Radiation imaging apparatus, information processing apparatus, information processing method, and program

Abstract

To acquire an energy spectrum of irradiated radiation.SOLUTION: A radiation imaging apparatus includes a processing unit that executes processing of aggregating energy information obtained by dividing radiation photon energy acquired with time in a time direction, thereby acquiring an energy spectrum.SELECTED DRAWING: Figure 8

Description

Technical Field

[0001] The disclosed technology relates to a radiation imaging device, an information processing device, an information processing method, and a program. More specifically, it relates to a radiation imaging device, an information processing device, an information processing method, and a program used for still image shooting such as general shooting in medical diagnosis and moving image shooting such as fluoroscopy shooting.

Background Art

[0002] Currently, as an imaging device used for medical image diagnosis by X-rays and non-destructive inspection, a radiation imaging device using a flat panel detector (Flat Panel Detector, hereinafter abbreviated as FPD) formed of a semiconductor material has become widespread.

[0003]

[0004] In the energy subtraction process, as a shooting method using an FPD, by processing a plurality of images with different energies obtained by irradiating X-rays with different tube voltages, a substance separation image with reduced contrast, for example, a bone image or a soft tissue image can be obtained (Patent Document 1). However, since the time difference between a plurality of images is determined by the shooting time of the FPD, if the subject moves during this time, for example, artifacts due to movement may occur in the images obtained by the energy subtraction process.

[0005] Patent Document 2 describes a dual energy imaging system that generates X-ray pulses with different kV values on a millisecond time scale, sample-holds the signal integration for the first sub-image at the first kV value, and performs signal integration corresponding to the second sub-image at the second kV value in parallel with the reading out of the first sub-image.

Prior Art Documents

Patent Documents

[0006]

Patent Document 1

Patent Document 2

Summary of the Invention

Problems to be Solved by the Invention

[0007] However, in energy subtraction processing, since information on a plurality of X-ray spectra with different X-ray energies is required, in an imaging method (hereinafter also referred to as time-division imaging) for obtaining a plurality of sub-images by sampling with a very short time difference within one pulse of X-ray irradiation period, it is difficult to actually measure the X-ray spectrum of each image, and processing using an X-ray spectrum estimated based on pixel values of sub-images and the like has been performed, which has been a factor reducing the accuracy of subtraction processing.

[0008] The disclosed technology provides a technology capable of obtaining the energy spectrum of irradiated radiation.

Means for Solving the Problems

[0009] A radiation imaging apparatus according to an aspect of the disclosed technology includes processing means for performing a process of obtaining an energy spectrum by aggregating energy information obtained by dividing radiation photon energy acquired in time series in the time direction.

Effects of the Invention

[0010] According to the disclosed technology, the energy spectrum of irradiated radiation can be obtained. By using the obtained energy spectrum for energy subtraction processing, it becomes possible to improve the accuracy of the processing.

Brief Description of the Drawings

[0011] [Figure 1] A diagram showing a configuration example of an X-ray imaging system according to the first embodiment. [Figure 2] A pixel equivalent circuit diagram of an X-ray imaging apparatus according to the first embodiment. [Figure 3]Timing chart of an X-ray imaging device according to the first embodiment. [Figure 4] Timing chart of an X-ray imaging device according to the first embodiment. [Figure 5] A diagram illustrating the correction process according to the first embodiment. [Figure 6] Block diagram of signal processing according to the first embodiment. [Figure 7] Block diagram of image processing according to the first embodiment. [Figure 8] A diagram illustrating the principle of time-resolved X-ray spectroscopy according to the first embodiment. [Figure 9] A diagram showing the relationship between the acquisition of the X-ray spectrum and the driving timing of the X-ray imaging system according to the first embodiment. [Figure 10] A diagram showing the processing flow of the X-ray imaging system according to the first embodiment. [Figure 11] A diagram illustrating the waveform of an X-ray. [Figure 12] A diagram illustrating the principle of image quality simulation. [Figure 13] This figure exemplifies the correlation between the timing of sample hold and the image quality obtained by energy subtraction processing. [Figure 14] A diagram showing the processing flow of the X-ray imaging system according to the second embodiment. [Figure 15] A diagram showing the processing flow of the X-ray imaging system according to the third embodiment. [Modes for carrying out the invention]

[0012] The embodiments will be described in detail below with reference to the attached drawings. Note that the following embodiments do not limit the invention to the claims. While the embodiments describe multiple features, not all of these features are essential to the invention, and the features may be combined in any way. Furthermore, in the attached drawings, the same or similar configurations are given the same reference numerals, and redundant descriptions are omitted.

[0013] Furthermore, the radiation in the disclosed technology includes not only alpha rays, beta rays, and gamma rays, which are beams created by particles (including photons) emitted by radioactive decay, but also beams with energy of equal or greater magnitude, such as X-rays, particle beams, and cosmic rays. In the following embodiments, an apparatus using X-rays will be described as an example of radiation. Therefore, in the following, radiation imaging apparatus and radiation imaging system will be described as X-ray imaging apparatus and X-ray imaging system, respectively.

[0014] [First Embodiment] Figure 1 is a block diagram showing an example configuration of an X-ray imaging system as an example of a radiation imaging system according to the first embodiment. The X-ray imaging system of the first embodiment comprises an X-ray generator 101, an X-ray control device 102, an imaging control device 103, and an X-ray imaging device 104.

[0015] The X-ray generator 101 generates X-rays and irradiates the subject with them. The X-ray control device 102 controls the generation of X-rays in the X-ray generator 101. The imaging control device 103 has, for example, one or more processors (CPUs) and memory, and the processor executes a program stored in memory to acquire X-ray images and perform image processing. Each process, including image processing by the imaging control device 103, may be implemented by dedicated hardware or by the cooperation of hardware and software. The X-ray imaging device 104 has a phosphor 105 that converts X-rays into visible light and a two-dimensional detector 106 that detects visible light. The two-dimensional detector is a sensor in which pixels 20 that detect X-ray quanta are arranged in an X column × Y row array and outputs image information.

[0016] The imaging control device 103 functions as an information processing device that processes radiation images using the processor described above. The acquisition unit 131, correction unit 132, signal processing unit 133, and image processing unit 134 show an example of the functional configuration as an information processing device.

[0017] The acquisition unit 131 acquires multiple radiation images with different energies obtained by irradiating the subject with radiation and taking a photograph. The acquisition unit 131 acquires radiation images as multiple radiation images obtained by performing multiple sample-and-hold operations during the irradiation of one shot of radiation.

[0018] The correction unit 132 corrects multiple radiation images acquired by the acquisition unit 131 to generate multiple images used in energy subtraction processing.

[0019] The signal processing unit 133 aggregates energy information obtained by dividing the X-ray photon energy acquired in time series in the time direction and performs a process to acquire an energy spectrum. The signal processing unit 133 acquires the energy spectrum by counting the number of photons in the time series in the energy information. Furthermore, the signal processing unit 133 uses the energy spectrum obtained by aggregating the energy information obtained by dividing the radiation photon energy acquired in time series in the time direction to irradiate the subject with radiation and take an image, thereby acquiring multiple images corresponding to multiple different radiation energies, and performs energy subtraction processing using the multiple images. When applying the acquisition of the energy spectrum to energy subtraction processing, the signal processing unit 133 performs energy subtraction processing using multiple images corresponding to multiple different X-ray energies obtained by irradiating the subject with X-rays and taking an image, and the energy spectrum acquired based on the X-ray photon energy acquired in time series. Here, the signal processing unit 133 aggregates the energy information obtained by dividing the X-ray photon energy at the timing when the X-ray energy signal is sampled and held, and acquires the energy spectrum.

[0020] Furthermore, the signal processing unit 133 generates material property images using multiple images generated by the correction unit 132. Material property images are images acquired in energy subtraction processing, such as material separation images that separate and represent materials, such as bone and soft tissue, and material identification images that represent the effective atomic number and its surface density. Based on multiple radiation images taken at different radiation energies, the signal processing unit 133 generates, for example, a first material separation image showing the thickness of a first material and a second material separation image showing the thickness of a second material. The signal processing unit 133 also generates a thickness image that combines the thicknesses of the first and second materials. Here, the first material includes at least calcium, hydroxyapatite, or bone, and the second material includes at least water, or fat or soft material that does not contain calcium. Details of the signal processing unit 133 will be described later. The image processing unit 134 generates a display image using the material property images acquired by the signal processing of the signal processing unit 133.

[0021] Figure 2 is an equivalent circuit diagram of a pixel 20 according to the first embodiment. The pixel 20 includes a photoelectric conversion element 201 and an output circuit section 202. The photoelectric conversion element 201 can typically be a photodiode. The output circuit section 202 includes an amplification circuit section 204, a clamping circuit section 206, a sample-and-hold circuit 207, and a selection circuit section 208.

[0022] The photoelectric conversion element 201 includes a charge storage unit, which is connected to the gate of the MOS transistor 204a of the amplification circuit 204. The source of the MOS transistor 204a is connected to the current source 204c via the MOS transistor 204b. The MOS transistor 204a and the current source 204c constitute a source follower circuit. The MOS transistor 204b is an enable switch that turns on when the enable signal EN supplied to its gate reaches an active level, thereby putting the source follower circuit into operation.

[0023] In the example shown in Figure 2, the charge storage unit of the photoelectric conversion element 201 and the gate of the MOS transistor 204a form a common node, which functions as a charge-voltage conversion unit that converts the charge stored in the charge storage unit into a voltage. That is, a voltage V (=Q / C) appears in the charge-voltage conversion unit, which is determined by the charge Q stored in the charge storage unit and the capacitance value C of the charge-voltage conversion unit. The charge-voltage conversion unit is connected to a reset potential Vres via a reset switch 203. When the reset signal PRES reaches an active level, the reset switch 203 turns on, and the potential of the charge-voltage conversion unit is reset to the reset potential Vres.

