when the ratio is large, that is, when the number of photons in the low energy range is relatively large, the dose is reduced.

  Next, the dose of the second imaging is changed to the determined dose (S717). The imaging conditions for realizing the second dose include, for example, tube current, view time, and imaging time. Performing the second imaging with the set dose (S715) and performing the same multi-energy arithmetic processing S703 as in the first embodiment using the measurement data obtained in the first imaging and the second imaging, This is the same as the flow in FIG.

  By controlling such doses, it is possible to improve the SNR and CNR of data by supplementing X-ray photons in the energy range insufficient for X-rays in the first spectrum with X-rays in the second spectrum. And the accuracy of the dual energy image can be improved. In addition, unnecessary irradiation of the subject with X-rays can be prevented, and the amount of invalid exposure can be reduced.

  Both the processing shown in FIG. 12 and the processing shown in FIG. 13 may be performed to control both the X-ray spectrum and the dose in the second imaging.

  The illustrated embodiment is an example of the present embodiment, and the X-ray CT apparatus of the present embodiment is used in multi-energy imaging in which a plurality of energy spectrum distributions are separately measured for a low energy range and a high energy range. It is characterized by having a function of determining imaging conditions based on multi-energy measurement data, and a method for determining imaging conditions can be variously changed.

  For example, in the above description, the second spectrum is determined using the ratio of the average number of X-ray photons, but this is merely an example and does not limit the present invention. For example, the number of X-ray photons in one of the low energy range and the high energy range may be used. Further, even when the number of X-ray photons in both energy ranges is used, there can be various usage methods other than the ratio, such as a method using a sum or an absolute value of each.

  Further, in the present embodiment, the second spectrum is determined using the result of the first imaging, but this is an example and does not limit the present invention. For example, the determination may be performed by performing a pre-scan before the actual shooting. Further, the first spectrum and its dose may be determined. By controlling both the first spectrum and the second spectrum in this way, the spectrum of the X-ray can be further optimized, and the image quality can be improved with high accuracy and the exposure can be reduced.

  Further, in the present embodiment, the case where the first and second shootings are performed in the first and second laps, respectively, is described, but this is an example and does not limit the present invention. The first and second shootings may be performed a plurality of times, or may be performed for a half time. Furthermore, the first and second shootings may be performed a plurality of times by switching for each single or multiple views. At this time, the second imaging condition may be determined using the result of the first first imaging, or may be determined based on the result of the immediately preceding first imaging.

  In the present embodiment, the second imaging condition is determined using the ratio of the average number of X-ray photons in the low energy range and the high energy range, but this is an example and does not limit the present invention. For example, the projection data of one of the low energy range and the high energy range may be used. As an example, there is a case where the imaging condition is determined by comparing the count of the low energy range with a threshold.

  There is also a case where the values of the projection data in the low energy range and the values of the projection data in the high energy range are used without being compared. Further, when there are three or more energy ranges, there may be various cases in which one or a plurality of the projection data is used. At this time, since the reduction of the number of photons in the low energy range may be particularly large depending on the subject, it is desirable to use projection data in the low energy range.

<Third embodiment>
In the first embodiment, the energy distribution of X-rays detected by the X-ray detector in each energy range (hereinafter, referred to as detected X-ray energy distribution) was changed by switching a plurality of types of irradiation X-ray spectra, In the present embodiment, switching of the detected X-ray energy distribution is realized by dynamically changing the energy threshold that divides the energy range in imaging.

That is, the X-ray CT apparatus of the present embodiment includes, as a detected energy distribution changing unit, an energy threshold changing unit that changes an energy threshold that determines the energy range of the X-ray detector, and changes the energy threshold by the energy threshold changing unit. Change the detected X-ray energy distribution.
The multi-energy calculating unit creates multi-energy projection data by using a plurality of projection data acquired when the energy thresholds are different.

  FIG. 14 shows an example of the configuration of the X-ray CT apparatus of the present embodiment. 14, the same elements as those of FIG. 1 are denoted by the same reference numerals, and a detailed description will be omitted. As illustrated, the X-ray CT apparatus according to the present embodiment includes an energy threshold changing unit 112.

  The energy threshold changing unit 112 changes, for example, a threshold (126 in FIG. 4) when the signal collecting unit 108 discriminates between the low energy range and the high energy range. The threshold can be changed by changing the comparison voltage of one comparator or, in the case of a digital comparator, by switching the comparison value. An example of the function of the energy threshold changing unit 112 will be described with reference to FIG.

  In FIG. 15, a range 130 represents a low energy range, and a range 131 represents a high energy range. The energy threshold changing unit 112 switches an energy threshold for separating a low energy range and a high energy range between the first imaging and the second imaging. That is, in the first imaging, the energy threshold is set to the threshold 132, and in the second imaging, the energy threshold is set to the threshold 133. After the energy threshold changing unit 112 switches the energy threshold between the threshold 132 and the threshold 133, the X-ray detector 104 classifies the X-ray photons into respective energy ranges and performs coefficients.

  The calculation unit 105 performs a multi-energy calculation using projection data obtained by two imagings having different energy thresholds. FIG. 16 shows an example of the configuration of the calculation unit 105 of the present embodiment. In FIG. 16, the same elements as those in FIG. 5 showing the configuration example of the calculation unit 105 of the first embodiment are denoted by the same reference numerals, and detailed description will be omitted.

  As shown, the calculation unit 105 includes a main control unit 1050, a correction processing unit 1051, a multi-energy calculation unit 1052, and an image reconstruction unit 1053, and the correction processing unit 1051, the multi-energy calculation unit 1052, and the image reconstruction unit 1053. Operates under the control of the main control unit 1050.

Although omitted in FIG. 16, the correction processing unit 1051 includes a defective element correction unit, an air correction unit, and the like, as in the first embodiment. Similarly to the first embodiment, the multi-energy calculation unit 1052 includes a density image calculation unit 1056 and a multi-energy projection data calculation unit 1057 , but further includes a narrow energy range data calculation unit 1058. The narrow energy range data calculation unit 1058 calculates the projection data of the area indicated by the area 134 in FIG. 15 using the projection data obtained by the first imaging and the projection data obtained by the second imaging.

