Radiation imaging processing device, method, and program

Abstract

To provide a radiation image processing device, a method, and a program capable of executing calibration of an imaging device at an appropriate timing.SOLUTION: A processor displays a subtraction image derived by executing weighting subtraction for a plurality of radiation images acquired by imaging a subject with radioactive rays with different energy distributions and subjected to predetermined image processing, in which a specific tissue in the subject is extracted, receives an instruction for changing a subtraction level corresponding to a weighting coefficient used when executing the weighting subtraction, determines whether or not an amount of change in the subtraction level has exceeded a predetermined threshold value, and issues a warning if the determination is affirmative.SELECTED DRAWING: Figure 3

Description

Technical Field

[0001] The present disclosure relates to a radiation image processing apparatus, method, and program.

Background Art

[0002] Conventionally, energy subtraction processing using two radiation images obtained by irradiating a subject with two types of radiation having different energy distributions is known by utilizing the fact that the attenuation amounts of radiation transmitted through substances constituting the subject are different. Energy subtraction processing is a method of obtaining an image in which specific structures included in a radiation image are extracted by corresponding each pixel of the two radiation images obtained as described above, multiplying appropriate weighting factors between the pixels, and then performing subtraction (subtraction).

[0003] In addition, a method has been proposed for estimating the bone density of a subject by correcting a bone image representing the bone part of the subject obtained by energy subtraction processing using a correction factor set based on the body thickness of the subject and the tube voltage (see Patent Document 1). Also, a method for removing the scattered radiation component of the radiation included in a radiation image using the body thickness of the subject has been proposed in order to improve the image quality of the radiation image (see Patent Document 2).

[0004] Here, the body thickness of the subject used for correcting the bone image and removing scattered radiation in the methods described in Patent Documents 1 and 2 above is derived based on the dose of radiation that directly reaches the radiation detector and the dose of radiation that reaches through the subject. The dose of radiation that directly reaches the radiation detector is calculated based on imaging conditions such as tube voltage (kV), dose (mAs), and the distance between the radiation source and the radiation detector (SID (Source-to-Image receptor Distance)).

Prior Art Documents

Patent Documents

[0005]

Patent Document 1

[0006] Incidentally, the sensitivity of the radiation tube and the radiation detector used in radiography equipment to detect radiation transmitted through the subject and generate a radiographic image deteriorates over time. When the tube and radiation detector deteriorate, the imaging conditions set in the radiography equipment deviate from the set values. When the imaging conditions deviate from the set values, the correction coefficient changes because the tube voltage and body thickness change in the method described in Patent Document 1, and as a result, it becomes impossible to accurately estimate bone density. Also, in the method described in Patent Document 2, it becomes impossible to accurately remove scattered radiation components. In this way, if bone density cannot be estimated accurately or scattered radiation components cannot be accurately removed, the image quality of the acquired image deteriorates. In this case, the set values ​​in the radiography equipment can be matched with the actual imaging conditions by performing calibration of the radiography equipment. However, it is difficult for the operator to grasp the appropriate timing for calibration.

[0007] This disclosure is made in light of the above circumstances and aims to enable calibration of the imaging device at an appropriate time. [Means for solving the problem]

[0008] The radiation image processing device according to this disclosure comprises at least one processor, The processor displays a subtraction image in which a specific tissue within the subject is extracted, derived by weighting and subtracting multiple radiation images obtained by photographing the subject with radiation of different energy distributions, and then subjected to predetermined image processing. It accepts instructions to change the subtraction level corresponding to the weight coefficient used when performing weighted subtraction. Determine whether the amount of change in the subtraction level exceeds a predetermined threshold. A warning is issued if the judgment is affirmative.

[0009] Furthermore, in the radiation image processing apparatus according to this disclosure, the processor acquires multiple radiation images, The thickness distribution of the subject is derived based on one of several radiographic images and the imaging conditions when all of the radiographic images were acquired. Based on the body thickness distribution and imaging conditions, the scattered radiation component of the radiation scattered by the subject is removed from multiple radiation images. By weighting and subtracting multiple radiation images from which scattered radiation components have been removed, a bone image is derived in which the bone portion of the subject is extracted. A correction coefficient is obtained to correct the pixel values ​​of bone images based on body thickness distribution and imaging conditions. The subtraction image may be derived by correcting the bone image using a correction coefficient.

[0010] Furthermore, in the radiation image processing apparatus according to this disclosure, the processor acquires multiple radiation images, The thickness distribution of the subject is derived based on one or more radiographic images and the imaging conditions when all of the radiographic images were acquired. Based on the body thickness distribution and imaging conditions, the scattered radiation component of the radiation scattered by the subject is removed from multiple radiation images. A subtraction image may also be derived by weighting and subtracting multiple radiation images from which scattered radiation components have been removed.

[0011] Furthermore, in the radiation image processing apparatus according to this disclosure, the processor acquires multiple radiation images, The thickness distribution of the subject is derived based on one of several radiographic images and the imaging conditions when all of the radiographic images were acquired. By performing weighted subtraction of a plurality of radiation images, a bone image in which the bone part of the subject is extracted is derived. A correction coefficient for correcting the pixel value of the bone image is obtained based on the body thickness distribution and the imaging conditions. A subtraction image may be derived by correcting the bone image with the correction coefficient.

[0012] The radiation image processing method according to the present disclosure displays a subtraction image in which a specific tissue in a subject is extracted, which is derived by performing weighted subtraction of a plurality of radiation images obtained by imaging the subject with radiation having different energy distributions and to which predetermined image processing has been applied. Receives an instruction to change the subtraction level corresponding to the weighting coefficient used when performing weighted subtraction. Determines whether or not the amount of change in the subtraction level exceeds a predetermined threshold value. When the determination is affirmative, a warning is issued.