[0024] The clamp circuit 206 clamps the noise output by the amplification circuit 204 in accordance with the potential of the reset charge-voltage conversion section using the clamp capacitor 206a. In other words, the clamp circuit 206 is a circuit for canceling this noise from the signal output from the source follower circuit in accordance with the charge generated by photoelectric conversion in the photoelectric conversion element 201. This noise includes the kTC noise at the time of reset. Clamping is performed by setting the clamp signal PCL to an active level to turn on the MOS transistor 206b, and then setting the clamp signal PCL to an inactive level to turn off the MOS transistor 206b. The output side of the clamp capacitor 206a is connected to the gate of the MOS transistor 206c. The source of the MOS transistor 206c is connected to the current source 206e via the MOS transistor 206d. The source follower circuit is composed of the MOS transistor 206c and the current source 206e. The MOS transistor 206d is an enable switch that turns on when the enable signal EN0 supplied to its gate reaches an active level, putting the source follower circuit into operation.

[0025] The signal output from the clamp circuit 206 in response to the charge generated by photoelectric conversion in the photoelectric conversion element 201 is written as an optical signal to the capacitor 207Sb via switch 207Sa when the optical signal sampling signal TS reaches an active level. The signal output from the clamp circuit 206 when the MOS transistor 206b is turned on immediately after resetting the potential of the charge-voltage conversion unit is the clamp voltage. The noise signal is written to the capacitor 207Nb via switch 207Na when the noise sampling signal TN reaches an active level. This noise signal includes the offset component of the clamp circuit 206. The signal sample-and-hold circuit 207S is formed by switch 207Sa and capacitor 207Sb, and the noise sample-and-hold circuit 207N is formed by switch 207Na and capacitor 207Nb. The sample-and-hold circuit 207 includes the signal sample-and-hold circuit 207S and the noise sample-and-hold circuit 207N.

[0026] When the drive circuit drives the row selection signal to an active level, the signal (optical signal) held in capacitor 207Sb is output to signal line 21S via MOS transistor 208Sa and row selection switch 208Sb. At the same time, the signal (noise) held in capacitor 207Nb is output to signal line 21N via MOS transistor 208Na and row selection switch 208Nb. MOS transistor 208Sa forms a source follower circuit with a constant current source (not shown) provided on signal line 21S. Similarly, MOS transistor 208Na forms a source follower circuit with a constant current source (not shown) provided on signal line 21N. The signal selection circuit section 208S is formed by MOS transistor 208Sa and row selection switch 208Sb, and the noise selection circuit section 208N is formed by MOS transistor 208Na and row selection switch 208Nb. The selection circuit section 208 includes a signal selection circuit section 208S and a noise selection circuit section 208N.

[0027] Pixel 20 may have an adder switch 209S for adding the optical signals of multiple adjacent pixels 20. In adder mode, the adder mode signal ADD becomes active, and the adder switch 209S turns ON. As a result, the capacitances 207Sb of adjacent pixels 20 are interconnected by the adder switch 209S, and the optical signals are averaged. Similarly, pixel 20 may have an adder switch 209N for adding the noise of multiple adjacent pixels 20. When the adder switch 209N turns ON, the capacitances 207Nb of adjacent pixels 20 are interconnected by the adder switch 209N, and the noise is averaged. The adder unit 209 includes an adder switch 209S and an adder switch 209N.

[0028] Furthermore, the pixel 20 may have a sensitivity changing unit 205 for changing the sensitivity. The pixel 20 may include, for example, a first sensitivity changing switch 205a and a second sensitivity changing switch 205'a, as well as circuit elements associated with them. When the first change signal WIDE reaches an active level, the first sensitivity changing switch 205a is turned on, and the capacitance value of the first additional capacitor 205b is added to the capacitance value of the charge-voltage conversion unit. This reduces the sensitivity of the pixel 20. When the second change signal WIDE2 reaches an active level, the second sensitivity changing switch 205'a is turned on, and the capacitance value of the second additional capacitor 205'b is added to the capacitance value of the charge-voltage conversion unit. This further reduces the sensitivity of the pixel 20. By adding a function to reduce the sensitivity of the pixel 20 in this way, it becomes possible to receive a larger amount of light, and the dynamic range can be widened. When the first change signal WIDE reaches an active level, the enable signal ENw may be set to an active level, and the MOS transistor 204'a may be used in source follower operation instead of the MOS transistor 204a.

[0029] The X-ray imaging device 104 reads the output of the pixel circuit described above from the two-dimensional detector 106, converts it to a digital value using an AD converter (not shown), and then transfers the image to the imaging control device 103.

[0030] Next, the operation of the X-ray imaging system of the first embodiment having the above-described configuration will be explained. Figure 3 shows the drive timing of the X-ray imaging device 104 when energy subtraction is performed in the X-ray imaging system according to the first embodiment. The waveforms in Figure 3, with time on the horizontal axis, show the timing of X-ray irradiation, synchronization signal, reset of the photoelectric conversion element 201, sample-and-hold circuit 207, and image readout from signal line 21.

[0031] X-rays are emitted after the photoelectric conversion element 201 is reset by a reset signal. Ideally, the X-ray tube voltage is a square wave, but the rise and fall times of the tube voltage are finite. In particular, when pulsed X-rays are emitted for a short irradiation time, the tube voltage can no longer be considered a square wave and takes on waveforms as shown in X-rays 301-303. The energy of the X-rays differs in the rising phase (X-ray 301), the stabilizing phase (X-ray 302), and the falling phase (X-ray 303). Therefore, by obtaining X-ray images corresponding to the radiation during periods separated by the sample-and-hold mechanism, multiple types of X-ray images with different energies can be obtained.

[0032] The X-ray imaging device 104 performs sampling with the noise sample-and-hold circuit 207N after irradiation with the rising phase X-ray 301, and then performs sampling with the signal sample-and-hold circuit 207S after irradiation with the stable phase X-ray 302. Subsequently, the X-ray imaging device 104 reads out the difference between signal line 21N and signal line 21S as an image. At this time, the noise sample-and-hold circuit 207N holds the signal (R1) of the rising phase X-ray 301, and the signal sample-and-hold circuit 207S holds the sum (R1+B) of the signal (B) of the rising phase X-ray 301 and the stable phase X-ray 302. Therefore, an image 304 corresponding to the signal of the stable phase X-ray 302 is read out.

[0033] Next, the X-ray imaging device 104 performs sampling again with the signal sample-and-hold circuit 207S after the irradiation of the falling-end X-ray 303 and the reading of the image 304 are completed. After that, the X-ray imaging device 104 resets the photoelectric conversion element 201 and performs sampling again with the noise sample-and-hold circuit 207N, reading out the difference between the signal line 21N and the signal line 21S as an image. At this time, the noise sample-and-hold circuit 207N holds the signal of the state when no X-rays are irradiated, and the signal sample-and-hold circuit 207S holds the signal of the rising-end X-ray 301, the stable-end X-ray 302, and the sum of the signals (R2) of the falling-end X-ray 303 (R1+B+R2). Therefore, the image 306 corresponding to the signal of the rising-end X-ray 301, the stable-end X-ray 302, and the falling-end X-ray 303 is read out. Subsequently, by calculating the difference between image 306 and image 304, image 305 is obtained, which corresponds to the sum of the rising phase X-ray 301 and the falling phase X-ray 303. This calculation may be performed by the X-ray imaging device 104 or by the imaging control device 103.

[0034] The timing for resetting the sample-and-hold circuit 207 and the photoelectric conversion element 201 is determined using a synchronization signal 307 that indicates the start of X-ray irradiation from the X-ray generator 101. As a method for detecting the start of X-ray irradiation, a configuration can be used in which the tube current of the X-ray generator 101 is measured and it is determined whether or not the current value exceeds a preset threshold, but this is not the only method. For example, after the reset of the photoelectric conversion element 201 is completed, a configuration can be used in which the start of X-ray irradiation is detected by repeatedly reading out the pixels 20 and determining whether or not the pixel value exceeds a preset threshold.

[0035] Alternatively, for example, the X-ray imaging device 104 may be equipped with an X-ray detector different from the two-dimensional detector 106, and the start of X-ray irradiation may be detected by determining whether the measured value exceeds a preset threshold. In either method, after a predetermined time has elapsed since the input of the synchronization signal 307 indicating the start of X-ray irradiation, sampling of the signal sample-and-hold circuit 207S, sampling of the noise sample-and-hold circuit 207N, and resetting of the photoelectric conversion element 201 are performed.

[0036] As described above, an image 304 corresponding to the stable phase of the pulsed X-ray and an image 305 corresponding to the sum of the rise and fall phases are obtained. Since the energies of the X-rays irradiated when forming these two X-ray images are different, energy subtraction processing can be performed by performing calculations between these X-ray images.

[0037] Figure 4 shows the drive timing of the X-ray imaging device 104 when energy subtraction is performed in the X-ray imaging system according to the first embodiment. The drive timing shown in Figure 4 differs from the drive timing in Figure 3 in that the tube voltage of the X-ray generator 101 is actively switched.