  Hereinafter, the flow of the processing of the arithmetic unit will be described together with the shooting procedure. Also in the present embodiment, a case will be described in which two rounds of shooting are performed, and the first round is performed in the first round and the second round is performed in the second round. According to the first embodiment, the spectrum changing unit 111 performs the shooting by switching the shooting in the first spectrum in the first cycle and the shooting in the second spectrum in the second cycle. After the imaging, the energy threshold is changed by the energy threshold changing unit 112, and the second round imaging is performed with an energy threshold different from the first imaging.

  The calculation unit 105 creates a multi-energy image using the projection data obtained in the first and second imagings. FIG. 17 shows an example of the processing flow of the calculation unit. In FIG. 17, the same processes as those in FIG. 6 are denoted by the same reference numerals, and detailed description will be omitted.

  As shown in FIG. 17, first, correction processing such as defect element correction S701 and air correction S702 is performed on the projection data 143. After the correction processing, the narrow energy range data calculation unit 1058 performs the narrow energy range data calculation processing S7033 to create multi-energy projection data. In this processing, projection data of a region (hereinafter, referred to as a narrow energy range) 134 sandwiched between the energy threshold 132 and the energy threshold 133 is created using the imaging results at different energy thresholds.

Specifically, the number of X-ray photons in the low energy range obtained in the first imaging is X _L1 , the number of X-ray photons in the high energy range is X _H1 , and the number of X-rays in the low energy range obtained in the second imaging is X _H1 . Assuming that the number of X-ray photons is X _L2 and the number of X-ray photons in the high energy range is X _H2 , the number of X-ray photons X _{LH in the} narrow energy range 134 is obtained by the equation (14-1) or (14-2). Can be

但し、第2の撮影のエネルギー閾値133＞第1の撮影のエネルギー閾値132とする。

However, it is assumed that the second imaging energy threshold 133> the first imaging energy threshold 132.

  The narrow energy range data calculation unit 1058 performs the above difference calculation to create projection data of the narrow energy range 134, the low energy range 135, and the high energy range 136, that is, multi-energy projection data 144.

  The narrow energy range data calculation processing S7033 is performed using one or both of the equation (14-1) and the equation (14-2) .To determine the multi-energy projection data 144 with higher accuracy, It is desirable to compare the noise of the first imaging and the second imaging and use a combination that suppresses the noise of the projection data 143 in the narrow energy range obtained by the difference.

  Specifically, the noise of the projection data 143 of the low energy range 130 obtained in the first imaging and the noise of the projection data 143 of the high energy range 131 obtained in the second imaging are compared, and the low noise obtained in the first imaging is compared. When the projection data 143 in the energy range 130 is lower, the projection data 143 in the low energy range 130 is used. On the other hand, when the noise value of the projection data 143 in the high energy range 131 obtained by the second imaging is lower, the projection data 143 in the high energy range 131 is used.

  The reason is as follows. Although the multi-energy projection data 144 in the narrow energy range 134 is calculated by subtracting the projection data 143 obtained in the first imaging and the second imaging, the count number of only the narrow energy range 134 can be calculated as the signal level. On the other hand, as for the noise level, noise in the entire energy range used remains.

  As shown in FIG. 15, the energy range remaining as noise includes a low energy range 135 and a narrow energy range 134 when using projection data in a low energy range, and a high energy range when using projection data in a high energy range. There is an energy range 136 and a narrow energy range 134. Here, since the narrow energy range 134 is common, the other low energy range 135 and high energy range 136 remain. Therefore, by comparing the projection data 143 in the low energy range 135 and the projection data 143 in the high energy range 136 and using the smaller one, it is possible to reduce the noise in the multi-energy projection data 144 in the narrow energy range 134.

  Therefore, instead of comparing the noise of the projection data 143 of the low energy range 130 obtained in the first imaging with the noise of the projection data 143 of the high energy range 131 obtained in the second imaging, the noise of the narrow energy range 134 is added thereto. In addition, the projection data 143 of the low energy range 130 obtained in the second imaging and the noise of the projection data 143 of the high energy range 131 obtained in the first imaging are compared, and the low energy range obtained in the second imaging is compared. When the noise of the projection data 143 of 130 is smaller, the projection data 143 of the low energy range 130 obtained in the first and second imagings is converted to the projection data 143 of the high energy range 131 obtained in the first imaging. When the noise is smaller, the same effect can be obtained by selecting the projection data 143 in the high energy range 131 obtained in the first and second imaging.

  Further, instead of the noise of the projection data 143 in the low energy range 130 and the high energy range 131, one having a smaller output value (the number of X-ray photons) may be used. This is because the X-ray quantum noise decreases as the number of X-ray photons decreases.

  Next, a reconstruction process S704 is performed on the multi-energy projection data 144 obtained in this way, and a reconstructed image 145 using X-ray photons in the narrow energy range 134 is created. Reconstructed images can be similarly created for the low energy range 135 and the high energy range 136.

  In the above description, the multi-energy projection data 144 in the narrow energy range 134 is created by switching to the two energy thresholds 132 and 133, but may be switched to three or more energy thresholds, for example. At this time, the first to the Mth (M is an integer of 3 or more) energy thresholds are switched in the order of energy, and the Nth (N is an integer of 1 or more and less than M) energy thresholds and (N + 1) By subtracting the measurement results of the energy thresholds of the above, multi-energy projection data 144 of (M−1) narrow energy ranges can be created.

  According to the present embodiment, a reconstructed image in an arbitrary narrow energy range sandwiched between two energy thresholds can be obtained. Further, by making the energy threshold value 132 and the threshold value 133 close to each other and setting the narrow energy range 134 to a very narrow range, a monochromatic X-ray equivalent image can be obtained.

  In addition, in the conventional photon counting CT apparatus, for example, when obtaining a multi-energy image separated into three energy ranges, two comparators are required.According to the present embodiment, one comparator is used. X-ray photons in one energy range can be counted, and three projection data and a reconstructed image can be obtained. Therefore, the circuit scale for the comparator can be reduced as compared with the conventional device, and accordingly, the power consumption of the circuit can be reduced and the increase in the price of the device can be suppressed. Reducing the power consumption of this circuit leads to extending the life of the device.