[0013] The radiation image processing program according to the present disclosure includes a procedure for displaying a subtraction image in which a specific tissue in a subject is extracted, which is derived by performing weighted subtraction of a plurality of radiation images obtained by imaging the subject with radiation having different energy distributions and to which predetermined image processing has been applied, a procedure for receiving an instruction to change the subtraction level corresponding to the weighting coefficient used when performing weighted subtraction, a procedure for determining whether or not the amount of change in the subtraction level exceeds a predetermined threshold value, and a procedure for issuing a warning when the determination is affirmative, and causes the computer to execute them.

Advantages of the Invention

[0014] According to the present disclosure, the operator can calibrate the imaging apparatus at an appropriate timing by the warning.

Brief Description of the Drawings

[0015] [Figure 1] Schematic block diagram showing the configuration of a radiation image imaging system to which a radiation image processing apparatus according to a first embodiment of the present disclosure is applied [Figure 2] Diagram showing a schematic configuration of a radiation image processing apparatus according to a first embodiment [Figure 3] Diagram showing a functional configuration of a radiation image processing apparatus according to a first embodiment [Figure 4] Diagram for explaining linear approximation [Figure 5] Diagram showing a bone image [Figure 6] Diagram showing the relationship between the contrast of bone and soft tissues with respect to body thickness [Figure 7] Diagram showing a look-up table for obtaining a correction coefficient [Figure 8] Diagram showing a display screen [Figure 9] Diagram showing a display screen on which a warning is displayed [Figure 10] Flowchart showing the processes performed in a first embodiment [Figure 11] Diagram showing a functional configuration of a radiation image processing apparatus according to a second embodiment [Figure 12] Flowchart showing the processes performed in a second embodiment [Figure 13] Diagram showing a functional configuration of a radiation image processing apparatus according to a third embodiment [Figure 14] Flowchart showing the processes performed in a third embodiment

MODE FOR CARRYING OUT THE INVENTION

[0016] Hereinafter, embodiments of the present disclosure will be described with reference to the drawings. FIG. 1 is a schematic block diagram showing the configuration of a radiation image imaging system to which a radiation image processing apparatus according to a first embodiment of the present disclosure is applied. As shown in FIG. 1, the radiation image imaging system according to the present embodiment includes an imaging apparatus 1 and a radiation image processing apparatus 10 according to the first embodiment.

[0017] The imaging device 1 is an imaging device for performing energy subtraction using the so-called one-shot method, in which radiation such as X-rays emitted from the radiation source 3 and transmitted through the subject H is irradiated onto the first radiation detector 5 and the second radiation detector 6 with varying energies. During imaging, as shown in Figure 1, the first radiation detector 5, a radiation energy conversion filter 7 made of a copper plate or the like, and the second radiation detector 6 are arranged in order from the side closest to the radiation source 3, and the radiation source 3 is driven. The first and second radiation detectors 5 and 6 and the radiation energy conversion filter 7 are in close contact.

[0018] As a result, the first radiation detector 5 acquires a first radiation image G1 of subject H using low-energy radiation, including so-called soft rays. The second radiation detector 6 acquires a second radiation image G2 of subject H using high-energy radiation, from which soft rays have been removed. The first and second radiation images G1 and G2 are input to the radiation image processing device 10.

[0019] The first and second radiation detectors 5 and 6 are capable of repeatedly recording and reading radiation images. They may be so-called direct-type radiation detectors that generate an electric charge by directly receiving radiation, or they may be so-called indirect-type radiation detectors that first convert radiation into visible light and then convert that visible light into an electric charge signal. Furthermore, as a method for reading out the radiation image signal, it is desirable to use a so-called TFT (thin film transistor) readout method, in which the radiation image signal is read out by turning a TFT switch on and off, or a so-called optical readout method, in which the radiation image signal is read out by irradiating it with reading light. However, other methods may also be used.

[0020] In the imaging device 1, when acquiring the first radiation image G1 and the second radiation image G2, imaging conditions such as tube voltage, imaging dose, beam quality, SID (Source Image receptor Distance), which is the distance between the radiation source 3 and the surfaces of the first and second radiation detectors 5 and 6, SOD (Source Object Distance), which is the distance between the radiation source 3 and the surface of the subject H, and the presence or absence of a scatter removal grid are set.

[0021] SOD and SID are used to calculate body thickness distribution, as described later. For SOD, it is preferable to acquire it using, for example, a TOF (Time Of Flight) camera. For SID, it is preferable to acquire it using, for example, a potentiometer, ultrasonic rangefinder, or laser rangefinder.

[0022] The imaging conditions are pre-calibrated so that the set values ​​for imaging device 1 match the actual output. For example, the imaging dose and tube voltage are calibrated so that the output corresponds to the set values ​​set for imaging device 1. Similarly, the radiation detectors 5 and 6 are calibrated so that when the set dose of radiation is irradiated, the set signal value is output. The imaging conditions are set by the operator via input device 15.

[0023] Next, a radiation image processing apparatus according to the first embodiment will be described. First, with reference to Figure 2, the hardware configuration of the radiation image processing apparatus according to the first embodiment will be described. As shown in Figure 2, the radiation image processing apparatus 10 is a computer such as a workstation, server computer, or personal computer, and includes a CPU (Central Processing Unit) 11, non-volatile storage 13, and memory 16 as a temporary storage area. The radiation image processing apparatus 10 also includes a display 14 such as a liquid crystal display, input devices 15 such as a keyboard and mouse, and a network I / F (Interface) 17 connected to a network (not shown). The CPU 11, storage 13, display 14, input devices 15, memory 16, and network I / F 17 are connected to a bus 18. Note that the CPU 11 is an example of a processor in this disclosure.