[0038] First, after the photoelectric conversion element 201 is reset, the X-ray generator 101 irradiates with low-energy X-rays 401. In this state, the X-ray imaging device 104 performs sampling using the noise sample-and-hold circuit 207N. Subsequently, the X-ray generator 101 switches the tube voltage to irradiate with high-energy X-rays 402. In this state, the X-ray imaging device 104 performs sampling using the signal sample-and-hold circuit 207S. Subsequently, the X-ray generator 101 switches the tube voltage to irradiate with low-energy X-rays 403. The X-ray imaging device 104 reads out the difference between signal line 21N and signal line 21S as an image. At this time, the noise sample-and-hold circuit 207N holds the signal of low-energy X-rays 401 (R1), and the signal sample-and-hold circuit 207S holds the sum of the signals of low-energy X-rays 401 and high-energy X-rays 402 (B) (R1+B). Therefore, image 404, which corresponds to the high-energy X-ray signal 402, is read out.

[0039] Next, the X-ray imaging device 104 performs sampling again with the signal sample-and-hold circuit 207S after the irradiation of low-energy X-rays 403 and the reading of image 404 are completed. After that, the X-ray imaging device 104 resets the photoelectric conversion element 201 and performs sampling again with the noise sample-and-hold circuit 207N, reading out the difference between signal line 21N and signal line 21S as an image. At this time, the noise sample-and-hold circuit 207N holds the signal of the state when X-rays are not irradiated, and the signal sample-and-hold circuit 207S holds the signal of low-energy X-rays 401 and the sum of the signals (R2) of high-energy X-rays 402 and low-energy X-rays 403 (R1+B+R2). Therefore, image 406 corresponding to the signals of low-energy X-rays 401, high-energy X-rays 402 and low-energy X-rays 403 is read out.

[0040] Subsequently, by calculating the difference between image 406 and image 404, image 405 corresponding to the sum of low-energy X-rays 401 and low-energy X-rays 403 is obtained. This calculation may be performed by the X-ray imaging device 104 or by the imaging control device 103. The synchronization signal 407 is the same as in Figure 3. In this way, by acquiring images while actively switching the tube voltage, the energy difference between the low-energy and high-energy radiation images can be made larger compared to the method in Figure 3.

[0041] Next, the energy subtraction processing by the imaging control device 103 will be described. In the first embodiment, the energy subtraction processing is divided into three stages: correction processing by the correction unit 132, signal processing by the signal processing unit 133, and image processing by the image processing unit 134. Each of these processes will be described below.

[0042] The correction process is a process that processes multiple radiation images acquired from the X-ray imaging device 104 to generate multiple images used in the signal processing described later in the energy subtraction process. Figure 5 shows a block diagram of the correction process for energy subtraction according to the first embodiment. First, the acquisition unit 131 causes the X-ray imaging device 104 to perform imaging without irradiating with X-rays, and acquires images using the drive shown in Figure 3 or Figure 4. Two images are read out by this drive. Hereinafter, the first image (image 304 or image 404) will be referred to as F_ODD, and the second image (image 306 or image 406) as F_EVEN. F_ODD and F_EVEN are images corresponding to the fixed pattern noise (FPN) of the X-ray imaging device 104.

[0043] Next, the acquisition unit 131 irradiates the X-ray imaging device 104 with X-rays when there is no subject to be photographed and acquires a gain correction image output from the X-ray imaging device 104 by the drive shown in Figure 3 or Figure 4. This drive reads out two images in the same way as above. Hereafter, the first gain correction image (image 304 or image 404) will be referred to as W_ODD, and the second gain correction image (image 306 or image 406) will be referred to as W_EVEN. W_ODD and W_EVEN are images corresponding to the sum of the FPN and the X-ray signal of the X-ray imaging device 104. The correction unit 132 subtracts F_ODD from W_ODD and F_EVEN from W_EVEN to obtain images WF_ODD and WF_EVEN from which the FPN of the X-ray imaging device 104 has been removed. This is called offset correction.

[0044] WF_ODD is the image corresponding to the stable phase of X-ray 302, and WF_EVEN is the image corresponding to the sum of the rising phase X-ray 301, the stable phase X-ray 302, and the falling phase X-ray 303. Therefore, the correction unit 132 subtracts WF_ODD from WF_EVEN to obtain the image corresponding to the sum of the rising phase X-ray 301 and the falling phase X-ray 303. This process of obtaining an image corresponding to a specific period of X-rays separated by sample-and-hold by subtracting multiple images is called color correction. The energies of the rising phase X-ray 301 and the falling phase X-ray 303 are lower than the energy of the stable phase X-ray 302. Therefore, by subtracting WF_ODD from WF_EVEN through color correction, a low-energy image W_Low is obtained when there is no subject. Also, a high-energy image W_High is obtained from WF_ODD when there is no subject.

[0045] Next, the acquisition unit 131 irradiates the X-ray imaging device 104 with X-rays while a subject is present to perform imaging, and acquires the image output from the X-ray imaging device 104 by the drive shown in Figure 3 or Figure 4. At this time, two images are read out. Hereafter, the first image (image 304 or image 404) will be referred to as X_ODD, and the second image (image 306 or image 406) as X_EVEN. The correction unit 132 performs offset correction and color correction in the same way as when there is no subject to obtain a low-energy image X_Low when a subject is present and a high-energy image X_High when a subject is present.

[0046] Here, if we let d be the thickness of the subject, μ be the linear attenuation coefficient of the subject, I0 be the output of pixel 20 when there is no subject, and I be the output of pixel 20 when there is a subject, then the following equation [Equation 1] holds.

[0047]

number

[0048] By rearranging equation [Equation 1], we obtain the following equation [Equation 2]. The right-hand side of equation [Equation 2] represents the attenuation rate of the subject. The attenuation rate of the subject is a real number between 0 and 1.

[0049]

number

[0050] Therefore, the correction unit 132 obtains an image L (hereinafter also referred to as "low-energy image L") of the attenuation rate at low energy by dividing the low-energy image X_Low (when a subject is present) by the low-energy image W_Low (when a subject is absent). Similarly, the correction unit 132 obtains an image H (hereinafter also referred to as "high-energy image H") of the attenuation rate at high energy by dividing the high-energy image X_High (when a subject is present) by the high-energy image W_High (when a subject is absent). This process of obtaining images (L,H) of the attenuation rate at low energy or high energy by dividing an image obtained based on a radiation image obtained in the state with a subject by an image obtained based on a radiation image obtained in the state without a subject is called gain correction.

[0051] Figure 6 shows a block diagram of the signal processing of the energy subtraction process according to the first embodiment. The signal processing unit 133 generates a material property image using multiple images obtained from the correction unit 132. Below, the process of generating a material separation image consisting of an image B of bone thickness (hereinafter also referred to as bone image B) and an image S of soft tissue thickness (hereinafter also referred to as soft tissue image S) will be explained. The signal processing unit 133 obtains the image B of bone thickness and the image S of soft tissue thickness from the image L of the attenuation rate at low energy and the image H of the attenuation rate at high energy obtained by the correction shown in Figure 5 through the following process.

[0052] First, let E be the energy of the X-ray photon, N(E) be the number of photons at energy E, B be the thickness in the bone thickness image, S be the thickness in the soft tissue thickness image, and μ be the linear attenuation coefficient of bone at energy E. B (E) The linear attenuation coefficient of soft tissue at energy E is μ S (E) If the attenuation rate is I / I0, then the following equation [Equation 3] holds.

[0053]

number

[0054] The number of photons N(E) in energy E is the spectrum of X-rays. The spectrum of X-rays can be obtained by simulation or actual measurement. Also, the linear attenuation coefficient μ B (E) of bone at energy E and the linear attenuation coefficient μ S (E) of soft tissue at energy E are obtained from databases such as NIST (National Institute of Standards and Technology), respectively. Therefore, according to Equation [3], it is possible to calculate the thickness B in an arbitrary bone thickness image, the thickness S in a soft tissue thickness image, and the attenuation rate I / I0 in the X-ray spectrum N(E).

[0055] Here, let the spectrum of X-rays with low energy be N L (E), and the spectrum of X-rays with high energy be N H (E). Then, for the attenuation rate in image L and the attenuation rate in image H, the following equations of Equation [4] hold. In the following description, the attenuation rate in image L shown in Equation [4] is also simply referred to as the low-energy attenuation rate L, and the attenuation rate in image H is also simply referred to as the high-energy attenuation rate H.

[0056]

Equation

[0057] By solving the non-linear simultaneous equations of Equation [4], the thickness B in the bone thickness image and the thickness S in the soft tissue thickness image can be obtained. As a typical method for solving non-linear simultaneous equations, the case of using the Newton-Raphson method is explained here. First, let the number of iterations of the Newton-Raphson method be m, the bone thickness after the m-th iteration be B m , and the soft tissue thickness after the m-th iteration be S m . When the high-energy attenuation rate H m and the low-energy attenuation rate L m after the m-th iteration are represented by the following Equation [5].[[]]END]]

[0058]

Equation

[0059] Furthermore, the rate of change in the attenuation rate when the thickness changes slightly is expressed by the following equation [Equation 6].

[0060]

number

[0061] At this time, the bone thickness B after the m+1th iteration. m+1 and the thickness S of the soft tissue m+1 This can be expressed by the following equation [Equation 7], using the high-energy attenuation rate H and the low-energy attenuation rate L.

[0062]

number

[0063] The inverse of a 2x2 matrix, where determinant is det, can be expressed by Cramer's rule as follows [Equation 8].

[0064]

number

[0065] Therefore, substituting equation [Equation 8] into equation [Equation 7] yields the following equation [Equation 9].