  However, in the present embodiment, it is not limited to switching one comparator to create the multi-energy projection data 144 in the narrow energy range 134, and a plurality of comparators may be used. For example, two energy thresholds may be realized by two comparators. At this time, the circuit scale of the comparator is increased as compared with the case where switching is performed by one comparator, but the number of read circuits after the comparator can be reduced. Further, a plurality of comparators may be provided, and a plurality of energy thresholds may be switched and read out.

  Although the third embodiment has been described with reference to FIGS. 15 to 17, the present embodiment is not limited to the configuration and operation described above, and various modifications are possible. For example, the case where two shootings are performed by two rounds of shooting has been described. However, this is an example, and a scan range for performing two shootings may be various, such as a plurality of rounds, a single round, and a half round. Instead of switching between the first and second shootings in one round, switching may be performed in a plurality of rounds, or may be performed in a single view, in a plurality of views, or in a half round. Further, the switching may be performed not only once but also a plurality of times.

  However, in these cases, the first and second imagings may not be signals integrated along the same path. This is because, for example, when the first view and the second view are switched between the odd view and the even view and the shooting is performed in one round, the projection directions obtained in the first and second shooting are different.

  In this case, after estimating the count value from the signal value of the adjacent view, it is better to create the multi-energy projection data 144 in the narrow energy range by performing the count obtained in the other imaging and the narrow energy range data calculation processing S7033. good. This is because, for example, the multi-energy projection data 144 in the narrow energy range of the second view estimates the value of the second view from the count obtained in the first imaging of the first view and the third view, and actually calculates the value of the second view. Means to create by taking the difference from the count obtained in the second imaging.

  In the present embodiment, the description has been made on the assumption that the same energy threshold is switched in all the X-ray detection elements 400.However, this is an example, and for example, only some of the X-ray detection elements 400 perform. Is also good. Further, different energy thresholds may be switched for each of the X-ray detection elements 400 and groups thereof. At this time, the switching of the different energy thresholds may or may not be synchronized.

  Further, in the present embodiment, the case where the energy threshold changing unit 112 switches the energy threshold (126 in FIG. 4) for separating the low energy range and the high energy range is described, but this is an example, and the minimum The energy value (127 in FIG. 4, 127) or the maximum energy value may be used.

<Fourth embodiment>
This embodiment is an embodiment in which the X-ray CT apparatus of the third embodiment is applied to K-edge imaging.

  The apparatus configuration of the present embodiment is the same as the X-ray CT apparatus of the third embodiment, and includes an energy threshold changing unit 112. As shown in FIG. 18, the calculation unit 105 includes a correction processing unit 1051, a multi-energy calculation unit 1052 including a narrow energy range data calculation unit 1058, and an image reconstruction unit 1053, and a K edge image for creating a K edge image. A calculation unit 1059 is provided. The multi-energy operation unit 1052 creates multi-energy projection data by multi-energy operation and creates a multi-energy image as in the third embodiment. The K-edge image calculation unit 1059 creates a K-edge image by subtracting the two reconstructed images before and after the K-edge energy.

  First, an example of how to set an energy threshold by the energy threshold changing unit 112 will be described with reference to FIG. In the example shown in the figure, the first shooting is performed on the first lap, the second shooting is performed on the second lap, and the third shooting is performed on the third lap. The energy threshold changing unit 112 distinguishes between the low energy range and the high energy range using the energy threshold, at the energy threshold 132 in the first imaging, at the energy threshold 133 in the second imaging, and in the third imaging. This is performed at the energy threshold 138. The height of the threshold value is in a relationship of threshold value 132 <threshold value 133 <threshold value 138, and the energy threshold value 133 is set to be similar to the energy of the K edge of the metal on which the K edge imaging is to be performed.

  Next, data processing performed by the arithmetic unit 105 of the present embodiment will be described with reference to the flow of FIG. As illustrated in FIG. 20, the calculation unit 105 performs, for the respective projection data 143 obtained in the first imaging, the second imaging, and the third imaging, for example, as in the third embodiment, Defective element correction S701 and air correction S702 are performed.

  Next, the multi-energy calculation unit 1052 performs a narrow-energy image calculation process S7033. In this process, the number of X-ray photons in the narrow energy range 134 (hereinafter, referred to as the first narrow energy range) is calculated from the first imaging and the second imaging, and the third imaging and the second imaging are performed. From this, the number of X-ray photons in the narrow energy range 137 (hereinafter referred to as the second narrow energy range) is calculated.

  At this time, similarly to the case of the third embodiment, the measurement results of one or both of the low energy range and the high energy range can be used, but as described in the third embodiment, the result of the low noise level is obtained. It is desirable to use. That is, when calculating the first narrow energy range, the projection data obtained in the low energy range 130 obtained in the first imaging and the projection data obtained in the high energy range 131 obtained in the second imaging. Compare the noise level, if the projection data obtained in the low energy range 130 obtained in the first imaging is smaller, use the projection data obtained in the low energy range 130 obtained in the first and second imaging. If the projection data obtained in the high energy range 131 obtained in the second imaging is smaller, the projection data obtained in the high energy range 131 obtained in the first and second imagings is used.

  Similarly, when calculating the second narrow energy range, projection data obtained in the low energy range 130 obtained in the second imaging, and projection data obtained in the high energy range 131 obtained in the third imaging. Comparing the noise level of the projection data obtained in the low energy range 130 obtained in the second imaging, if the projection data obtained in the low energy range 130 obtained in the second imaging is smaller, If the projection data obtained in the high energy range 131 obtained in the third imaging is smaller than the projection data obtained in the third imaging, the projection data obtained in the high energy range 131 obtained in the second and third imaging is used.

  Next, in the reconstruction processing S704, a reconstruction operation is performed on each of the first and second narrow energy range projection data obtained as the multi-energy projection data 144 to obtain a reconstructed image.

  Next, in K edge image calculation processing S7041, the reconstructed images in the first and second narrow energy ranges 134 and 137 are subtracted to obtain a K edge image. With this K-edge image, an image having only a target substance (hereinafter, referred to as a target substance) having a high contrast can be obtained. This principle will be described with reference to FIG.

  In FIG. 21, the horizontal axis represents energy, the vertical axis represents the value of the mass absorption coefficient at that energy, the curve 150 represents the mass absorption coefficient of the target substance to be subjected to K-edge imaging, and the energy 151 represents the K-edge energy. That is, the K-edge energy is the energy at which the mass absorption coefficient changes abruptly, and the energy at which the X-ray absorption rate changes abruptly according to the energy.