[0024] The storage 13 is implemented using an HDD (Hard Disk Drive), an SSD (Solid State Drive), and flash memory, etc. The storage 13, as a storage medium, stores the radiation image processing program 12 installed in the radiation image processing device 10. The CPU 11 reads the radiation image processing program 12 from the storage 13, expands it into memory 16, and executes the expanded radiation image processing program 12.

[0025] The radiation image processing program 12 is stored in a memory device of a server computer connected to the network, or in network storage, in a state that allows external access, and is downloaded and installed on the computers comprising the radiation image processing device 10 upon request. Alternatively, it is recorded on a recording medium such as a DVD (Digital Versatile Disc) or CD-ROM (Compact Disc Read Only Memory) and distributed, and then installed from that recording medium on the computers comprising the radiation image processing device 10.

[0026] Next, the functional configuration of the radiation image processing apparatus according to the first embodiment will be described. Figure 3 is a diagram showing the functional configuration of the radiation image processing apparatus according to the first embodiment. As shown in Figure 3, the radiation image processing apparatus 10 includes an image acquisition unit 21, a scattered radiation removal unit 22, a bone image extraction unit 23, a correction unit 24, a display control unit 25, and a warning unit 26. The CPU 11 then executes the radiation image processing program 12 and functions as the image acquisition unit 21, the scattered radiation removal unit 22, the bone image extraction unit 23, the correction unit 24, the display control unit 25, and the warning unit 26.

[0027] The image acquisition unit 21 causes the imaging device 1 to perform energy subtraction imaging of the subject H, thereby acquiring the first radiation image G1 and the second radiation image G2 of the subject H from the first and second radiation detectors 5 and 6. The imaging conditions are set as described above when acquiring the first radiation image G1 and the second radiation image G2.

[0028] The scattered radiation removal unit 22 removes scattered radiation components from the first radiation image G1 and the second radiation image G2 acquired by the image acquisition unit 21. The process of removing scattered radiation components is an example of the image processing predetermined in this disclosure. The removal of scattered radiation components will be described below. Any method can be used to remove scattered radiation components, such as the method described in Japanese Patent Application Publication No. 2015-043959. The scattered radiation removal process for the first radiation image G1 using the method described in Japanese Patent Application Publication No. 2015-043959 will be described below. In the following description, the first and second radiation images from which scattered radiation components have been removed will also be referred to as G1 and G2, respectively.

[0029] First, the scatter removal unit 22 acquires a virtual model of the subject H having an initial thickness distribution Ts(x,y). The virtual model is data that virtually represents the subject H, in which the thickness according to the initial thickness distribution Ts(x,y) is associated with the coordinate position of each pixel in the first radiation image G1. The virtual model of the subject H having an initial thickness distribution Ts(x,y) is assumed to be stored in storage 13 in advance, but it may also be acquired from an external server where the virtual model is stored. The derivation of the initial thickness distribution Ts(x,y) will be explained below.

[0030] First, when the radiation source 3 is driven to irradiate the radiation detector 5 with radiation in the absence of a subject H, the dose I0(x,y) of the radiation emitted from the radiation source 3 that reaches the radiation detector 5 is expressed by the following equation (1). In equation (1), mAs included in the imaging conditions is the dose, and kV is the tube voltage. Here, F is a linear or nonlinear function that represents the amount of radiation that reaches the radiation detector 5 when a reference dose (e.g., 1 mAs) is irradiated to the radiation detector 5 at a reference SID (e.g., 100 cm) in the absence of a subject H. F changes depending on the tube voltage. Also, since the received dose I0 is derived for each pixel of the radiation image G0 acquired by the radiation detector 5, (x,y) represents the pixel position of each pixel. I0(x,y) = mAs × F(kV) / SID 2 (1)

[0031] Furthermore, if we denote the initial thickness distribution as Ts, the attenuation coefficient of the subject H with the initial thickness distribution Ts as μ(Ts), and the scatter-to-primary ratio (STPR(Ts)), which is the ratio of the scattered dose to the primary dose in the radiation after passing through the subject H with the initial thickness distribution Ts, without considering the spread of scattered radiation, then the dose I1 after passing through the subject H is expressed by the following equation (2). Note that in equation (2), the initial thickness distribution Ts, the reached dose I0, and the dose I1 are derived for each pixel of the radiation image G0, but (x,y) is omitted. Also, STPR is a nonlinear function that depends not only on the thickness but also on the tube voltage (kV), but the kV notation is omitted in equation (2). STPR(Ts) is determined in advance experimentally or by simulation. I1=I0×exp{-μ(Ts)×Ts}×{1+STPR(Ts)} (2)

[0032] In equation (2), dose I1 is the pixel value at each pixel of the radiation image G1, and the received dose I0 is derived from equation (1) above. On the other hand, F may be a nonlinear function, and since STPR is a nonlinear function, equation (2) cannot be solved algebraically for Ts. For this reason, the scatter removal unit 22 defines an error function E1 shown in equation (3) or equation (3-1) below. Then, it derives an initial thickness distribution Ts that minimizes the error function E1 or makes the error function E1 less than a predetermined threshold Th1. In this case, the scatter removal unit 22 derives the initial thickness distribution Ts using optimization algorithms such as the steepest descent method and the conjugate gradient method. E1=[I1-I0×exp{-μ(Ts)×Ts}×{1+STPR(Ts)}] 2 (3) E1=|I1-I0×exp{-μ(Ts)×Ts}×{1+STPR(Ts)}| (3-1)