[0066]

number

[0067] By repeating the above calculations, the attenuation rate H of the high energy after the mth iteration can be calculated. m The difference in the measured high-energy attenuation rate H approaches zero. The same is true for the low-energy attenuation rate L. As a result, the bone thickness B after the mth repetition... m It converges to the bone thickness B, and to the soft tissue thickness S at the mth time. mThis converges to the thickness S of the soft tissue. In this way, the nonlinear system of equations shown in [Equation 4] can be solved. Therefore, by calculating [Equation 4] for all pixels, an image of bone thickness B and an image of soft tissue thickness S can be obtained from an image L of the attenuation rate at low energy and an image H of the attenuation rate at high energy.

[0068] In the first embodiment, an image B of bone thickness and an image S of soft tissue thickness were calculated, but the disclosed technology is not limited to this form. For example, the thickness W of water and the thickness I of contrast agent may be calculated. That is, the thickness may be decomposed into any two types of substances. Furthermore, an image of the effective atomic number Z and an image of the surface density D may be obtained from the image L of the attenuation rate at low energy and the image H of the attenuation rate at high energy obtained by the correction shown in Figure 5. The effective atomic number Z is the equivalent atomic number of the mixture, and the surface density D is the density of the subject [g / cm³]. 3 It is the product of [ ] and the thickness of the subject [cm].

[0069] Furthermore, in the first embodiment, the Newton-Raphson method was used to solve the nonlinear simultaneous equations. However, the disclosed techniques are not limited to this form. For example, iterative methods such as the least squares method or the bisection method may be used. Also, in the first embodiment, the nonlinear simultaneous equations were solved using an iterative method, but the disclosed techniques are not limited to this form. A configuration may be used in which bone thickness B and soft tissue thickness S for various combinations of high-energy attenuation rates H and low-energy attenuation rates L are determined in advance to generate a table, and bone thickness B and soft tissue thickness S are determined quickly by referring to this table.

[0070] Figure 7 shows a block diagram of the image processing of the energy subtraction process according to the first embodiment. The image processing unit 134 of the first embodiment performs image processing to obtain a virtual monochromatic X-ray image from the bone thickness image B and the soft tissue thickness image S obtained by the signal processing shown in Figure 6. A virtual monochromatic X-ray image is an image that is expected to be obtained when irradiated with X-rays of a single energy. For example, if the energy of the virtual monochromatic X-ray is E VIn this case, the virtual monochromatic X-ray image V can be obtained by the following [Equation 10]. [Number 10] V = exp{-μ} B (E V )B-μ S (E V )S} Virtual monochromatic X-ray images are used in Dual Energy CT, which combines energy subtraction and three-dimensional reconstruction. In this case, to improve the contrast-to-noise ratio (CNR) of the virtual monochromatic X-ray image, the energy of the virtual monochromatic X-ray is reduced. V Modify the linear attenuation coefficient μ of bone. For example, the linear attenuation coefficient μ of bone. B (E) is the linear attenuation coefficient μ of soft tissue. S It is larger than (E). However, the energy of the virtual monochromatic X-ray E V The larger the value, the smaller the difference becomes. Therefore, the increase in noise in the virtual monochromatic X-ray image due to noise in the bone image is suppressed. On the other hand, the energy E of the virtual monochromatic X-ray V The smaller μ becomes, B (E) and μ S As the difference in (E) increases, the contrast of the virtual monochromatic X-ray image increases. That is, the energy E of the virtual monochromatic X-ray image increases. V There is an appropriate value for it.

[0071] In this embodiment, a virtual monochromatic X-ray image was generated from the bone thickness B and the soft tissue thickness S, but the embodiment is not limited to this. As described above, the effective atomic number Z and the surface density D may be calculated first, and then the virtual monochromatic X-ray image may be generated using the effective atomic number Z and the surface density D. Also, multiple energies E V A composite X-ray image may be generated by combining multiple virtual monochromatic X-ray images produced by the same method. A composite X-ray image is an image that is expected to be obtained when irradiated with X-rays of an arbitrary spectrum.

[0072] Furthermore, although the image processing in this embodiment generated a virtual monochromatic X-ray image, it is not limited to this form. The bone thickness image B and the soft tissue thickness image S may be displayed as they are. Alternatively, the bone thickness image B and the soft tissue thickness image S may be displayed after applying a time-direction filter such as a recursive filter or a spatial-direction filter such as a Gaussian filter. Alternatively, a DSA image (Digital Subtraction Angiography) of the bone may be obtained using low-energy images (attenuation rate) and high-energy images (attenuation rate) before and after the injection of contrast agent, and this may be displayed. In other words, the image processing in this embodiment can be said to be a process that performs arbitrary calculations on the image after signal processing.

[0073] DSA images are acquired, for example, as follows: First, before injecting the contrast agent, X-ray imaging is performed to obtain an image of the attenuation rate at low energy L. M Image H of the attenuation rate at high energy M Obtain image L. M and image H M Bone thickness mask image B M and mask image S of soft tissue thickness M Next, we obtain the attenuation rate image L at low energy, which was taken after the injection of the contrast agent. L Image H of the attenuation rate at high energy L Live image B of bone thickness L Live image S of the thickness of soft tissue L To find the bone thickness (live image B) L Bone thickness mask image B M By pulling, the DSA image of the bone B DSA You can obtain this.

[0074] The energy subtraction process in this embodiment consists of three steps: correction processing, signal processing, and image processing, as shown in Figures 5 to 7. At this time, the bone thickness B and soft tissue thickness S obtained by solving equation [Equation 4] are used as estimated thickness values. The thickness measured by a measuring device is used as the true thickness value. If the correction processing and signal processing are performed appropriately, the estimated thickness value and the true thickness value should match. However, the inventors of this application have found that the estimated thickness value obtained by the above energy subtraction process does not necessarily match the true thickness value. If the error between the estimated thickness value and the true thickness value becomes large, artifacts may occur in the image after image processing.

[0075] As a result of the inventor's investigation, it was found that the causes of errors include scattered radiation, dose dependence of the attenuation rate, and the relationship between the attenuation rate, thickness, and X-ray spectrum. In this embodiment, a method for reducing errors caused by the X-ray spectrum is proposed.

[0076] Figure 8 is a diagram illustrating the principle of time-resolved X-ray spectrum according to this embodiment. Time-resolved X-ray spectrum refers to acquiring X-ray photon energy (radiation photon energy) in a time series and dividing the acquired X-ray photon energy (radiation photon energy) in the time direction. The X-ray photon energy divided in the time direction is called energy information. Here, as an example, the waveform of a convex X-ray pulse is used. The X-ray generator 101 (radiation generator) generates radiation by switching the tube voltage. When generating a convex X-ray pulse, the X-ray generator 101 (radiation generator) generates an X-ray pulse (convex X-ray pulse) by switching between the first tube voltage, a second tube voltage that is higher than the first tube voltage, and the first tube voltage.

[0077] Figure 8(A) shows the relationship between time and tube voltage. In Figure 8(A), 801 to 803 show X-ray waveforms with time on the horizontal axis and tube voltage on the vertical axis. Section 811 represents the low-energy section, section 812 represents the high-energy section, and section 813 represents the low-energy section.

[0078] Figure 8(B) shows the relationship between time and the energy of X-ray photons. In Figure 8(B), 804-806 show the energy information of each individual X-ray photon, with time on the horizontal axis and the energy of the X-ray photon on the vertical axis. The signal processing unit 133 divides the X-ray photon energy at two timings: one for sampling and holding the signal corresponding to the first radiation energy (low energy) (e.g., SH_N), and another for sampling and holding the signal corresponding to the second radiation energy (e.g., SH_S), which is higher than the first radiation energy.

[0079] 804 and 806 represent the X-ray photon group (low-energy X-ray photon group) in the low-energy sections 811 and 813 of the convex X-ray pulse. Similarly, 805 represents the X-ray photon group (high-energy X-ray photon group) in the high-energy section 812 of the convex X-ray pulse.

[0080] Figure 8(C) shows the relationship between X-ray energy and the number of X-ray photons. By aggregating the energy information of X-ray photons in X-ray photon groups 804 and 806, and plotting the X-ray energy on the horizontal axis and the number of X-ray photons on the vertical axis, a low-energy X-ray spectrum 807 can be obtained as shown in Figure 8(C). Similarly, by aggregating the energy information of X-ray photons in X-ray photon group 805, and plotting the X-ray energy on the horizontal axis and the number of X-ray photons on the vertical axis, a high-energy X-ray spectrum 808 can be obtained. The shape of the X-ray spectrum 807 corresponds to the X-ray energies (801, 803) in the aggregation interval (811, 813). The shape of the X-ray spectrum 808 corresponds to the X-ray energies (802) in the aggregation interval (812).

[0081] As described above, by time-resolved analysis of time-X-ray photon energy (time-series X-ray photon energy) data, time-varying X-ray spectra 807 and 808 within a single pulse can be obtained.

[0082] Figure 9 shows the relationship between the acquisition of the X-ray spectrum and the driving timing of the X-ray imaging system according to this embodiment. The driving of the X-ray imaging system is as described in Figure 4, but in Figure 9, the illustrations of the reset and readout timings shown in Figure 4 are omitted to avoid repetition of the explanation. Also, the explanation of reference numerals that are the same as those in Figure 4 is omitted.

[0083] Figure 9, section 901, shows the relationship between X-rays, synchronization signal, sample-hold SH_N, SH_S, and X-ray photon energy (time - X-ray photon energy), with time on the horizontal axis. Also, as shown in Figure 8, sections 807 and 808 show the distribution of the X-ray spectrum, with X-ray energy on the horizontal axis and the number of X-ray photons on the vertical axis. Section 807 shows the distribution of low-energy X-ray spectra, while section 808 shows the distribution of high-energy X-ray spectra.