The K-edge energy 151 of the target substance is set as the second energy threshold 133, and before and after that, the first narrow energy range 134 and the second narrow energy range 137 are realized, and the number of X-ray photons in that range is measured. In this case, the X-rays are easily transmitted in the first narrow energy range 134 lower than the K edge energy, whereas the X-rays are hardly transmitted in the second narrow energy range 137 higher than the K edge energy. Therefore, when the projection data of the second narrow energy range 137 is subtracted from the projection data of the first narrow energy range 134, a high contrast can be obtained for the target substance.

  On the other hand, for substances other than the target substance, the mass absorption coefficient does not change significantly between the first narrow energy range 134 and the second narrow energy range 137, so that a difference between the projection data results in a very small contrast. Therefore, in the K-edge image, only the target substance can be extracted with high contrast.

  In the above description, the second energy threshold 133 is set to match the K-edge energy of the target substance.However, an error (hereinafter referred to as energy In this case, it is desirable to set the energy threshold 133 so that the energy is slightly higher than the K edge as shown in FIG.

  This is because even if the X-ray photons with energy equal to or higher than the energy threshold are actually measured as being equal to or lower than the energy threshold due to the energy resolution error, the number of X-ray photons below the energy threshold is much larger than the X-ray photons above the energy threshold. Although the amount of counting is relatively small, if the X-ray photons of energy below the energy threshold are actually measured above the energy threshold due to energy resolution errors, the amount of incorrect counting will be relatively large. It is because.

  By setting the energy threshold to a value slightly larger than the K edge energy, the possibility of erroneously measuring an X-ray having a K edge energy or less as an energy threshold or more due to such a resolution error can be reduced. The degree to which the energy threshold is set to a value larger than the K edge energy is, for example, an energy determination error of the X-ray detection element 400. This energy determination error can be measured in advance depending on variations in the signal amount and output voltage value in the detection layer 401, a measurement error in the reading circuit 405, and the like.

  According to the present embodiment, the K-edge imaging of the target substance can be performed with high accuracy by appropriately changing the energy threshold by the energy threshold changing unit 112 and performing imaging. Also, compared to the case where a comparator is provided for each of the first, second, and third energy thresholds to separate them, the circuit scale of the comparator can be reduced, and the same effect as that described in the third embodiment can be obtained. The effect is obtained.

  In the present embodiment, counting of two narrow energy ranges is calculated using three energy thresholds, but this is an example, and using four or more energy thresholds, three or more narrow energy ranges are calculated. The count may be calculated. Further, a comparator that switches the energy threshold and a comparator that does not switch the energy threshold may be mixed, and each energy range may be counted to calculate the count of the narrow energy range. Alternatively, a plurality of comparators may switch between a plurality of different energy thresholds to count each energy range and calculate a count for a narrow energy range.

  Further, in the present embodiment, the projection data of the two narrow energy ranges 134 and 137 are created for calculating the K-edge image, but one of them may be substituted with the projection data of the actually counted energy range. For example, when calculating the K edge image, the projection data of the high energy range 131 obtained by the second imaging can be used instead of the projection data of the high narrow energy range 137 shown in FIG. When this method is adopted, the third imaging in FIG. 19 is unnecessary, and only two energy thresholds can be switched, so that the imaging can be speeded up. Such a method is useful especially when the actually counted energy range is narrow, such as when the high energy range 131 is narrow.

  Further, also in the present embodiment, as described in the third embodiment, the switching between the first imaging, the second imaging, and the third imaging is not limited to every lap, but every plural laps, half a lap. There may be various cases, such as every one, every plural views, every one view, etc.

  Further, in the present embodiment, the description has been made on the assumption that the imaging is performed by switching the energy range in synchronization with one view by all the X-ray detection elements, but this is an example and does not limit the present invention. For example, switching may be performed for each of a plurality of views. Alternatively, the detection may be performed by only some of the X-ray detection elements. Furthermore, the energy range in which images are taken by the X-ray detection element may be different.

  This is, for example, the projection data of the K-edge imaging in one group (group A) X-ray detection elements, the projection data other than the K-edge imaging in the other group (group B) X-ray detection elements, respectively You may get it. Specifically, for example, in the group A X-ray detection element, the energy thresholds 132, 133, and 138 are switched to separate the two energy ranges separated by the energy threshold, and the group B X-ray detection element It can be realized by a method such as separating two energy ranges separated by one energy threshold.

<Fifth embodiment>
This embodiment includes, as a detected energy spectrum distribution changing unit, an irradiation spectrum changing unit adopted in the first embodiment, and an energy threshold changing unit adopted in the third embodiment, and a function of a control unit. It is characterized by including a timing control unit that controls the irradiation spectrum changing unit and the energy threshold changing unit.

FIG. 23 shows the overall configuration of the X-ray CT apparatus of the present embodiment, and FIG. 24 shows a functional block diagram of the control unit 107. 23, the same elements as those in FIGS. 1 and 14 are denoted by the same reference numerals, and description thereof will be omitted. As shown in FIG. 23, the X-ray CT apparatus includes a spectrum changing unit 111 that changes a tube voltage, a thickness and a type of a filter, and the like, and changes an energy spectrum of irradiated X-rays, and an X-ray detector 104 . An energy threshold changing unit 112 for changing an energy threshold at the time of reading a line photon is provided.

  As shown in FIG. 24, the control unit 107 controls the operation of the scanner control unit 1071, which controls the scanner equipped with the X-ray source 100 and the X-ray detector 104, the readout circuit of the X-ray detector 104, and the operation of the signal collection unit 108. An X-ray detector control unit 1072 for controlling, a display control unit 1073 for controlling display by the display unit 106, and a timing control unit 1074 for controlling timing of changing the detected energy spectrum are provided. The scanner control unit 1071, the X-ray detector control unit 1072, and the display control unit 1073 are common to the first to third embodiments. The X-ray CT apparatus according to the second embodiment may have a function (an imaging condition control unit 1075) for controlling an imaging condition corresponding to the imaging condition determination unit 1060. Here, the operation of the timing control unit 1074, which is a feature of the present embodiment, will be mainly described.

