Medical image processing apparatus and medical image processing method

Abstract

To compare medical images captured under different imaging conditions in a CT apparatus at the same energy level.SOLUTION: A medical image processing apparatus includes an acquisition unit, a determination unit, and a generation unit. The acquisition unit acquires imaging conditions including a tube voltage of an X-ray tube used when capturing a past CT image of a subject. The determination unit determines, based on the acquired imaging condition, a range of energy bins to be used for generating a photon counting CT image corresponding to a past CT image, among a plurality of energy bins set for counting, for each energy band, X-ray photons detected by new imaging performed on a subject using a photon counting CT apparatus. The generation unit generates a photon counting CT image using an X-ray photon count in the determined range of energy bins.SELECTED DRAWING: Figure 1

Description

Technical Field

[0001] The embodiments disclosed in this specification and the drawings relate to a medical image processing apparatus and a medical image processing method.

Background Art

[0002] In follow-up examinations after treatment, regular health check-ups, etc., it may be necessary to compare a previously taken medical image with a newly taken medical image to observe changes in the state of the subject. When performing such a comparison of medical images, it is important to align the imaging conditions of both to eliminate differences in the appearance of the images due to differences in imaging conditions and make it easier to confirm changes in the state of the subject.

[0003] By the way, in recent years, in the field of medical imaging diagnostic devices, in addition to conventional CT (Computed Tomography) devices using detectors composed of scintillators and photodiodes, new photon counting CT devices (Photon Counting Computed Tomography; PCCT) using photon counting detectors, which are semiconductor detectors, have come to be used. In photon counting CT devices, various images (integral images (conventional images), Mono, k-edge, etc.) can be created.

Prior Art Documents

Patent Documents

[0004]

Patent Document 1

Summary of the Invention

Problems to be Solved by the Invention

[0005] When there is a need to compare a medical image captured by a conventional CT device with a newly captured medical image by a photon counting CT device, it is not easy to compare the two at the same energy level. In a conventional CT device, the tube voltage is arbitrarily changed according to the subject for imaging. However, in imaging using a photon counting CT device, it is common to set the tube voltage of the X-ray tube to 120 kVp or 140 kVp for imaging. Therefore, if a past medical image was captured with the tube voltage set to 80 kVp by a conventional CT device, even if a CT image (integrated image) is created using the sum data of the count values in a plurality of energy bands (energy bins) collected by the photon counting CT device, there will be a difference in the appearance of the images due to the difference in tube voltage, and the two images cannot be simply compared. Also, even when comparing medical images captured by a photon counting CT device, if they were captured under imaging conditions with different energy bin settings, etc., it is not easy to compare the two at the same energy level.

[0006] The problem to be solved by the embodiments disclosed in this specification and the drawings is to enable comparison of medical images captured under different imaging conditions in a CT device at the same energy level. However, the problems to be solved by the embodiments disclosed in this specification and the drawings are not limited to the above problems. The problems corresponding to the effects of each configuration shown in the embodiments described later can also be regarded as other problems.

Means for Solving the Problem

[0007] The medical image processing apparatus according to the embodiment includes an acquisition unit, a determination unit, and a generation unit. The acquisition unit acquires imaging conditions including the tube voltage of the X-ray tube used during past CT imaging of a subject. The determination unit determines, based on the acquired imaging conditions, a range of energy bins to be used for generating a photon counting CT image corresponding to the past CT image from among a plurality of energy bins set for counting X-ray photons detected by a new imaging using a photon counting CT apparatus for the subject for each energy band. The generation unit generates a photon counting CT image using the count values of X-ray photons in the determined range of energy bins.

Brief Description of the Drawings

[0008]

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Best Mode for Carrying Out the Invention

[0009] Hereinafter, a medical image processing apparatus and a medical image processing method according to an embodiment will be described with reference to the drawings.

[0010] <First Embodiment> [Configuration of Photon Counting CT Apparatus] FIG. 1 is a diagram showing an example of a photon counting CT apparatus 1 according to the first embodiment. The photon counting CT apparatus 1 can generate image data that discriminates substances of an inspection object through which X-rays have passed, using a direct detector such as a semiconductor detector having excellent energy resolution. Further, the photon counting CT apparatus 1 can generate image data similar to a conventional CT image using the total data of the count values in a plurality of energy bands (energy bins).