[0033] The function F in equation (1) above can be derived as follows. First, at a reference SID (e.g., 100 cm) for each tube voltage used during imaging, radiation is irradiated to the radiation detector 5 with a reference dose (e.g., 1 mAs), and the signal value output from the radiation detector 5 is derived. Then, the signal value is considered as the acquired dose I0, and the tube voltage (kV) is plotted on the x-axis and the acquired dose I0 (QL / mAs) on the y-axis, as shown in Figure 4. Since the plot shown in Figure 4 can be linearly approximated, the plot is linearly approximated. This allows us to find the relationship I0 = a·x + b, where x is the tube voltage. Having found such a relationship in advance, the acquired dose I0 can be derived from equation (1) using equation (4) below, using the tube voltage x1 (kV), tube current x2 (mAs), and SID x3 (cm) that are actually used. If the plot changes nonlinearly, the least squares method can be used to derive the nonlinear function. I0=x2(mAs)×(a x1+b)×(100 / c) 2 (4)

[0034] The scattered radiation removal unit 22 derives an estimated primary radiation image Ip(x,y) based on a virtual model, which is an estimated primary radiation image obtained by imaging the virtual model, and an estimated scattered radiation image Is(x,y) based on a scattered radiation image obtained by imaging the virtual model, as shown in equations (5) and (6) below. Furthermore, as shown in equation (7) below, the scattered radiation removal unit 22 derives an estimated image Im(x,y) by combining the estimated primary radiation image Ip(x,y) and the estimated scattered radiation image Is(x,y) as an estimated image obtained by estimating the first radiation image G1 obtained by imaging the subject H. Ip(x,y) = Io(x,y)×exp(-μSoftT(x,y))×T(x,y)) (5) Is(x,y) = Io(x,y)×STPR(T(x,y))*PSF(T(x,y)) (6) Im(x,y) = Is(x,y)+Ip(x,y) (7)

[0035] Here, (x,y) is the coordinate of the pixel position of the first radiation image G1, Io(x,y) is the pixel value of the first radiation image G1 at pixel position (x,y), Ip(x,y) is the primary component at pixel position (x,y), and Is(x,y) is the scattered component at pixel position (x,y). Note that when deriving the first estimated image Im(x,y), the initial thickness distribution Ts(x,y) is used as the thickness distribution T(x,y) in equations (5) and (6).

[0036] Furthermore, μSoft(T(x,y)) in equation (5) is an attenuation coefficient corresponding to the body thickness distribution (x,y) of the soft tissue of the human body at the pixel position (x,y). μSoft(T(x,y)) can be determined in advance experimentally or by simulation and stored in storage 13. In this embodiment, the attenuation coefficient μSoft(T(x,y)) is the attenuation coefficient of the standard soft tissue described later.

[0037] Furthermore, PSF(T(x,y)) in equation (6) is a point spread function that represents the distribution of scattered rays spreading from a single pixel according to the body thickness distribution T(x,y), and is defined according to the energy characteristics of the radiation. Also, * is an operator that indicates a convolution operation. PSF also changes depending on the distribution of the irradiation field in the imaging device 1, the composition distribution of the subject H, the irradiation dose during imaging, the tube voltage, the imaging distance, and the characteristics of the radiation detectors 5 and 6. For this reason, PSF should be experimentally determined in advance for each energy characteristic of the radiation used by the imaging device 1, according to the irradiation field information, subject information, and imaging conditions.

[0038] Next, the scattered radiation removal unit 22 modifies the initial body thickness distribution Ts(x,y) of the virtual model so that the difference between the estimated image Im and the first radiation image G1 is small. The scattered radiation removal unit 22 updates the body thickness distribution T(x,y), scattered radiation component Is(x,y), and primary radiation component Ip(x,y) by repeatedly deriving the body thickness distribution T(x,y), scattered radiation component Is(x,y), and primary radiation component Ip(x,y) until the difference between the estimated image Im and the first radiation image G1 satisfies a predetermined termination condition. When the termination condition is met, the scattered radiation removal unit 22 subtracts the scattered radiation component Is(x,y) derived by equation (7) from the first radiation image G1. This removes the scattered radiation component contained in the first radiation image G1. The body thickness distribution T(x,y) derived when the termination condition is met is used when deriving bone density, which will be described later.

[0039] Meanwhile, the scattered radiation removal unit 22 performs the scattered radiation removal process on the second radiation image G2 in the same manner as on the first radiation image G1. In the following description, the reference numerals G1 and G2 will also be used for the first and second radiation images from which the scattered radiation component has been removed.

[0040] The bone image extraction unit 23 generates a bone image Gb representing the bone region of subject H from the first radiation image G1 and the second radiation image G2, from which the scattered radiation component has been removed by the scattered radiation removal unit 22. Figure 5 shows an example of a bone image Gb generated by the bone image extraction unit 23. The bone image extraction unit 23 generates a bone image Gb in which only the bone region of subject H contained in the first radiation image G1 and the second radiation image G2 is extracted by performing weighted subtraction between corresponding pixels on the first radiation image G1 and the second radiation image G2, as shown in equation (8) below. In equation (8) below, α1 is a weighting coefficient, which is derived based on attenuation coefficients corresponding to the radiation energy of the soft tissue and bone region of subject H. Also, x and y are the coordinates of each pixel in the bone image Gb. Gb(x,y)=G1(x,y)-α1×G2(x,y) (8)

[0041] The correction unit 24 corrects the bone image Gb generated by the bone image derivation unit 23 to derive a bone density image representing the bone density in the bone region of the subject H. The process of correcting the bone image Gb is an example of the image processing in this disclosure. The correction unit 24 derives a bone density image B by deriving the bone density B(x,y) for each pixel (x,y) of the bone image Gb. Note that the bone density B(x,y) may be derived for all bones included in the bone image Gb, or it may be derived only for predetermined bones.

[0042] Specifically, the correction unit 24 converts each pixel value Gb(x,y) of the bone region in the bone image Gb to the pixel value of the bone image acquired under standard imaging conditions, and considers this value as bone density, thereby deriving the bone density B(x,y) corresponding to each pixel of the bone image Gb. More specifically, the correction unit 24 derives the bone density B(x,y) for each pixel by correcting each pixel value Gb(x,y) of the bone image Gb using a correction coefficient obtained from a lookup table described later.