  The timing control unit 1074 performs switching of the irradiation energy spectrum and switching of the energy threshold at different periods, and acquires a plurality of data having different combinations of the irradiation energy spectrum and the energy threshold. Here, different periods mean, for example, a case where the period width is the same and the phase is different, or a case where the period width is different. The type and energy threshold of the irradiation energy spectrum used for the timing control (value and its type), and the switching cycle may be programmed in advance, or the operator may operate the input unit (UI) 110. The setting may be arbitrarily set via the Internet.

  FIG. 25 shows an example of switching between the X-ray energy spectrum and the energy threshold controlled by the timing control unit 1074. FIG. 25 shows a case where two irradiation energy spectra and two energy thresholds (for example, thresholds 132 and 133 in FIG. 15) are switched at different periods.

  In FIG. 25, the horizontal axis 185 is the elapsed time, for example, the time 176 is the start of shooting and the start time of the first view, and the times 177 to 181 are 2 views, 3 views, 4 views, 5 views and 6 Start time of the view. The vertical axis in FIG. 25 indicates the content of the selection of the X-ray spectrum and the energy threshold. In the range in which the timing curve 170 of the X-ray spectrum has the value 172, the spectrum changing means 111 selects the first spectrum and irradiates it with X-rays. Selectively emit X-rays. In the range where the energy threshold timing curve 171 has the value 175, the energy threshold changing unit 112 selects the threshold 132 (hereinafter, referred to as the low energy threshold 132) shown in FIG. Then, the threshold 133 shown in FIG. 15 (hereinafter referred to as the high energy threshold 133) is selected.

  That is, in this timing control, the irradiation X-ray of the first spectrum is used for the first view, and data is acquired with the energy threshold of the X-ray detector as the low energy threshold 132. Next, at time 177, the energy threshold changing unit 112 switches the energy threshold to the high energy threshold 133, and thereby, in the second view, the irradiation X-ray is the first spectrum, and the energy threshold is the high energy threshold 133. Data is obtained.

  Next, at time 178, the energy threshold is switched to the low energy threshold 132, and the spectrum changing unit 111 switches the irradiation X-ray spectrum to the low energy threshold 132. Thus, in the third view, the irradiated X-ray is the second spectrum, and the data is acquired with the energy threshold value of the low energy threshold value 132.

  Next, at time 179, the energy threshold changing unit 112 switches the energy threshold to the high energy threshold 133. Thus, in the fourth view, the X-ray is the second spectrum, and the data is acquired with the energy threshold value of the high energy threshold value 133. Next, at time 181, the energy threshold is switched to the low energy threshold 132, and the spectrum changing unit 111 switches the irradiation X-ray spectrum to the high energy threshold 133.

  Thus, the X-ray is the second spectrum in the fifth view and the data is acquired with the low energy threshold 132 as the energy threshold, as in the first view. Thereafter, by sequentially switching in the same order, the sixth view is similar to the second view, the seventh view is similar to the third view, and the eighth view is similar to the fourth view in the X-ray spectrum and energy threshold. Data can be obtained in combination. Furthermore, in the subsequent views, the i-th view (i is an integer of 1 to 4) and the (4n + i) -th view (n is an integer of 1 or more) can acquire data with the same combination of the X-ray spectrum and the energy threshold. FIG. 26 shows data acquired by such imaging.

When acquiring the data of all the views with each combination of the X-ray spectrum and the energy threshold, shift the time 176 of FIG. 25 by one view, perform the same imaging as the start time of the second view, and then Shooting may be repeated such that similar shooting is performed using 176 as the start time of the third view, and similar shooting is performed using time 176 as the start time of the fourth view.

  As a result, data can be obtained for all combinations of X-ray spectra and energy thresholds. By performing multi-energy arithmetic processing on these, a multi-energy image of, for example, [the number of combinations × 2] can be obtained, and a plurality of narrow energy range data can be obtained by applying the third embodiment. The multi-energy calculation processing is the same as the processing described in the first to fourth embodiments, and a duplicate description will be omitted.

  In order to obtain a large amount of data, one of the two detection spectrum changing means (the spectrum changing unit 111 and the energy threshold changing unit 112) may be performed at a slower cycle than the other, but the change of the X-ray spectrum may be performed. However, there is a problem that the load on the X-ray tube is large and it is technically difficult to change the X-ray tube at a higher speed. Therefore, it is desirable to switch the X-ray spectrum at a period slower than the energy threshold as shown in FIG. However, this change is an example, and the present embodiment includes a case where the energy threshold is switched at a cycle that is slower than the irradiation X-ray spectrum or a case that the energy threshold is switched at the same cycle.

  FIG. 25 illustrates a case where two types of X-ray spectra and two types of energy thresholds are switched, but this is an example and does not limit the present invention. In some cases, there are cases in which m (m is an integer of 2 or more) types of X-ray spectra and n (n is an integer of 2 or more) types of energy thresholds are switched. At this time, after switching n kinds of energy thresholds for one kind of X-ray spectrum, there is a method of changing the X-ray spectrum and similarly switching n kinds of energy thresholds. In another example, switching of n types of energy thresholds may be repeatedly performed for one type of X-ray spectrum. This means that, for example, when the number of repetitions is c (c is an integer of 2 or more), the switching of the energy threshold is repeated (n × c) times for one type of X-ray spectrum.

  FIG. 25 illustrates a case where at least one of the X-ray spectrum and the energy threshold is switched for each view, but this is an example and does not limit the present invention. For example, there may be various cases such as every plural views, every half circle, every one circle, every plural circles, and the like. That is, in FIG. 25, the time 176 to 177, the time 177 to 178, the time 178 to 179, the time 179 to 180, and the time 180 to 181 are not limited to the data acquisition interval of one view. , One time or a plurality of times.

  According to the present embodiment, the change in the detected X-ray energy distribution is performed by combining the change in the energy threshold that divides the energy range of the irradiated X-rays and the change in the spectrum of the irradiated X-rays, thereby reducing the number of comparators. Many multi-energy images can be obtained without increasing. In addition, by performing calculations between multi-energy projection data or multi-energy images, it is possible to improve the accuracy of the multi-energy image or obtain a plurality of narrow energy range data.

  In the present embodiment, the case where a combination of the X-ray spectrum and the energy threshold is obtained in one view is described, but this is an example and does not limit the present invention. In one view, only one or a part of the data of the combination of the X-ray spectrum and the energy threshold may be obtained, and all data may be obtained in a plurality of adjacent views. For example, in the photographing of FIG. 25, photographing of all combinations is performed in four views, but these may be used as one set.