[0084] During the X-ray irradiation period, sampling and holding are performed according to the synchronization signal, and the energy of each individual X-ray photon (time-X-ray photon energy) is measured. The timing of the time-resolved analysis of the acquired time-X-ray photon energy is set to coincide with the sampling and holding SH_N and SH_S in the operation of the X-ray imaging device 104, as shown by the dotted line. As a result, the X-ray energy of the image obtained by sampling and holding (sub-image) matches the X-ray energy of the time-resolved X-ray spectrum. That is, measured X-ray spectra of multiple images (sub-images (H, L)) with different X-ray energies are obtained. Here, as shown in Figure 4, sub-image 405 is defined as image 405, which corresponds to the sum of low-energy X-rays 401 and low-energy X-rays 403. Also, sub-image 406 is defined as image 406, which corresponds to the signals of low-energy X-rays 401, high-energy X-rays 402 and low-energy X-rays 403.

[0085] Here, image 405 (sub-image 405) corresponds to image L of the attenuation rate at low energy in equation [Equation 4], and image 406 (sub-image 406) corresponds to image H of the attenuation rate at high energy. Also, the low-energy X-ray spectrum 807 shown in Figures 8 and 9 corresponds to the low-energy X-ray spectrum N in equation [Equation 4]. L (E) corresponds to the high-energy X-ray spectrum 808 in equation [4], which corresponds to the spectrum N in high-energy X-rays. H Corresponds to (E). Linear attenuation coefficient μ of bone at energy E. B (E) and the linear attenuation coefficient μ of soft tissue at energy E S (E) is obtained from databases such as NIST.

[0086] By solving the simultaneous equations in [Equation 4] using sub-image 405 and X-ray spectrum 807, and sub-image 406 and X-ray spectrum 808, a highly accurate energy subtraction image can be obtained. For example, as material separation images, a highly accurate image B of bone thickness and an image S of soft tissue thickness can be obtained.

[0087] For measuring time-X-ray photon energy, an X-ray photon detector such as cadmium telluride (CdTe) can be used. It is also possible to use an X-ray spectrometer with a built-in CdTe detector. Normally, an X-ray spectrometer automatically aggregates X-ray photon energy information and outputs an X-ray spectrum, but by acquiring the stored data before aggregation, time-X-ray photon energy data can be obtained. In this embodiment, the X-ray photon detector or X-ray spectrometer can function as an acquisition unit that acquires X-ray photon energy (radiation photon energy) in a time series.

[0088] To improve the timing consistency between the time-resolved X-ray photon group and the sample-and-hold timing of the X-ray imaging device 104, the signal processing unit 133 performs signal processing to synchronize the X-ray generator 101, the X-ray imaging device 104, and the X-ray photon detector. For example, the synchronization signal 407 used to synchronize the X-ray generator 101 and the X-ray imaging device 104 may be used for synchronizing the X-ray photon detection. Through the signal processing of the signal processing unit 133, the X-ray photon detector acquires X-ray photon energy (radiation photon energy) in synchronization with the irradiation of radiation based on the time-varying tube voltage.

[0089] When acquiring an X-ray spectrum, the intensity of the X-rays may be limited by conditions such as the tube voltage, tube current, and irradiation time of the X-ray generator 101, as well as constraints due to pile-up in the X-ray photon detector. The energy information of X-ray photons obtained from one pulse of X-rays is only about tens to hundreds of them, and it is preferable to increase the amount of energy information to obtain a more accurate X-ray spectrum.

[0090] Therefore, when X-rays are irradiated multiple times, the signal processing unit 133 obtains energy information by summing the X-ray photon energies, which are divided at the timing of sampling and holding the X-ray energy signal, according to the X-ray energy.

[0091] The X-ray photon detector is capable of acquiring the X-ray photon energy in a time series after X-ray irradiation at least once or multiple times. When X-rays are irradiated multiple times, the signal processing unit 133 acquires energy information by summing the X-ray photon energies, which are divided at the timing of sample-holding the signal corresponding to the X-ray energy, according to the energy of the radiation.

[0092] For example, as shown in Figures 902 and 903, the information from multiple X-ray pulses may be summed by energy to generate X-ray spectra 807 and 808. In this case, the timing of the time-resolved X-ray pulses (division timing) must be consistent for each irradiation. Therefore, when X-rays are irradiated multiple times, the signal processing unit 133 divides the X-ray photon energy at a consistent timing for each irradiation. For example, it is preferable to synchronize the time-resolved timing for each irradiation by resetting the timer of the X-ray photon detector after each X-ray irradiation. If the energy information of the X-ray photons is insufficient even after summing the information from multiple X-ray pulses by energy, and the shape of the resulting X-ray spectrum is rough, it may be corrected to a smooth shape by smoothing correction.

[0093] The multiple X-ray spectra obtained by the above method of acquiring X-ray spectra are measured values ​​and may contain erroneous X-ray photon energy information due to pile-up in the X-ray photon detector. Therefore, it is preferable to remove only the erroneous X-ray photon energy information by pile-up correction before using it for energy subtraction calculations. It is preferable to perform this pile-up correction on each spectrum after time division. The signal processing unit 133 can perform a correction (pile-up correction) to exclude the X-ray spectrum in the excess region if the X-ray spectrum (low-energy spectrum, high-energy spectrum) contains energy spectra in the excess region that exceeds the energy range of the irradiated X-rays. For example, if the spectrum contains information of spectra counted in the excess region that exceeds the upper or lower limit of the energy range of the irradiated X-rays, the signal processing unit 133 performs a pile-up correction to exclude the energy spectrum in the excess region from the X-ray spectrum, treating the information of spectra counted in this excess region as an error caused by pile-up in the X-ray photon detector. Note that the above smoothing correction may be performed after the pile-up correction.

[0094] Figure 10 shows a flowchart of the processing flow in the X-ray imaging system according to the first embodiment. The flowchart in Figure 10 shows an example where the acquisition of the X-ray spectrum, as explained in Figures 8 and 9, is performed before image acquisition. Note that the timing of X-ray spectrum acquisition is not limited to the processing flow in Figure 10.

[0095] (S1001: Acquisition of time-X-ray photon energy information) In S1001, the X-ray photon detector acquires information on time-series X-ray photon energy.

[0096] (S1002: Time-X-ray photon energy information decomposition) Next, in S1002, the signal processing unit 133 divides the time-X-ray photon energy information acquired by the X-ray photon detector. The signal processing unit 133 divides the time-X-ray photon energy information into X-ray photon groups 804, 805, and 806, for example, as shown in Figure 8, in synchronization with the timing of the sample-and-hold (SH_N, SH_S). X-ray photon groups 804 and 806 correspond to the low-energy sections 811 and 813 of the convex X-ray pulse, and X-ray photon group 805 corresponds to the high-energy section 812 of the convex X-ray pulse.

[0097] (S1003: Aggregation, multiple spectra obtained) In S1003, the signal processing unit 133 aggregates the divided X-ray photon energy information and acquires multiple spectra with different X-ray energies. Using the X-ray photon energy information of X-ray photon groups 804 and 806 corresponding to the low-energy sections 811 and 813, and the X-ray photon energy information of X-ray photon group 805 corresponding to the high-energy section 812, the signal processing unit 133 acquires, for example, a low-energy X-ray spectrum 807 and a high-energy X-ray spectrum 808, as shown in Figure 8 or Figure 9. The signal processing unit 133 then stores the acquired data of multiple spectra with different X-ray energies in the memory of the imaging control device 103.

[0098] Here, among the acquired spectral data, for example, the low-energy X-ray spectrum 807 shown in Figure 8 or Figure 9 is the spectrum N for low-energy X-rays in equation [Equation 4]. L Corresponding to (E), the high-energy X-ray spectrum 808 is the spectrum N in high-energy X-rays in equation [Equation 4]. H Corresponds to (E).

[0099] (S1004: Time-division imaging) In S1004, image acquisition is performed by time-resolved imaging using an X-ray imaging system. The signal processing unit 133 of the X-ray imaging system irradiates and takes images by time-resolved imaging based on, for example, the timing chart in Figure 4 or Figure 9, and acquires multiple images (sub-images: H, L) corresponding to multiple different radiation energies.

[0100] (S1005: Energy subtraction process) Then, in S1005, the signal processing unit 133 performs energy subtraction processing using the multiple images (sub-images: H, L) corresponding to the multiple radiation energies obtained in S1004 and the data of multiple spectra with different X-ray energies stored in memory in S1003.

[0101] By solving the simultaneous equations in [Equation 4] using sub-image 405 and X-ray spectrum 807, and sub-image 406 and X-ray spectrum 808, a highly accurate energy subtraction image can be obtained. For example, as material separation images, a highly accurate image B of bone thickness and an image S of soft tissue thickness can be obtained.

[0102] Furthermore, the procedure for acquiring the X-ray spectrum may be performed before the product is provided to the customer. For example, the service department may acquire the data in advance and store it in memory before shipment. Multiple patterns of X-ray spectrum data may be prepared by changing the imaging conditions and stored in memory, and the X-ray spectrum data may be changed according to the imaging conditions. Preferably, the imaging conditions include a range of X-ray irradiation conditions (e.g., tube voltage, tube current, storage time, etc.) and sample-hold conditions that are expected to be used at the customer's site. Examples of changes to the imaging conditions are not limited to X-ray irradiation conditions and sample-hold conditions, but various other conditions may be set.

[0103] According to this embodiment, the energy spectrum of the irradiated radiation can be obtained. By using the obtained energy spectrum in energy subtraction processing, the accuracy of the processing can be improved.

[0104] [Second Embodiment] In the second embodiment, in addition to the configuration for time-resolved X-ray spectrum analysis, a configuration for optimizing imaging conditions will be described. The configuration and operation of the X-ray imaging system in the second embodiment are the same as in the first embodiment.