[0043] Here, the higher the tube voltage at radiation source 3 and the higher the energy of the radiation emitted from radiation source 3, the smaller the contrast (i.e., the difference in pixel values) between soft tissue and bone in the radiation image. Also, during the process of radiation passing through subject H, the low-energy component of the radiation is absorbed by subject H, resulting in beam hardening, which increases the energy of the radiation. The increase in radiation energy due to beam hardening increases with the thickness of subject H.

[0044] Figure 6 shows the relationship between the contrast between bone and soft tissue and the thickness of subject H. Figure 6 shows the relationship between the contrast between bone and soft tissue and the thickness of subject H at three tube voltages: 80kV, 90kV, and 100kV. As shown in Figure 6, the higher the tube voltage, the lower the contrast. Furthermore, beyond a certain thickness of subject H, the greater the thickness, the lower the contrast. Note that the larger the pixel value Gb(x,y) in the bone region of the bone image Gb, the greater the contrast between bone and soft tissue. Therefore, the relationship shown in Figure 6 shifts towards higher contrast as the pixel value Gb(x,y) in the bone region of the bone image Gb increases.

[0045] In this embodiment, a lookup table for obtaining correction coefficients to compensate for the difference in contrast in the bone image Gb according to the tube voltage during acquisition, and the decrease in contrast due to the effect of beam hardening, is stored in storage 13. The correction coefficient is a coefficient for correcting each pixel value Gb(x,y) of the bone image Gb.

[0046] Figure 7 shows an example of a lookup table stored in storage 13. In Figure 7, a lookup table LUT1 is shown as an example, with the reference imaging conditions set to a tube voltage of 90kV. As shown in Figure 7, in lookup table LUT1, a larger correction coefficient is set as the tube voltage increases and the thickness of the subject increases. In the example shown in Figure 7, since the reference imaging conditions are a tube voltage of 90kV, the correction coefficient is 1 when the tube voltage is 90kV and the body thickness is 0. Note that although lookup table LUT1 is shown in two dimensions in Figure 7, the correction coefficient differs depending on the pixel value of the bone region. Therefore, lookup table LUT1 is actually a three-dimensional table with an axis representing the pixel value of the bone region.

[0047] Here, the correction coefficient is derived by taking radiographic images using a phantom that simulates the human body, containing materials corresponding to the soft tissues and bone tissues of the human body. For example, acrylic or urethane can be used as the material corresponding to the soft tissues. For example, hydroxyapatite can be used as the material corresponding to the bone tissues of the human body. The radiation attenuation coefficient and fat content of the material corresponding to the soft tissues in the phantom are predetermined values ​​depending on the material. As mentioned above, the greater the thickness of the subject's body, the smaller the contrast between bone and soft tissue. As a result, even if the bone density is the same, the bone density derived from the radiographic image will be smaller as the body thickness increases, given the same body thickness.

[0048] Therefore, by imaging phantoms of various thicknesses, pixel values ​​for the bone region corresponding to the body thickness are derived from the acquired radiographic images, and a correction coefficient corresponding to the body thickness is derived so that the derived pixel values ​​correspond to the same bone density. Furthermore, the correction coefficient is derived according to various tube voltages. This allows for the derivation of the lookup table LUT1 shown in Figure 7.

[0049] The correction unit 24 obtains a pixel-by-pixel correction coefficient C0(x,y) from the lookup table LUT1, which corresponds to the imaging conditions, including the body thickness distribution T(x,y) of the subject H and the tube voltage setting value stored in the storage 13. Then, as shown in equation (9) below, the correction unit 24 derives the bone density B(x,y) for each pixel of the bone region in the bone region image Gb by multiplying it by the correction coefficient C0(x,y). This derives a bone density image B in which the bone density B(x,y) is the pixel value. The bone density B(x,y) thus derived represents the pixel value of the bone region in the bone region included in the radiographic image obtained by imaging the subject H with a tube voltage of 90kV, which is the standard imaging condition, and from which the effects of beam hardening have been removed. In this embodiment, the unit of bone density is g / cm 2 It is assumed that this is the case. Bone density image B is an example of a subtraction image in this disclosure. B(x,y) = C0(x,y) × Gb(x,y) (9)

[0050] The display control unit 25 displays the bone density image B on the display 14. Figure 8 is a diagram showing the display screen of the bone density image B. As shown in Figure 8, the bone density image B is displayed on the display screen 30. The display screen 30 also displays a subtraction level adjustment interface 31. The subtraction level is a value that reflects the weighting coefficient α1 in the above equation (8), and is set so that the subtraction level with respect to the weighting coefficient α1 is 1.0. The adjustment interface 31 has a "-" button 31A and a "+" button 31B. When the operator operates the "-" button 31A, the subtraction level decreases, and as a result, the weighting coefficient α1 also decreases. When the operator operates the "+" button 31B, the subtraction level increases, and as a result, the weighting coefficient α1 also increases. In this embodiment, the subtraction level changes in units of 0.1 by operating the "-" button 31A and the "+" button 31B, and the weighting coefficient α1 changes accordingly. The initial value of the subtraction level is 1.0. Figure 8 shows the state where the subtraction level has been changed from 1.0 to 1.2.

[0051] When the operator changes the subtraction level value by operating the adjustment interface 31, the bone image output unit 23 changes the weighting coefficient α1 in equation (8). When the weighting coefficient α1 in equation (8) changes, the degree of soft tissue removal in the bone image Gb changes. Therefore, when the operator changes the subtraction level value by operating the adjustment interface 31, the degree of soft tissue removal in the bone density image B changes.