  In other words, the four views may be regarded as one view, and the photographing data obtained between the four views may be treated as being simultaneously acquired. Furthermore, the deviation of those views may be corrected. As a method thereof, for example, data for every four views may be used to estimate and create data for the view between them, and data for all combinations may be created for all views. As another method, for example, when a reconstructed image is created from projection data obtained in different views, a reconstructed image at the same position may be created by correcting a difference in the acquisition angle.

<Sixth embodiment>
The X-ray CT apparatus of the present embodiment is characterized in that it has a function of simulating conditions set by the X-ray CT apparatuses of the above-described first to fifth embodiments and images to be acquired. That is, the X-ray CT apparatus of the present embodiment includes a simulation unit (1064) that creates a multi-energy image of a simulated subject by simulation, and an input unit (110) that inputs information including subject information used in the simulation unit. And a storage unit (109) for storing a plurality of subject models.

  The simulation unit determines a simulated object to be simulated from the object model stored in the storage unit based on the object information input via the input unit, and determines a multi-energy image of the determined simulated object. create. The energy range used for the simulation may be a value input from the input unit, or may be determined based on the determined simulated subject. For example, the simulation unit calculates an optimal energy range by calculating and comparing physical quantities from the created multi-energy image.

  FIG. 27 shows a configuration example of the calculation unit 105 for implementing the present embodiment. The correction processing unit 1051, the multi-energy calculation unit 1052, and the image reconstruction unit 1053 have the same functions as those of the calculation unit 105 illustrated in FIG. 5 and FIG. In the present embodiment, in addition to these, a simulated subject selecting unit 1063 and a simulation unit 1064 are provided. Further, the storage unit 109 stores a plurality of object models so that a simulated object can be selected according to the input object information. The simulation result formed by the simulation unit 1064 is passed to the imaging condition control unit 1075.

  Hereinafter, the X-ray CT apparatus according to the present embodiment has a function realized by changing the irradiation X-ray spectrum described in the first embodiment and a K-edge realized by switching the energy threshold described in the third embodiment. The operation of the X-ray CT apparatus according to the present embodiment will be described with reference to a display example of a graphical user interface (hereinafter, referred to as a GUI) shown in FIG. 28 and a flow of FIG. 29 in a case where the apparatus has both functions of creating an image. I will explain while.

  The GUI shown in FIG. 28 also serves as the input unit 110 displayed on the display unit 106. FIG. 28 shows only a part of the entire GUI for the sake of simplicity, and similarly shows only a part of the entire shooting method, shooting parameters, and functions in the software. . Similarly, the shooting and input procedures described below also show a part of the whole.

  First, before starting photographing, photographing conditions are input. The shooting conditions are set, for example, using the setting area 230.

  As an imaging method, first, ON / OFF of dual imaging and ON / OFF of K-edge imaging are selected (S721). For the ON / OFF selection, for example, a checkbox 200 shown in FIG. 28 is checked to select dual imaging, and a checkbox 201 is checked to perform K-edge imaging. When the dual shooting is turned on, the input fields for the first and second shooting parameters can be input, and when the dual shooting is turned off, only the input field for the first shooting parameter can be input. FIG. 28 shows a case where both shootings are ON. However, instead of selecting both K edge imaging and dual imaging, either one may be used.

  For example, the first shooting is performed on the first lap and the second shooting is performed on the second lap. However, the present invention is not limited to this. For example, after the first imaging is performed for several or half of the circumference, the second imaging for the same circumference may be performed. Further, the first imaging and the second imaging may be repeatedly performed for every other view or every plural views. Further, the view amount and the number of revolutions of the first and second photographing may be different.

  Next, a photographing parameter is input or selected. In the present embodiment, the imaging parameters in the first imaging and the second imaging are the tube voltages 204 and 206 and the tube currents 205 and 207. The tube voltage input fields 204 and 206 can be input, for example, from 60 kV to 140 kV every 1 kV, and the tube current input fields 205 and 207 can be input every 10 mA from 10 mA to 1200 mA, for example.

  However, these input ranges and steps are merely examples. Next, the part 202 to be imaged and the body type 203 of the subject are input. The imaging region selection field 202 can be selected from, for example, head and neck, chest, heart, abdomen, and pelvis, and the body type selection field 203 can be selected from, for example, lean, normal, obese, and children. Note that the subject information to be input is not limited to the imaging site and body type, but may include various items of subject information such as gender and weight. Further, it is also possible to obtain object information obtained by imaging means other than pre-scanning and X-ray CT, for example, information such as the size and transmission amount of the object, and let the simulation unit 1064 take in the object information as object information. good.

  A simulated subject is selected and determined from a plurality of subject models stored in the storage unit 109 according to the input subject information (S722). In the example shown in FIG. 28, the heart is selected as the imaging site 202, and a simulated subject 274 as shown in FIG. 30 is selected as the simulated attenuator at that time. The simulated subject 274 simulates a chest and a heart, and a heart 270, a lung 271 and a bone 272 are arranged in a fat 273 simulating a human body. Further, an artery (not shown) is simulated in the heart 270, and blood and iodine contrast agent are mixed in the artery.

When the input of the imaging parameters is completed as described above and the structure of the attenuated body to be simulated is determined, the simulation unit 1064 sets the X-ray spectrum distribution (irradiated X-ray spectrum distribution) 210 in the first and second imagings as the target. The X-ray spectrum distribution after passing through the sample (irradiation X-ray spectrum distribution) 211 is calculated and displayed on the spectrum display units 208 and 209 (S723). These X-ray spectrum distributions 210 and 211 are represented, for example, with the horizontal axis 217 as energy and the vertical axis 216 as the number of X-ray photons of the energy.

  The irradiation X-ray spectrum distribution 210 is obtained by calculation using the input tube voltage and the value of the tube current, and is calculated using the tube voltage value 204 and the tube voltage value 205 in the first imaging. In the second photographing, it is calculated using the tube voltage value 206 and the tube voltage value 207. X-ray spectrum distribution 211 after transmission, using the attenuation coefficient and density determined in advance, determine the attenuation in the structure of the simulated attenuated body determined in S722, the irradiation X-ray spectrum distribution 210 previously determined to It is calculated by multiplication.