[0011] The photon counting CT apparatus 1 includes, for example, a gantry device 10, a couch device 30, and a console device 40. In FIG. 1, for convenience of explanation, both a view of the gantry device 10 as seen from the Z-axis direction and a view as seen from the X-axis direction are shown, but actually, there is only one gantry device 10. In the present embodiment, the rotation axis of the rotating frame 17 or the longitudinal direction of the top plate 33 of the couch device 30 in a non-tilted state is defined as the Z-axis direction, an axis orthogonal to the Z-axis direction and horizontal with respect to the floor surface is defined as the X-axis direction, and a direction orthogonal to the Z-axis direction and perpendicular to the floor surface is defined as the Y-axis direction.

[0012] <Gantry Device 10> The gantry device 10 includes, for example, an X-ray tube 11, a wedge 12, a collimator 13, an X-ray high voltage device 14, an X-ray detector 15, a data acquisition system (hereinafter, DAS: Data Acquisition System) 16, a rotating frame 17, and a control device 18. The X-ray detector 15 and the DAS 16 constitute a detector module 20.

[0013] The X-ray tube 11 generates X-rays by irradiating thermoelectrons from the cathode (filament) toward the anode (target) by applying a high voltage from the X-ray high-voltage device 14. The X-ray tube 11 includes a vacuum tube. For example, the X-ray tube 11 is a rotating anode type X-ray tube that generates X-rays by irradiating a rotating anode with thermoelectrons.

[0014] The wedge 12 is a filter for adjusting the amount of X-rays irradiated from the X-ray tube 11 to the subject P. The wedge 12 attenuates the X-rays passing through itself so that the distribution of the amount of X-rays irradiated from the X-ray tube 11 to the subject P becomes a predetermined distribution. The wedge 12 is also called a wedge filter or a bow-tie filter. The wedge 12 is, for example, made by processing aluminum to have a predetermined target angle and a predetermined thickness.

[0015] The collimator 13 is a mechanism for narrowing down the irradiation range of the X-rays that have passed through the wedge 12. The collimator 13 narrows down the irradiation range of the X-rays, for example, by forming a slit by combining a plurality of lead plates. The collimator 13 may also be called an X-ray aperture. The narrowing range of the collimator 13 may be mechanically drivable.

[0016] The X-ray high-voltage device 14 has, for example, a high-voltage generator (not shown) and an X-ray control device (not shown). The high-voltage generator has an electric circuit including a transformer and a rectifier, etc., and generates a high voltage to be applied to the X-ray tube 11. The X-ray control device controls the output voltage of the high-voltage generator according to the amount of X-rays to be generated by the X-ray tube 11. The high-voltage generator may be one that boosts the voltage by the above-described transformer, or may be one that boosts the voltage by an inverter. The X-ray high-voltage device 14 may be provided on the rotating frame 17, or may be provided on the side of a fixed frame (not shown) of the gantry device 10.

[0017] The X-ray detector 15 detects the intensity of the X-rays generated by the X-ray tube 11 and incident after passing through the subject P. The X-ray detector 15 outputs an electrical signal (which may also be an optical signal, etc.) corresponding to the detected intensity of the X-rays to the DAS 16. The X-ray detector 15 has, for example, a plurality of X-ray detection element arrays. Each of the plurality of X-ray detection element arrays has a plurality of X-ray detection elements arranged in the channel direction along an arc centered on the focal point of the X-ray tube 11. The plurality of X-ray detection element arrays are arranged in the slice direction (column direction, row direction).

[0018] The X-ray detector 15 is, for example, a direct detection type detector. As the X-ray detector 15, for example, a semiconductor diode with electrodes attached to both ends of the semiconductor is applicable. X-ray photons incident on the semiconductor are converted into electron-hole pairs. The number of electron-hole pairs generated by the incidence of one X-ray photon depends on the energy of the incident X-ray photon. The electrons and holes are respectively attracted to a pair of electrodes formed at both ends of the semiconductor. The pair of electrodes generates an electrical pulse having a pulse height value corresponding to the charge of the electron-hole pairs. One electrical pulse has a pulse height value corresponding to the energy of the incident X-ray photon.

[0019] The DAS 16 collects, for example, count data (count value) indicating the count number of X-ray photons detected by the X-ray detector 15 for a plurality of energy bins according to a control signal from the control device 18. The count values for the plurality of energy bins correspond to the energy spectrum of the X-rays incident on the X-ray detector 15, which is deformed according to the response characteristics of the X-ray detector 15. The DAS 16 outputs detection data based on digital signals to the console device 40. The detection data is the digital value of the count value identified by the channel number, column number of the generating X-ray detection element, and the view number indicating the collected view. The view number is a number that changes according to the rotation of the rotary frame 17, and is, for example, a number incremented according to the rotation of the rotary frame 17. Therefore, the view number is information indicating the rotation angle of the X-ray tube 11. The view period is the period that falls between the rotation angle corresponding to a certain view number and the rotation angle corresponding to the next view number.