[0052] Here, the sensitivity of the radiation tube and radiation detectors 5 and 6 used in the radiation source 3 of the imaging device 1 deteriorates over time. When the tube deteriorates, the imaging conditions set in the imaging device 1 deviate from the set values. When the imaging conditions deviate from the set values, the tube voltage (kV) and dose (mAs) in equation (1) deviate from the set values, causing the received dose I0 to shift, and as a result, the initial body thickness distribution Ts also shifts. Therefore, it becomes impossible to accurately derive the body thickness distribution of the subject H. In addition, when the sensitivity of the radiation detectors 5 and 6 deteriorates, the accuracy of the signal values ​​output from the radiation detectors 5 and 6 decreases. When the accuracy of the body thickness distribution and the accuracy of the signal values ​​decrease, the accuracy of the scattered radiation component derived in equation (6) above decreases, and as a result, it becomes impossible to accurately remove the scattered radiation component from the radiation images G1 and G2.

[0053] Furthermore, if the tube voltage deviates from the set value, and the accuracy of the body thickness distribution decreases, the accuracy of the correction coefficient C0(x,y) in equation (9) also decreases, making it impossible to accurately correct the bone image Gb.

[0054] When the accuracy of scatter radiation component removal and the correction accuracy of bone image Gb decrease, the image quality of the displayed bone density image B deteriorates. This deterioration in image quality of bone density image B manifests as unnatural image quality due to the appearance of soft tissue or excessive removal of soft tissue.

[0055] In this embodiment, the operator can change how soft tissue appears in bone density image B by viewing the bone density image B displayed on the display screen 30 and changing the subtraction level. By changing how soft tissue appears in bone density image B, it is possible to compensate for the deterioration of image quality of bone density image B caused by a decrease in the accuracy of scatter radiation component removal and a decrease in the accuracy of bone image correction. Therefore, by changing the subtraction level, the operator can display bone density image B with the desired image quality on the display screen 30.

[0056] Additionally, an exit button 32 is displayed on the display screen 30. After changing the subtraction level, the operator can save the bone density image B with the modified image quality to the storage 13 by operating the exit button 32.

[0057] On the other hand, as the tube and radiation detectors 5 and 6 deteriorate, the deviation from the set values ​​of the imaging conditions increases, resulting in a greater degree of deterioration in the image quality of bone density image B, and thus a larger change in the subtraction level. When the change in the subtraction level becomes large enough, it is necessary to calibrate the imaging device 1.

[0058] In this embodiment, the warning unit 26 determines whether the amount of change in the subtraction level exceeds a predetermined threshold. Specifically, it determines whether the amount of change in the subtraction level exceeds ±0.3, that is, whether the rate of change in the subtraction level exceeds 30%. In this case, the warning unit 26 determines whether the value displayed on the adjustment interface 31 becomes less than 0.7 or greater than 1.3. If the determination is affirmative, the warning unit 26 issues a warning. Figure 9 shows the display screen with the warning displayed. As shown in Figure 9, the value on the adjustment interface 31 is 1.4, which exceeds 1.3. Therefore, the warning unit 26 displays the warning 33 "Please perform calibration." on the display screen 30. The warning may be given by voice in addition to the display, or instead of the display. Furthermore, the threshold for determining the amount of change is not limited to the above.

[0059] Next, the process performed in the first embodiment will be described. Figure 10 is a flowchart showing the process performed in the first embodiment. First, the image acquisition unit 21 causes the imaging device 1 to perform energy subtraction imaging of the subject H, thereby acquiring first and second radiation images G1 and G2 (radiation image acquisition; step ST1). Next, the scattered radiation removal unit 22 derives the body thickness distribution of the subject H (step ST2) and removes the scattered radiation component from the first and second radiation images G1 and G2 (step ST3). Steps ST2 and ST3 are performed in parallel. Then, the bone image extraction unit 23 derives a bone image Gb from which the bone portion of the subject H has been extracted from the first and second radiation images G1 and G2 from which the scattered radiation component has been removed (step ST4).

[0060] Next, the correction unit 24 corrects the bone image Gb to derive the bone density image B (step ST5), and the display control unit 25 displays the bone density image B (step ST6). Next, the bone image deriving unit 23 determines whether or not the subtraction level has been changed (step ST7). If step ST7 is affirmative, the warning unit 26 determines whether or not the amount of change in the subtraction level exceeds a predetermined threshold (change amount determination: step ST8). If step ST8 is negative, the process returns to step ST4, and the bone image Gb is derived using a weighting coefficient α1 corresponding to the changed subtraction level, and the bone density image B is derived.

[0061] If step ST8 is affirmed, the warning unit 26 issues a warning (step ST9) and terminates the process. On the other hand, if step ST7 is denied, it is determined whether or not a termination instruction has been given (step ST10), and if step ST10 is denied, the process returns to step ST7. If step ST10 is affirmed, the process terminates.

[0062] Thus, in this embodiment, a warning is issued when the amount of change in the subtraction level exceeds a predetermined threshold. This allows the operator to calibrate the imaging device 1 at the appropriate time. Furthermore, by calibrating the imaging device 1 at the appropriate time, degradation of the acquired radiation images can be prevented.

[0063] Next, a second embodiment of the present disclosure will be described. Figure 11 is a diagram showing the functional configuration of a radiation image processing apparatus according to the second embodiment. In Figure 11, the same reference numerals are used for components identical to those in Figure 3, and detailed explanations are omitted. The second embodiment differs from the first embodiment in that only scatter removal processing is performed, and no processing is performed to correct the bone image Gb. For this reason, the radiation image processing apparatus 10A according to the second embodiment does not include a correction unit 24. In the second embodiment, the bone image Gb corresponds to the subtraction image of the present disclosure.