Instead of displaying both the irradiation X-ray spectrum distribution 210 and the spectrum of transmitted X-rays after passing through the subject, only one of them may be displayed. Although the X-ray spectrum varies depending on the position and view of the X-ray detection element, here, a representative position determined in advance, and the X-ray spectrum in the imaging direction determined in advance is displayed.

  Next, an energy threshold value for determining an energy range to be used in K-edge imaging and dual imaging, and a K-edge energy are determined (S724). This is set, for example, by the operator using the spectrum display unit 208 as shown in FIG. In FIG. 28, energy thresholds 213 and 215 limit the energy range used in K-edge imaging. The energy threshold 212 is for dividing the energy range for purposes other than K-edge imaging. The setting energy 214 is for setting the K edge energy.

  When the threshold value is input, the simulation unit 1064 simulates what kind of image can be actually obtained, creates and displays the image (S725). However, the image at this time is a reconstructed image of the simulated subject. Further settings required for the simulation are performed using, for example, the setting area 231 in FIG. First, a display image 220 and an image type 221 are selected.

  The display image 220 can be selected from, for example, a K-edge image, a reconstructed image photographed with a predetermined tube voltage, and a multi-energy image such as a monochromatic X-ray equivalent image and a density image. At this time, in the reconstructed image taken at a predetermined tube voltage, a column for inputting the tube voltage of the image to be created is displayed, and in the monochromatic X-ray equivalent image, an input column for inputting the energy of the image to be created is displayed. In the image, columns for selecting a reference substance are displayed, and the user performs input and selection in these columns. The image type 221 can be selected from volume rendering, MIP, MPR, and the like. At this time, if necessary, input fields for the cutout position and direction are displayed, and the user performs an input thereto. In the example of FIG. 28, first, a K-edge image is selected.

  When the execute button 222 is pressed after the above input, a simulation image 223 is displayed on the display unit 106. In FIG. 28, a reconstructed image of the heart 270 and the coronary artery 275 is displayed. Here, an image in which the coronary artery 275 is extracted for K-edge imaging is obtained. Varies by. The simulation unit changes these setting values as appropriate (S726), and confirms the image quality such as the contrast, noise, and CNR of the image to find the optimum value.

Next, the display image 220 is changed to, for example, a monochromatic X-ray equivalent image. In this case, by checking the images while changing the energy range, it is possible to find the optimal setting. These processes are not shown in FIG. 29, but are similar to processes S725 and S726. Although the case where a monochrome X-ray equivalent image is displayed after displaying a K-edge image as a simulation image has been described, the display order is arbitrary. Also, one or more various multi-energy images may be displayed. Furthermore, not only the reconstructed image but also various multi-energy projection data may be displayed.

  Then, imaging is performed using the optimized energy threshold value and the set energy value (S727, S728). In the actual imaging, the energy range is switched to the energy thresholds 212, 213, 214, and 215 to perform the imaging, and as shown in the third embodiment and the fourth embodiment, from the projection data acquired respectively, The number of X-ray photons between energy ranges can be calculated, and a multi-energy image such as a K-edge image can be obtained.

  As described above, by determining the optimum energy range and the set energy by simulation in advance of photographing, it is possible to acquire an image with good image quality such as contrast, noise, and CNR in actual photographing.

  Although the present embodiment has been described with reference to FIGS. 28 and 29, the present embodiment is not limited to the illustrated embodiment and can be variously modified.

  For example, the GUI shown in FIG. 28 is merely an example, and it goes without saying that there may be a case where there are no items such as input and display in the GUI. Furthermore, it goes without saying that a function for performing another input, selection, or the like, or another display may be added. For example, it is also possible to input a cycle for switching between the first imaging and the second imaging with the GUI.

  The selection items and input rules in the GUI shown in FIG. 28 are also examples, and do not limit the present invention. For example, in the X-ray CT apparatus of the fourth embodiment, the projection data of the K-edge imaging with the X-ray detection elements of one group (Group A) of the X-ray detectors, the X-ray of another group (Group B) When imaging is performed to acquire projection data other than K-edge imaging with the detection element, there may be various modes such as designating the former with the thresholds 213 to 215 of the spectrum display unit 208 and the latter with the threshold 212.

  Further, the imaging procedure shown in FIG. 29 is also an example, and it goes without saying that there may be cases where some procedures are not present, the order is different, and other procedures may be added. For example, in the flow illustrated in FIG. 29, the energy range is manually determined.However, a case in which a value determined in advance is used, or a case in which the energy range is automatically determined using the transmitted X-ray spectrum 211 or the like may be used. It is possible.

  Further, one or a part of the energy thresholds of the K edge imaging and the dual imaging may be determined manually or automatically. In this case, as shown in FIG. 31, for example, in the same manner as in step S722 of FIG. 29, after selecting a simulated subject from the subject model based on the input subject information, the multi-energy image is changed by changing the energy range. Are created (S731).

  Next, the physical quantities at predetermined positions of these reconstructed images are calculated to determine the best energy range (S732, S733). As the physical quantities used here, contrast, noise, CNR, etc. can be adopted, and the energy range when a multi-energy image with the best physical quantities is obtained is determined as the optimal energy range. The position on the reconstructed image at the time of calculating the physical book may be specified in advance, or may be specified by the user.

  Similarly, instead of manually defining the energy of the K edge, the energy of the K edge may be automatically specified by specifying the metal to be used in the contrast inspection, or generally used in the contrast inspection The K-edge energy of a heavy metal, for example, iodine may be set as a default. In this case, it is needless to say that the default K-edge energy may be changed by specifying the energy of the K-edge or the metal.

  Further, in the present embodiment, the case where ON / OFF of the dual shooting is performed by inputting from the input field is described.For example, when the number of reference materials is large, when the dual shooting needs to be performed, the automatic shooting is performed. It may have a function to be turned on in a specific way. Parameters such as the type and number of reference substances may be automatically determined and used, or may be input by the operator via the input unit 110. In particular, since the type and the number of the reference substances are parameters that affect the image quality, it is preferable that a setting field is provided in the setting area 230 or the setting area 231.

  Further, in the present embodiment, the spectrum calculation and the image simulation are performed using the object model. However, for example, instead of the object model, the calculation may be performed based on an image of a real human body taken in advance.