[0105] Figure 11 shows an example of the X-ray waveform described in this embodiment. While the processing of this embodiment can be applied to various X-ray waveforms, here we will explain using a convex X-ray waveform as an example. Ideally, for convex X-rays, the voltage changes perpendicularly at the timing of the tube voltage switch, as shown in Figure 11(A), and remains constant otherwise. The image quality of the energy subtraction image in an X-ray imaging system depends on the X-ray energy difference ΔE and dose ratio between the subimages. By performing a sample-and-hold at the voltage switch timing, the X-ray energy difference ΔE and dose ratio can be easily determined, and an energy subtraction image with the expected image quality can be obtained.

[0106] However, in reality, as shown in the waveform in Figure 11(B), the voltage does not change perpendicularly at the timing of the tube voltage switch, and the rise and fall of the voltage are blunted. In the case of such a blunted waveform, it may be difficult to grasp the energy difference ΔE and dose ratio of the subimage. Furthermore, since the energy difference ΔE and dose ratio change significantly even by changing the sample-hold timing by just a few ms, setting the sample-hold timing can also be difficult.

[0107] Therefore, in this embodiment, the X-ray spectrum is time-resolved, and the image quality obtained by the energy subtraction process at that time is simulated to optimize the imaging conditions, such as the timing of the sample hold. Hereinafter, this simulation will also be referred to as the image quality simulation.

[0108] Figure 12 shows the principle of image quality simulation. In Figure 12, 1201 to 1203 are diagrams showing X-ray photon energy information and sample-hold timing, with time on the horizontal axis and X-ray photon energy on the vertical axis. The only difference between 1201 to 1203 is the timing of the sample-hold SH_N and SH_S, shown by the dotted lines. In this embodiment, CNR / √Dose is used as evaluation information (index) to indicate the image quality of the image acquired by image quality simulation. CNR / √Dose is the value obtained by dividing the ratio of the contrast between the region where the target substance is present and the region where the target substance is not present (contrast to noise ratio: CNR) to the noise in the image acquired by energy subtraction processing by the square root of the dose (Does). It can indicate the visibility of the target substance and can be suitably used as an index of the energy subtraction image. Here, the target substance refers to the substance contained in the subject that is separated by energy subtraction processing (for example, bone and soft tissue may be included).

[0109] In Figure 12, evaluation information 1204-1206 indicates the image quality obtained by energy subtraction processing using multiple sub-images (two images: a high-energy image H and a low-energy image L) obtained under the shooting conditions 1201-1203.

[0110] If energy information of X-ray photons is available, information such as the number of X-ray photons and the energy of the X-rays can be obtained not only from the X-ray spectrum of the subimage when the sample hold timing is changed. When determining evaluation information (CNR / √Dose), noise (N) and dose (Dose) can be determined from the number of X-ray photons and energy. Furthermore, the contrast can be determined by substituting the X-ray spectrum and X-ray energy information into equation [Equation 4] and solving the simultaneous equations to obtain the contrast between the region where the target material is present and the region where the target material is not present in the acquired image. Linear attenuation coefficient μ in equation [Equation 4] B (E) Linear attenuation coefficient μ of soft tissue S (E) The types and thicknesses of materials such as bone thickness B and soft tissue thickness S can be arbitrarily changed to suit the subject being photographed.

[0111] Therefore, if energy information of X-ray photons is available, the image quality of the image acquired by energy subtraction processing can be simulated. By changing the sample-hold timing as an imaging condition, for example, as shown in Figure 12, 1201-1203, and calculating the evaluation information CNR / √Dose (for example, 1204-1206), the sample-hold timing can be optimized by using imaging conditions that maximize the evaluation information CNR / √Dose, thereby obtaining an energy subtraction image with good image quality.

[0112] Figure 13 shows the correlation between the sample-and-hold timing obtained from the simulation and the image quality of the image acquired by energy subtraction processing. The horizontal axis represents the SH_S timing, and the vertical axis represents the evaluation information CNR / √Dose. Series 1302 in the graph represents different SH_N timings. The condition of the maximum value of the evaluation information 1301 in the vertical axis direction becomes the optimal sample-and-hold (SH_N, SH_S) condition. By acquiring images at the sample-and-hold timing of the maximum value of the evaluation information 1301, and using the time-resolved X-ray spectrum acquired by the method described in the first embodiment, images with good image quality can be obtained by energy subtraction processing.

[0113] Figure 14 shows a flowchart of the sample hold timing optimization process in the X-ray imaging system according to the second embodiment. In Figure 14, the processes S1402 to S1405 enclosed by dashed lines are simulation processes.

[0114] (S1401: Acquisition of time-X-ray photon energy information) In S1401, the X-ray photon detector acquires information on time-series X-ray photon energy.

[0115] (S1402: Time-X-ray photon energy information is divided and the division position is changed) Next, in S1402, the signal processing unit 133 processes the time-X-ray photon energy information acquired by the X-ray photon detector using a set sample-and-hold (SH_N) control. i SH_S j The sample is split at the timing of ). Here, the subscripts i and j are parameters that change the sample-hold setting in the loop. By changing the parameters to i=i+1..., j=j+1..., etc., the sample-hold (SH_N i SH_S jThe timing of ) can be changed. In a loop, for example, changing parameter i changes the setting of SH_N for 1302 in Figure 13, and changing parameter j changes the setting of SH_S for the horizontal axis in Figure 13.

[0116] The signal processing unit 133 processes the time-X-ray photon energy information using a set sample-and-hold (SH_N) method. i SH_S j Based on the timing of ), the X-ray photons are divided into groups (e.g., 804, 805, 806).

[0117] (S1403: Aggregation, multiple spectra obtained) In S1403, the signal processing unit 133 aggregates the information of the divided X-ray photon energies and acquires multiple spectra with different X-ray energies. The signal processing unit 133 then stores the data of the multiple spectra with different X-ray energies that it has acquired into the memory of the imaging control device 103.

[0118] (S1404: Image Quality Evaluation) In S1404, time-resolved imaging using an X-ray imaging system is performed to irradiate and image the device, acquiring multiple images (sub-images) corresponding to multiple different radiation energies. The signal processing unit 133 then performs energy subtraction processing using the acquired multiple images (sub-images: H, L) and data from multiple spectra with different X-ray energies, calculates evaluation information for the acquired image (energy subtraction image: energy subtraction image), and stores it in memory.

[0119] The signal processing unit 133 performs a sample-and-hold operation (SH_N i SH_S j The settings of the sample-and-hold (SH_N) are changed and the process of S1402 to S1404 is repeatedly executed. The signal processing unit 133 changes the settings of the sample-and-hold (SH_N) and executes the process repeatedly. i SH_S jBased on this, the time-X-ray photon energy information is divided, the divided X-ray photon energy information is aggregated, and multiple spectra with different X-ray energies are obtained. Then, the signal processing unit 133 performs energy subtraction processing, calculates evaluation information for the acquired image (energy subtraction image: EneSub image), and stores it in memory.

[0120] (S1405: Determine the optimal sample-hold timing) In S1405, the signal processing unit 133 changes the timing of sample-holding the X-ray energy signal as an imaging condition to set the timing at which the evaluation information reaches its maximum value. The signal processing unit 133 compares multiple evaluation information obtained by repeatedly executing the processes in S1402 to S1404 and determines the optimal sample-hold timing at which the evaluation information CNR / √Dose is maximized. Using the optimal sample-hold timing obtained by simulation, the following time-resolved imaging (S1406) is performed.

[0121] (S1406: Time-division imaging) In S1406, based on the optimal sample-hold timing, images are acquired by time-resolved imaging using the X-ray imaging system. The X-ray imaging system irradiates with radiation and takes images by time-resolved imaging, and the signal processing unit 133 acquires multiple images (sub-images) corresponding to multiple different radiation energies.

[0122] (S1407: Energy subtraction process) In S1407, the signal processing unit 133 performs energy subtraction processing using multiple images (sub-images) corresponding to multiple radiation energies and data from multiple spectra with different X-ray energies acquired at the optimal sample-hold timing. By solving the system of equations [Equation 4] using the image L (sub-image) of the attenuation rate at low energy and the low-energy X-ray spectrum, and the image H (sub-image) of the attenuation rate at high energy and the high-energy X-ray spectrum, a highly accurate energy subtraction image can be obtained. For example, as material separation images, a highly accurate image B of bone thickness and an image S of soft tissue thickness can be obtained.

[0123] According to this embodiment, the energy spectrum of the irradiated radiation can be obtained. By using the obtained energy spectrum in energy subtraction processing, the accuracy of the processing can be improved.

[0124] [Third Embodiment] In the third embodiment, a configuration is described in which the timing of the sample hold and the X-ray irradiation conditions are optimized as imaging conditions by simulating the image quality obtained by time-resolved X-ray spectrum analysis and energy subtraction processing. The configuration and operation of the X-ray imaging system in the third embodiment are the same as in the first embodiment, and the configuration for evaluating and simulating the image quality obtained by energy subtraction processing is the same as in the second embodiment.

[0125] Figure 15 shows a flowchart of the optimization process for sample hold timing and X-ray irradiation conditions in the X-ray imaging system according to the third embodiment. The difference from the flowchart in Figure 14 is the addition of a loop for changing the X-ray irradiation conditions. In the processing flow of Figure 15, optimization of the X-ray irradiation conditions is performed in addition to the sample hold timing. In Figure 15, the processes S1502 to S1506 enclosed by dashed lines are simulation processes.