[0020] DAS16 may detect the view switching by a timing signal input from the control device 18, may detect it by an internal timer, or may detect it by a signal acquired from a sensor (not shown). When X-rays are continuously irradiated by the X-ray tube 11 during a full scan, DAS16 collects a detection data group for the entire circumference (360 degrees). When X-rays are continuously irradiated by the X-ray tube 11 during a half scan, DAS16 collects detection data for a half circumference (180 degrees + fan angle). DAS16 processes the detection data detected by the semiconductor detector.

[0021] FIG. 2 is a diagram showing an example of the configuration of DAS16 according to the first embodiment. DAS16 includes readout channels corresponding to the number of channels according to the number of X-ray detection elements. These multiple readout channels are mounted in parallel on an integrated circuit such as an ASIC. FIG. 2 shows only the configuration of DAS16-1 for one readout channel.

[0022] DAS16-1 includes a preamplifier circuit 61, a waveform shaping circuit 63, a plurality of pulse height discrimination circuits 65, a plurality of counting circuits 67, and an output circuit 69. The preamplifier circuit 61 amplifies a detection electrical signal DS (current signal) from the connected X-ray detection element. For example, the preamplifier circuit 61 converts a current signal from the connected X-ray detection element into a voltage signal having a voltage value (pulse height value) proportional to the charge amount of the current signal. The waveform shaping circuit 63 is connected to the preamplifier circuit 61. The waveform shaping circuit 63 shapes the waveform of the voltage signal from the preamplifier circuit 61. For example, the waveform shaping circuit 63 reduces the pulse width of the voltage signal from the preamplifier circuit 61.

[0023] A plurality of counting channels corresponding to the number of energy bands (energy bins) are connected to the waveform shaping circuit 63. When n energy bins are set, n counting channels are provided in the waveform shaping circuit 63. Each counting channel includes a pulse height discrimination circuit 65-n and a counting circuit 67-n.

[0024] Each of the pulse height discrimination circuits 65-n discriminates the energy of X-ray photons detected by the X-ray detection element, which is the pulse height value of the voltage signal from the waveform shaping circuit 63. For example, the pulse height discrimination circuit 65-n includes a comparison circuit 653-n. A voltage signal from the waveform shaping circuit 63 is input to one input terminal of each of the comparison circuits 653-n. A reference signal TH (reference voltage value) corresponding to a different threshold value is supplied from the control device 18 to the other input terminal of each of the comparison circuits 653-n.

[0025] For example, a reference signal TH-1 is supplied to the comparison circuit 653-1 for the energy bin B1, a reference signal TH-2 is supplied to the comparison circuit 653-2 for the energy bin B2, and a reference signal TH-n is supplied to the comparison circuit 653-n for the energy bin Bn. Each of the reference signals TH has an upper limit reference value and a lower limit reference value. Each of the comparison circuits 653-n outputs an electrical pulse signal when the voltage signal from the waveform shaping circuit 63 has a pulse height value corresponding to the energy bin corresponding to each of the reference signals TH. For example, the comparison circuit 653-1 outputs an electrical pulse signal when the pulse height value of the voltage signal from the waveform shaping circuit 63 is the pulse height value corresponding to the energy bin B1 (when it is between the reference signals TH-1 and TH-2). On the other hand, the comparison circuit 653-1 for the energy bin B1 does not output an electrical pulse signal when the pulse height value of the voltage signal from the waveform shaping circuit 63 is not the pulse height value corresponding to the energy bin B1. Also, for example, the comparison circuit 653-2 outputs an electrical pulse signal when the pulse height value of the voltage signal from the waveform shaping circuit 63 is the pulse height value corresponding to the energy bin B2 (when it is between the reference signals TH-2 and TH-3).

[0026] The counting circuit 67-n counts the electrical pulse signals from the pulse discriminator circuit 65-n at a readout period that matches the view switching period. For example, the control device 18 supplies a trigger signal TS to the counting circuit 67-n at the switching timing of each view. Triggered by the supply of the trigger signal TS, the counting circuit 67-n adds 1 to the count number stored in the internal memory every time an electrical pulse signal is input from the pulse discriminator circuit 65-n. Triggered by the supply of the next trigger signal, the counting circuit 67-n reads out the data of the count number (i.e., the count value) accumulated in the internal memory and supplies it to the output circuit 69. Also, the counting circuit 67-n resets the count number stored in the internal memory to the initial value every time the trigger signal TS is supplied. In this way, the counting circuit 67-n counts the count number for each view.