  In the present embodiment, the case where the X-ray spectrum after transmission in the imaging direction determined in advance is displayed at a representative position determined in advance is described, but this is an example, and the present invention is described. It is not limited. For example, an X-ray spectrum at a position where the number of transmitted X-ray photons is smallest and an imaging direction may be displayed. Further, an input unit that allows a user to set an arbitrary imaging direction may be provided in a GUI or the like, and an X-ray spectrum in the imaging direction set by the set X-ray detection element may be displayed.

  According to the present embodiment, a GUI that allows an operator to easily input imaging conditions and information necessary for imaging when performing imaging with the X-ray CT apparatus of the present invention is provided. In particular, by presenting a spectrum or a simulation image created by the simulation unit before shooting, the operator can set shooting conditions more appropriately.

<Application example>
In each of the above embodiments and the modifications thereof, the case where a photon-counting type detector is used as the energy sorting X-ray detector has been described as an example, but the present invention is not limited to this. For example, the X-ray detector 104 may be a detector that performs energy classification, and then integrates and outputs the signal amounts.

  In each of the above-described embodiments and the modifications thereof, as a detected X-ray energy distribution changing unit, a spectrum changing unit that changes the spectrum of the irradiated X-ray and an energy threshold when the X-ray detector performs energy classification are changed. Although the energy threshold changing unit has been exemplified, the detected X-ray energy distribution changing means is not limited to these, and various other means for changing the energy distribution of X-rays detected by the X-ray detector in each energy range may be used. I can take it.

  In each of the above embodiments and the modifications thereof, the medical X-ray CT apparatus has been described as an example.However, the present invention is not limited to this, and the radiation incident on the detection element is divided for each energy range. Needless to say, the present invention can be applied to any CT apparatus equipped with a photon counting type radiation detector for counting the number of photons. As an example, there may be an X-ray CT apparatus for nondestructive inspection, an X-ray cone beam CT apparatus, and the like.

  Further, the present invention is not limited to the above-described embodiment, and can be variously modified and implemented in an implementation stage without departing from the gist of the invention. Furthermore, the above embodiment includes various stages, and various inventions can be extracted by appropriately combining a plurality of disclosed components. For example, some components may be deleted from all the components shown in the embodiment.

  As an example, there may be various radiation imaging apparatuses that create and display the multi-energy projection data 144 and images calculated based on the multi-energy projection data 144 without performing the image reconstruction processing without having the reconstruction processing unit 1053. Examples include X-ray diagnostic imaging devices, X-ray imaging devices, X-ray fluoroscopy devices, mammography, digital subtraction devices, nuclear medicine screening devices, radiation therapy devices, and the like.

  ADVANTAGE OF THE INVENTION According to this invention, the imaging device provided with the photon-counting type detector and realizing high-precision multi-energy image, material identification and classification, and reduction of the number of circuits is provided.

100 X-ray source, 101 Gantry rotation unit, 103 bed top, 104 X-ray detector, 105 calculation unit, 106 display unit, 107 control unit, 108 signal collection unit, 109 storage unit, 110 input unit, 111 spectrum change unit , 112 Energy threshold changing unit, 123 to 125 sampling time, 130, 135 Low energy range, 131, 136 High energy range, 132, 133, 138 Energy threshold, 134, 137 Narrow energy range, 400 X-ray detector, 401 detection Layers, 402 , 403 electrodes, 405 readout circuit, 1050 main control unit, 1051 correction processing unit, 1052 multi-energy calculation unit, 1053 image reconstruction unit, 1054 defective element correction unit, 1055 air correction unit, 1056 density image calculation unit, 1057 Multi-energy projection data calculation unit, 1058 Narrow energy range data calculation unit, 1059 K edge image calculation unit, 1060 imaging condition determination unit, 1061 X-ray photon number ratio calculation unit, 1062 spectrum condition determination unit, 1063 simulated subject selection unit , 1064 Simulation unit, 1071 a scanner control unit, 1072 X-ray detector controller, 1073 a display controller, 1074 a timing control unit, 1075 photographing condition control unit

Claims (1)

An X-ray source for irradiating X-rays,
An X-ray detector that separates and detects incident X-rays into a plurality of energy ranges,
In each of the plurality of energy ranges, a detected energy distribution changing unit that changes the detected X-ray energy distribution detected by the X-ray detector,
A plurality of detected X-ray energy distributions, a signal collection unit that processes the output signals of the X-ray detectors in the plurality of energy ranges to obtain projection data,
A reconstruction unit that reconstructs the projection data to create a reconstructed image,
Create multi-energy projection data using the projection data obtained with the plurality of detected X-ray energy distributions, and / or a multi-energy image using the reconstructed image obtained with the plurality of detected X-ray energy distributions. A multi-energy operation unit for creating
With
The detected energy distribution changing unit includes an energy threshold changing unit that changes an energy threshold that determines the energy range of the X-ray detector.The detected X-ray energy is changed by changing the energy threshold by the energy threshold changing unit. Change the distribution,
The multi-energy operation unit, using a plurality of the projection data obtained when the energy threshold is different, to create the multi-energy projection data,
The X-ray detector separates the energy threshold and performs counting separately in a first energy range on the low energy side and a second energy range on the high energy side,
The energy threshold change unit, the energy threshold, at least, a first energy threshold, a second energy threshold higher than the first energy threshold, to change to take,
The signal collection unit, when the energy threshold is the first energy threshold, for each of the case of the second energy threshold, to create the projection data,
The multi-energy calculation unit is obtained in at least one energy range of the first energy range and the second energy range, the projection data in the case of the first energy threshold and the second energy threshold Using the projection data in the case, comprising a narrow energy range data calculation unit that calculates narrow energy range projection data is projection data of the energy range between the first energy threshold and the second energy threshold,
The narrow energy range data calculation unit,
The number of X-ray photons in the projection data of the first energy range in the case of the first energy threshold, and the number of X-ray photons in the projection data of the second energy range in the case of the second energy threshold. Is used to estimate the noise of each projection data,
X is characterized by calculating the narrow energy range projection data by subtracting the projection data in the case of the first energy threshold and the projection data in the case of the second energy threshold in the energy range of the smaller noise. X-ray CT device.

Images