[0126] (S1501: Acquisition of time-X-ray photon energy information) In S1501, the X-ray photon detector acquires information on time-series X-ray photon energy. Here, the X-ray irradiation conditions (k) in the X-ray generator 101 are set. Here, k is a parameter that changes the setting of the X-ray irradiation conditions in the iterative process. By changing the parameter k = k + 1, etc., the X-ray irradiation conditions can be changed.

[0127] (S1502: Time-X-ray photon energy information is divided, and the division position is changed.) Next, in S1502, the signal processing unit 133 processes the time-X-ray photon energy information acquired by the X-ray photon detector using a set sample-and-hold (SH_N) control. i SH_S j The program is split at the following timing.

[0128] (S1503: Aggregation, multiple spectra acquired) In S1503, the signal processing unit 133 aggregates the information of the divided X-ray photon energies and acquires multiple spectra with different X-ray energies. The signal processing unit 133 then stores the data of the multiple spectra with different X-ray energies that it has acquired into the memory of the imaging control device 103.

[0129] (S1504: Image Quality Evaluation) In S1504, time-resolved imaging using an X-ray imaging system is performed to irradiate and image the device, acquiring multiple images (sub-images) corresponding to multiple different radiation energies. The signal processing unit 133 then performs energy subtraction processing using the acquired multiple images (sub-images: H, L) and data from multiple spectra with different X-ray energies, calculates evaluation information for the acquired image (energy subtraction image: energy subtraction image), and stores it in memory.

[0130] The signal processing unit 133 performs a sample-and-hold operation (SH_N i SH_S jThe settings of ) are changed and processes S1502 to S1504 are executed repeatedly.

[0131] (S1505: Determine the optimal sample-hold timing) In S1505, the signal processing unit compares multiple evaluation pieces of information obtained by repeatedly executing the processes in S1502 to S1504, and determines the sample-and-hold timing that maximizes the evaluation information CNR / √Dose as the optimal sample-and-hold timing.

[0132] Then, the signal processing unit 133 returns to processing S1501, changes the X-ray irradiation conditions (k), and repeats the same processing. The signal processing unit 133 sets the sample-hold timing that maximizes the evaluation information CNR / √Dose, calculates the evaluation information CNR / √Dose under the changed X-ray irradiation conditions (k), and stores it in memory.

[0133] (S1506: Determine the optimal X-ray irradiation conditions) In S1506, the signal processing unit 133 evaluates the image quality obtained by simulating energy subtraction processing with changed imaging conditions based on evaluation information, and sets imaging conditions that maximize the evaluation information. In this step, the signal processing unit 133 changes the X-ray irradiation conditions as imaging conditions and sets irradiation conditions that maximize the evaluation information. The signal processing unit 133 repeatedly performs the process of calculating the evaluation information CNR / √Dose under the condition of changing the X-ray irradiation conditions (k), compares multiple acquired evaluation information, and finds the X-ray irradiation conditions (k) that maximize the evaluation information CNR / √Dose as the optimal X-ray irradiation conditions. Using the optimal sample-hold timing (S1505) and optimal X-ray irradiation conditions (S1506) obtained by simulation, the following time-resolved imaging (S1507) is performed.

[0134] (S1507: Time-division imaging) In S1507, time-resolved imaging is performed under optimal imaging conditions (X-ray irradiation conditions, sample hold timing). That is, based on the optimal X-ray irradiation conditions and optimal sample hold timing, images are acquired by time-resolved imaging using the X-ray imaging system. The X-ray imaging system irradiates with radiation and takes images by time-resolved imaging, and the signal processing unit 133 acquires multiple images (sub-images) corresponding to multiple different radiation energies.

[0135] (S1508: Energy subtraction process) In S1508, the signal processing unit 133 performs energy subtraction processing using multiple images (sub-images) corresponding to multiple radiation energies and data from multiple spectra with different X-ray energies acquired under optimal imaging conditions (X-ray irradiation conditions, sample hold timing).

[0136] By solving the system of equations [Equation 4] using the low-energy attenuation rate image L (secondary image) and low-energy X-ray spectrum, and the high-energy attenuation rate image H (secondary image) and high-energy X-ray spectrum, a highly accurate energy subtraction image can be obtained. For example, as material separation images, a highly accurate image of bone thickness B and an image of soft tissue thickness S can be obtained.

[0137] According to this embodiment, the energy spectrum of the irradiated radiation can be obtained. By using the obtained energy spectrum in energy subtraction processing, the accuracy of the processing can be improved, and an energy subtraction image with better image quality can be obtained.

[0138] (modified version) In the first to third embodiments, the X-ray imaging device 104 was an indirect type X-ray sensor using a phosphor. However, the disclosed technology is not limited to this form. For example, a direct type X-ray sensor using a direct conversion material such as CdTe may be used.

[0139] The first to third embodiments described the case where two X-ray energies, a first energy level and a second energy level higher than the first energy level, are used. However, the disclosed technology is not limited to such embodiments. For example, it is also applicable when there are three or more X-ray energy levels.

[0140] In the first to third embodiments, images with different energy levels were obtained by changing the energy of the X-rays. However, the disclosed technology is not limited to this form. By stacking multiple phosphors 105 and two-dimensional detectors 106, a configuration may be used to obtain images with different energy levels from a two-dimensional detector on the front and a two-dimensional detector on the back with respect to the direction of X-ray incidence. The two-dimensional detector 106 is not limited to medical use, but may also be an industrial two-dimensional detector.

[0141] In the first to third embodiments, energy subtraction processing was performed using the imaging control device 103 of the radiography system. However, the disclosed technology is not limited to these forms. Images acquired by the imaging control device 103 may be transferred to another computer for energy subtraction processing. For example, acquired images may be transferred to another personal computer (image viewer) via a medical PACS, energy subtraction processing may be performed, and then the results of the processing may be displayed on the radiography system.

[0142] In the first to third embodiments, an example of convex X-rays was described when the X-ray tube voltage was actively switched. However, the embodiments are not limited to this example, and for example, the tube voltage may be switched in a stepwise manner. The number of steps may be two steps, or three or more steps. Also, the tube voltage may be switched to increase, like an ascending staircase, or to decrease, like a descending staircase. For example, when the tube voltage increases, like an ascending staircase, the X-ray generator 101 generates X-rays by switching at least between a first tube voltage and a second tube voltage that is higher than the first tube voltage. Also, when the tube voltage decreases, like a descending staircase, the X-ray generator 101 generates X-rays by switching at least between a second tube voltage and a first tube voltage.

[0143] The sample-hold and time-resolved X-ray photon energy can be increased or decreased in accordance with the number of steps in the staircase. For example, as explained in Figure 4, the highest tube voltage may be set as the median value, the tube voltages to the left of the median value may be set to increase in a stepwise manner, and the tube voltages to the right of the median value may be set to decrease in a stepwise manner.

[0144] Furthermore, in the first to third embodiments, the tube voltage of the X-ray generator 101 (radiation generator) was changed. However, the disclosed technology is not limited to this form. The energy of the X-rays irradiated onto the X-ray imaging device 104 may be changed by, for example, switching the filter of the X-ray generator 101 over time.

[0145] The first to third embodiments described the technology using medical applications such as the separation of bone and soft tissue as examples. However, the disclosed technology is not limited to these forms. For example, it can also be applied to industrial applications such as the inspection of defective circuit boards.

[0146] Furthermore, in the second and third embodiments, CNR / √Dose was used as an index for evaluating the image quality of the energy subtraction image. However, the disclosed technology is not limited to embodiments using this index. For example, CNR may be used as an index without considering the dose. Alternatively, parameters other than image quality, such as the energy difference ΔE or dose ratio of the secondary images (image H and image L), may be used as an index. For example, the signal processing unit 133 may set imaging conditions in which evaluation information using the energy difference of the irradiated X-ray energies when acquiring multiple images (image H and image L), or evaluation information using the dose ratio of the irradiated X-rays when acquiring multiple images, is maximized.

[0147] The disclosures herein include the following radiation imaging devices, information processing devices, information processing methods (radiation imaging methods), and programs.

[0148] (Item 1) A radiation imaging apparatus equipped with processing means that aggregates energy information obtained by dividing the radiation photon energy acquired in a time series in the time direction and performs a process to acquire an energy spectrum.

[0149] (Item 2) The system includes a processing means that performs energy subtraction processing using multiple images corresponding to multiple different radiation energies obtained by irradiating a subject with radiation and taking photographs, and an energy spectrum obtained based on the radiation photon energy acquired in time series. The processing means is A radiation imaging device that, at the timing of sampling and holding the radiation energy signal, aggregates energy information obtained by dividing the radiation photon energy and acquires the energy spectrum.

[0150] (Item 3) The radiation imaging apparatus according to Item 1 or 2, wherein the processing means acquires the energy spectrum by counting the number of photons in the time series in the energy information.

[0151] (Item 4) The radiation imaging apparatus according to any one of Items 1 to 3, wherein the processing means divides the radiation photon energy between a timing for sampling and holding a signal corresponding to a first radiation energy and a timing for sampling and holding a signal corresponding to a second radiation energy higher than the first radiation energy.

[0152] (Item 5) The acquisition means further comprises acquiring the radiation photon energy in a time series after the radiation has been irradiated at least once or more times, The processing means is a radiation imaging device according to any one of items 1 to 4, wherein when the radiation is irradiated multiple times, the radiation photon energy, which is divided at the timing for sampling and holding the signal corresponding to the energy of the radiation, is summed up according to the energy of the radiation to acquire the energy information.

[0153] (Item 6) The processing means is a radiation imaging device according to item 5, wherein when the radiation is irradiated multiple times, the radiation photon energy is divided at the same timing for each irradiation.