[0064] Next, the process performed in the second embodiment will be described. Figure 12 is a flowchart showing the process performed in the second embodiment. First, the image acquisition unit 21 causes the imaging device 1 to perform energy subtraction imaging of the subject H, thereby acquiring the first and second radiation images G1 and G2 (radiation image acquisition; step ST11). Next, the scattered radiation removal unit 22 derives the body thickness distribution of the subject H (step ST12) and removes the scattered radiation component from the first and second radiation images G1 and G2 (step ST13). Note that the processes in steps ST12 and ST13 are performed in parallel. Then, the bone image extraction unit 23 derives a bone image Gb from which the bone portion of the subject H has been extracted from the first and second radiation images G1 and G2 from which the scattered radiation component has been removed (step ST14).

[0065] Next, the display control unit 25 displays the bone image Gb (step ST15). Subsequently, the bone image derivation unit 23 determines whether the subtraction level has been changed (step ST16). If step ST16 is affirmative, the warning unit 26 determines whether the amount of change in the subtraction level exceeds a predetermined threshold (change amount determination: step ST17). If step ST17 is negative, the process returns to step ST14, and the bone image Gb is derived using a weighting coefficient α1 corresponding to the changed subtraction level.

[0066] If step ST17 is affirmed, the warning unit 26 issues a warning (step ST18) and terminates the process. On the other hand, if step ST16 is denied, it is determined whether or not a termination instruction has been given (step ST19), and if step ST19 is denied, the process returns to step ST16. If step ST19 is affirmed, the process terminates.

[0067] Next, a third embodiment of the present disclosure will be described. Figure 13 is a diagram showing the functional configuration of the radiation image processing apparatus according to the third embodiment. In Figure 13, the same reference numerals are used for components identical to those in Figure 3, and detailed explanations are omitted. The third embodiment differs from the first embodiment in that it does not perform scattered radiation removal processing, but only performs processing to correct the bone image Gb. For this reason, the radiation image processing apparatus 10B according to the third embodiment does not have a scattered radiation removal unit 22. On the other hand, as described above, the thickness distribution of the subject H is necessary to derive the correction coefficient. In the third embodiment, the correction unit 24 derives the thickness distribution of the subject H. Alternatively, a separate thickness derivation unit for deriving the thickness distribution of the subject H may be provided. Furthermore, in the third embodiment, a scattered radiation removal unit 22 may be provided, and the scattered radiation removal unit 22 may derive the thickness distribution. In the third embodiment, the bone density image B corresponds to the subtraction image of the present disclosure.

[0068] Next, the process performed in the third embodiment will be described. Figure 14 is a flowchart showing the process performed in the third embodiment. First, the image acquisition unit 21 causes the imaging device 1 to perform energy subtraction imaging of the subject H, thereby acquiring the first and second radiation images G1 and G2 (radiation image acquisition; step ST21). Next, the scattered radiation removal unit 22 derives the body thickness distribution T(x,y) of the subject H (step ST22). Subsequently, the bone image deriving unit 23 derives a bone image Gb from the first and second radiation images G1 and G2, in which the bone portion of the subject H is extracted (step ST23).

[0069] Next, the correction unit 24 corrects the bone image Gb to derive the bone density image B (step ST24), and the display control unit 25 displays the bone density image B (step ST25). Next, the bone image deriving unit 23 determines whether the subtraction level has been changed (step ST26). If step ST26 is affirmative, the warning unit 26 determines whether the amount of change in the subtraction level exceeds a predetermined threshold (change amount determination: step ST27). If step ST27 is negative, the process returns to step ST23, and the bone image is derived using a weighting coefficient α1 corresponding to the changed subtraction level, and the bone density image B is derived.

[0070] If step ST27 is affirmed, the warning unit 26 issues a warning (step ST28) and terminates the process. On the other hand, if step ST26 is denied, it is determined whether or not a termination instruction has been given (step ST29), and if step ST29 is denied, the process returns to step ST26. If step ST29 is affirmed, the process terminates.

[0071] In the above embodiments, the display screen 30 shows either a bone density image B or a bone image Gb, but it is not limited to this. The soft tissue image Gs of the subject H may be derived using the following formula (10), and the soft tissue image Gs and the bone image Gb may be displayed in a switchable manner. Furthermore, in addition to the soft tissue image Gs and the bone image Gb, radiographic images G1 and G2 may be displayed in a switchable manner. In this case, the displayed radiographic images G1 and G2 may be those with scattered radiation removed or not. Gs(x, y)=G1(x, y)-α2×G2(x, y) (10)

[0072] Furthermore, while the above embodiments derive bone density images B for the chest region of subject H, the method is not limited to this. For example, bone density images B may be derived for any bone region of subject H, such as the femur, vertebrae, calcaneus, and metacarpals.

[0073] Furthermore, in each of the above embodiments, the first and second radiographic images G1 and G2 are acquired by the one-shot method when performing energy subtraction processing to derive bone density, but the method is not limited to this. The first and second radiographic images G1 and G2 may also be acquired by the so-called two-shot method, in which imaging is performed twice using only one radiation detector. In the case of the two-shot method, the position of subject H included in the first radiographic image G1 and the second radiographic image G2 may shift due to the movement of subject H. For this reason, it is preferable to perform positional alignment of the subject in the first radiographic image G1 and the second radiographic image G2 before performing the processing of this embodiment.

[0074] Furthermore, in the above embodiment, bone density is derived using radiation images acquired in a system that photographs subject H using first and second radiation detectors 5 and 6. However, the technology of this disclosure can also be applied when acquiring the first and second radiation images G1 and G2 using a accumulative phosphor sheet instead of radiation detectors. In this case, two accumulative phosphor sheets are stacked and radiation transmitted through subject H is irradiated onto them to accumulate and record the radiation image information of subject H on each accumulative phosphor sheet. The first and second radiation images G1 and G2 can then be acquired by photoelectrically reading the radiation image information from each accumulative phosphor sheet. Note that the two-shot method may also be used when acquiring the first and second radiation images G1 and G2 using a accumulative phosphor sheet.