[0154] (Item 7) A radiation imaging apparatus according to any one of Items 1 to 4, further comprising acquisition means for acquiring the radiation photon energy in synchronization with radiation irradiation based on a time-varying tube voltage.

[0155] (Item 8) The radiation imaging apparatus according to any one of Items 1 to 7, wherein the processing means performs pile-up correction to exclude the energy spectrum of an excess region that exceeds the energy range of the radiation when the energy spectrum includes such excess region.

[0156] (Item 9) The radiation imaging apparatus according to any one of Items 1 to 8, wherein the processing means performs smoothing correction to correct the waveform of the energy spectrum to a smooth shape.

[0157] (Item 10) Further comprising a radiation generating means that generates radiation by switching the tube voltage, The aforementioned radiation generating means is The radiation is generated by switching between the first tube voltage, a second tube voltage higher than the first tube voltage, and the first tube voltage, or At a minimum, the first tube voltage and the second tube voltage are switched to generate the radiation, or A radiation imaging device according to any one of items 1 to 9, wherein at least the second tube voltage and the first tube voltage are switched to generate the radiation.

[0158] (Item 11) The radiation imaging apparatus according to Item 2, wherein the processing means evaluates the image quality of an image obtained by simulating the energy subtraction processing with changed shooting conditions based on evaluation information, and sets the shooting conditions that result in the maximum value of the evaluation information.

[0159] (Item 12) The processing means is a radiation imaging apparatus according to Item 11, which changes the timing for sampling and holding the radiation energy signal as the imaging condition, and sets the timing at which the evaluation information reaches its maximum value.

[0160] (Item 13) The processing means is a radiation imaging device according to item 11, which changes the radiation irradiation conditions as the imaging conditions to set the irradiation conditions that result in the evaluation information being at its maximum value.

[0161] (Item 14) The processing means is Evaluation information using the ratio of contrast to noise in the aforementioned image, or A radiation imaging device according to any one of items 11 to 13 for setting the imaging conditions such that the evaluation information, which uses the ratio of contrast to noise divided by the square root of the radiation dose, is the maximum value.

[0162] (Item 15) The processing means is Evaluation information using the energy difference of the multiple radiation energies irradiated when acquiring the multiple images, or A radiation imaging device according to any one of items 11 to 14 for setting the shooting conditions such that the evaluation information using the dose ratio of the radiation irradiated when acquiring the multiple images is the maximum value.

[0163] (Item 16) A radiation imaging apparatus comprising a processing means that obtains multiple images corresponding to multiple different radiation energies by irradiating a subject with radiation and taking photographs using an energy spectrum obtained by aggregating energy information obtained by dividing the radiation photon energy acquired in a time series in the time direction, and performs energy subtraction processing using the multiple images.

[0164] (Item 17) An information processing device equipped with processing means for aggregating energy information obtained by dividing the radiation photon energy acquired in a time series in the time direction and performing a process to acquire an energy spectrum.

[0165] (Item 18) The system includes a processing means that performs energy subtraction processing using multiple images corresponding to multiple different radiation energies obtained by irradiating a subject with radiation and taking photographs, and an energy spectrum obtained based on the radiation photon energy acquired in time series. The processing means is An information processing device that, at the timing of sampling and holding the radiation energy signal, aggregates energy information obtained by dividing the radiation photon energy and acquires the energy spectrum.

[0166] (Item 19) An information processing apparatus comprising a processing means that obtains multiple images corresponding to multiple different radiation energies by irradiating a subject with radiation and taking photographs using an energy spectrum obtained by aggregating energy information obtained by dividing radiation photon energy acquired in a time series in the time direction, and performs energy subtraction processing using the multiple images.

[0167] (Item 20) An information processing method comprising a processing step of aggregating energy information obtained by dividing the radiation photon energy acquired in a time series in the time direction and performing a process to obtain an energy spectrum.

[0168] (Item 21) The system includes a processing step that performs energy subtraction processing using multiple images corresponding to multiple different radiation energies obtained by irradiating a subject with radiation and taking photographs, and an energy spectrum obtained based on the radiation photon energy acquired in time series. In the aforementioned processing step, An information processing method for obtaining the energy spectrum by aggregating energy information obtained by dividing the radiation photon energy at the timing of sampling and holding the radiation energy signal.

[0169] (Item 22) An information processing method comprising a processing step of obtaining multiple images corresponding to multiple different radiation energies by using an energy spectrum obtained by aggregating energy information obtained by dividing the radiation photon energy acquired in a time series in the time direction, irradiating a subject with radiation and taking an image, and performing energy subtraction processing using the multiple images.

[0170] (Item 23) A program that causes a computer to function as a processing means for a radiographic imaging apparatus as described in any one of items 1 through 16.

[0171] [Other embodiments] The disclosed technology can also be implemented by supplying a program that implements one or more of the functions of the embodiments described above to a system or device via a network or storage medium, and by having one or more processors in the computer of that system or device read and execute the program. It can also be implemented by a circuit (e.g., an ASIC) that implements one or more functions.

[0172] The disclosed technology is not limited to the embodiments described above, and various modifications and variations are possible without departing from the spirit and scope of the invention. Accordingly, claims are attached to make the scope of the invention public. [Explanation of symbols]

[0173] 101: X-ray generator, 102: X-ray control device, 103: Imaging control device (information processing device), 104: X-ray imaging device (radiation imaging device)

Claims

1. By using an energy spectrum obtained by aggregating energy information obtained by dividing the radiation photon energy acquired in time series in the time direction, irradiating a subject with radiation and performing imaging to acquire a plurality of images corresponding to a plurality of different radiation energies, and comprising processing means for performing energy subtraction processing using the plurality of images. A radiation imaging apparatus.

2. The processing means At the timing of sample-holding the signal of the radiation energy, aggregates the energy information obtained by dividing the radiation photon energy, and acquires the energy spectrum. The radiation imaging apparatus according to claim 1.

3. The processing means acquires the energy spectrum by counting the number of photons in time series in the energy information. The radiation imaging apparatus according to claim 1 or 2.

4. The processing means divides the radiation photon energy at the timing of sample-holding the signal corresponding to the first radiation energy and at the timing of sample-holding the signal corresponding to the second radiation energy higher than the first radiation energy. The radiation imaging apparatus according to claim 1 or 2.

5. Further comprising acquisition means for acquiring the radiation photon energy irradiated with radiation at least once or a plurality of times in time series, When the radiation is irradiated a plurality of times, the processing means totals the radiation photon energy divided at the timing of sample-holding the signal corresponding to the energy of the radiation for each energy of the radiation to obtain the energy information. The radiation imaging apparatus according to claim 1 or 2.

6. When the radiation is irradiated a plurality of times, the processing means divides the radiation photon energy at the same timing for each irradiation. The radiation imaging apparatus according to claim 5.

7. The radiation imaging apparatus according to claim 1 or 2, further comprising acquisition means for acquiring the radiation photon energy in synchronization with irradiation of radiation based on a tube voltage that changes over time.

8. The radiation imaging apparatus according to claim 1 or 2, wherein when the energy spectrum in the excess region exceeding the energy region of the radiation is included in the energy spectrum, the processing means performs pile-up correction to exclude the energy spectrum of the excess region.

9. The radiation imaging apparatus according to claim 1 or 2, wherein the processing means performs smoothing correction to correct the waveform of the energy spectrum into a smooth shape.

10. The radiation imaging apparatus according to claim 1 or 2, further comprising radiation generation means for generating radiation by switching the tube voltage, wherein the radiation generation means generates the radiation by switching between a first tube voltage, a second tube voltage higher than the first tube voltage, and the first tube voltage, or generates the radiation by switching at least between the first tube voltage and the second tube voltage, or generates the radiation by switching at least between the second tube voltage and the first tube voltage.

11. The radiation imaging apparatus according to claim 2, wherein the processing means evaluates the image quality of an image obtained by simulation of the energy subtraction process with the imaging conditions changed, based on evaluation information, and sets the imaging conditions that maximize the evaluation information.

12. The radiation imaging apparatus according to claim 11, wherein the processing means changes, as the imaging condition, the timing for sample-holding the signal of the radiation energy, and sets the timing that maximizes the evaluation information.

13. The radiation imaging apparatus according to claim 11, wherein the processing means changes, as the imaging condition, the irradiation condition of the radiation, and sets the irradiation condition that maximizes the evaluation information.

14. The processing means: Evaluation information using the ratio of the contrast to the noise of the image, or Evaluation information using a value obtained by dividing the ratio of the contrast to the noise by the square root of the dose of the radiation, sets the imaging condition that maximizes the evaluation information. The radiation imaging apparatus according to claim 11.

15. The processing means: Evaluation information using the energy difference of the plurality of radiation energies irradiated when acquiring the plurality of images, or Evaluation information using the dose ratio of the radiation irradiated when acquiring the plurality of images, sets the imaging condition that maximizes the evaluation information. The radiation imaging apparatus according to claim 11.

16. An information processing apparatus including a processing means that irradiates a subject with radiation and performs imaging using an energy spectrum obtained by aggregating energy information obtained by dividing radiation photon energy acquired in time series in the time direction, thereby acquiring a plurality of images corresponding to a plurality of different radiation energies, and performs energy subtraction processing using the plurality of images.

17. An information processing method including a processing step of irradiating a subject with radiation and performing imaging using an energy spectrum obtained by aggregating energy information obtained by dividing radiation photon energy acquired in time series in the time direction, thereby acquiring a plurality of images corresponding to a plurality of different radiation energies, and performing energy subtraction processing using the plurality of images.

18. A program for causing a computer to function as the processing means of the radiation imaging apparatus according to claim 1.

Images