[0075] Furthermore, the radiation in the above embodiment is not particularly limited, and in addition to X-rays, alpha rays or gamma rays can be used.

[0076] Furthermore, in the above embodiment, the hardware structure of the Processing Unit, which executes various processes such as the image acquisition unit 21, the scattered radiation removal unit 22, the bone image extraction unit 23, the correction unit 24, the display control unit 25, and the warning unit 26, can be any of the following types of processors. As mentioned above, these types of processors include a CPU, which is a general-purpose processor that executes software (programs) and functions as various processing units, as well as a Programmable Logic Device (PLD), which is a processor whose circuit configuration can be changed after manufacturing, such as an FPGA (Field Programmable Gate Array), and a dedicated electrical circuit, which is a processor with a circuit configuration specifically designed to execute a particular process, such as an ASIC (Application Specific Integrated Circuit).

[0077] A single processing unit may consist of one of these various processors, or it may consist of a combination of two or more processors of the same or different types (for example, a combination of multiple FPGAs or a combination of a CPU and an FPGA). Alternatively, multiple processing units may be composed of a single processor.

[0078] Examples of configuring multiple processing units with a single processor include, firstly, a configuration where one or more CPUs and software combine to form a single processor, as exemplified by client and server computers, and this processor functions as multiple processing units. Secondly, a configuration using a processor that realizes the functions of the entire system, including multiple processing units, on a single IC (Integrated Circuit) chip, as exemplified by System-on-a-Chip (SoC). Thus, various processing units are configured, in terms of hardware structure, using one or more of the above-mentioned processors.

[0079] Furthermore, the hardware structure of these various processors can more specifically utilize electrical circuits (Circuitry) that combine circuit elements such as semiconductor elements. [Explanation of Symbols]

[0080] 1. Imaging device 3 Radiation source 5, 6 Radiation detectors 7. Radiation energy conversion filter 10, 10A, 10B Radiation imaging processing device 11 CPU 12. Radiation Image Processing Program 13 Storage 14 displays 15 Input Devices 16 memory 17 Network Interface 18 bus 21 Image acquisition unit 22 Scattered radiation removal section 23 Bone Image Output Unit 24 Correction section 25 Display Control Unit 26 Warning section 30 display screen 31 Adjustment Interface 31A "-" button 31B "+" button 32. Exit button 33 Warning Gb bone image B Bone density image H Subject

Claims

1. Equipped with at least one processor, The aforementioned processor, A subtraction image is displayed, which is derived by weighting and subtracting multiple radiation images obtained by photographing the subject with radiation having different energy distributions, and then subjected to predetermined image processing, thereby extracting the bone or soft tissue within the subject. The system accepts instructions to change the subtraction level corresponding to the weight coefficient used when performing the aforementioned weighted subtraction. Determine whether the amount of change from the initial value of the subtraction level exceeds a predetermined threshold. A radiation image processing device that issues a warning if the aforementioned determination is confirmed.

2. The processor acquires the plurality of radiation images, The thickness distribution of the subject is derived based on any of the plurality of radiographic images and the shooting conditions when the plurality of radiographic images were acquired. Based on the body thickness distribution and the imaging conditions, the scattered radiation component of the radiation scattered by the subject is removed from the plurality of radiation images. By weighting and subtracting multiple radiation images from which the scattered radiation components have been removed, a bone image from which the bone portion of the subject is extracted is derived. A correction coefficient is obtained to correct the pixel values ​​of the bone image based on the body thickness distribution and the shooting conditions. The radiation image processing apparatus according to claim 1, wherein the subtraction image is derived by correcting the bone image with the correction coefficient.

3. The processor acquires the plurality of radiation images, The thickness distribution of the subject is derived based on any of the plurality of radiographic images and the shooting conditions when the plurality of radiographic images were acquired. Based on the body thickness distribution and the imaging conditions, the scattered radiation component of the radiation scattered by the subject is removed from the plurality of radiation images. The radiation image processing apparatus according to claim 1, wherein the subtraction image is derived by weighting and subtracting a plurality of radiation images from which the scattered radiation components have been removed.

4. The processor acquires the plurality of radiation images, The thickness distribution of the subject is derived based on any of the plurality of radiographic images and the shooting conditions when the plurality of radiographic images were acquired. By weighting and subtracting the aforementioned multiple radiographic images, a bone image is derived in which the bone portion of the subject is extracted. A correction coefficient is obtained to correct the pixel values ​​of the bone image based on the body thickness distribution and the shooting conditions. The radiation image processing apparatus according to claim 1, wherein the subtraction image is derived by correcting the bone image with the correction coefficient.

5. A subtraction image is displayed, which is derived by weighting and subtracting multiple radiation images obtained by photographing the subject with radiation having different energy distributions, and then subjected to predetermined image processing, thereby extracting the bone or soft tissue within the subject. The system accepts instructions to change the subtraction level corresponding to the weight coefficient used when performing the aforementioned weighted subtraction. Determine whether the amount of change from the initial value of the subtraction level exceeds a predetermined threshold. A radiation image processing method that issues a warning if the aforementioned determination is confirmed.

6. A procedure for displaying a subtraction image in which the bone or soft tissue within the subject is extracted, which is derived by weighting and subtracting multiple radiation images obtained by photographing the subject with radiation having different energy distributions, and which has undergone predetermined image processing; A procedure for receiving instructions to change the subtraction level corresponding to the weight coefficient used when performing the aforementioned weighted subtraction, A procedure for determining whether the amount of change from the initial value of the subtraction level exceeds a predetermined threshold, A radiation image processing program that causes a computer to execute a procedure to issue a warning if the aforementioned determination is confirmed.

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