[0027] The output circuit 69 is connected to the counting circuits 67-n for a plurality of readout channels mounted on the X-ray detector 15. The output circuit 69 integrates the count values from the counting circuits 67-n for the plurality of readout channels for each of the plurality of energy bins to generate the count values for the plurality of readout channels for each view. The count value for each energy bin is a set of count number data defined by the channel, segment (column), and energy bin. The count value for each energy bin is transmitted to the console device 40 in view units. The count value in view units is called the count data set CS. Further, the output circuit 69 transmits the data detected for each pixel detected by the X-ray detector 15 to the console device 40. The detected data includes at least one of the data detected for each pixel and the count value for each energy bin.

[0028] Returning to FIG. 1, the rotating frame 17 is an annular member that oppositely supports the X-ray tube 11, the wedge 12, and the collimator 13, and the X-ray detector 15. The rotating frame 17 is rotatably supported by a fixed frame about the subject P introduced therein. The rotating frame 17 further supports the DAS 16. The detection data output by the DAS 16 is transmitted from a transmitter having a light-emitting diode (LED) provided on the rotating frame 17 to a receiver having a photodiode provided on a non-rotating portion (e.g., the fixed frame) of the gantry device 10 by optical communication, and is transferred to the console device 40 by the receiver. Note that the method for transmitting the detection data from the rotating frame 17 to the non-rotating portion is not limited to the method using the above-described optical communication, and any non-contact type transmission method may be employed. The rotating frame 17 is not limited to an annular member as long as it can support and rotate the X-ray tube 11 and the like, and may be a member such as an arm.

[0029] The photon counting CT apparatus 1 is, for example, a Rotate / Rotate-Type X-ray CT apparatus (third-generation CT) in which both the X-ray tube 11 and the X-ray detector 15 are supported by the rotating frame 17 and rotate around the subject P. However, the present invention is not limited to this, and a Stationary / Rotate-Type X-ray CT apparatus (fourth-generation CT) in which a plurality of X-ray detection elements arranged in an annular shape are fixed to the fixed frame and the X-ray tube 11 rotates around the subject P may be used.

[0030] The control device 18 has a processing circuit having a processor such as a CPU (Central Processing Unit). The control device 18 receives an input signal from an input interface attached to the console device 40 or the gantry device 10 and controls the operations of the gantry device 10, the bed device 30, and the DAS 16. For example, the control device 18 rotates the rotating frame 17 or tilts the gantry device 10. When tilting the gantry device 10, the control device 18 rotates the rotating frame 17 about an axis parallel to the Z-axis direction based on the tilt angle (tilt angle) input to the input interface. The control device 18 grasps the rotation angle of the rotating frame 17 based on the output of a sensor (not shown) or the like. Further, the control device 18 controls the energy bin (reference signal TH) of the DAS 16. The control device 18 may be provided on the gantry device 10 or may be provided on the console device 40.

[0031] <Bed device 30> The bed device 30 is a device that places and moves the subject P to be scanned and introduces it into the rotating frame 17 of the gantry device 10. The bed device 30 includes, for example, a base 31, a bed driving device 32, a top plate 33, and a support frame 34. The base 31 includes a housing that supports the support frame 34 so as to be movable in the vertical direction (Y-axis direction). The bed driving device 32 includes a motor and an actuator. The bed driving device 32 moves the top plate 33 in the longitudinal direction (Z-axis direction) of the top plate 33 along the support frame 34. Further, the bed driving device 32 moves the top plate 33 in the vertical direction (Y-axis direction). The top plate 33 is a plate-like member on which the subject P is placed.

[0032] The bed driving device 32 may move not only the top plate 33 but also the support frame 34 in the longitudinal direction of the top plate 33. Conversely, the gantry device 10 may be movable in the Z-axis direction, and the rotation frame 17 may be controlled to come around the subject P by the movement of the gantry device 10. Also, a configuration in which both the gantry device 10 and the top plate 33 are movable may be adopted. Further, the photon counting CT device 1 may be a device in which the subject P is scanned in a standing or sitting position. In this case, the photon counting CT device 1 has a subject support mechanism instead of the bed device 30, and the gantry device 10 rotates the rotation frame 17 about an axis perpendicular to the floor surface.

[0033] <Console device 40> The console device 40 includes, for example, a memory 41, a display 42, an input interface 43, a network connection circuit 44, and a processing circuit 50. In the present embodiment, the console device 40 is described as a separate body from the gantry device 10, but a part or all of the components of the console device 40 may be included in the gantry device 10.

[0034] The memory 41 is realized by, for example, a semiconductor memory element such as a RAM (Random Access Memory) or a flash memory, a hard disk, an optical disk, or the like. The memory 41 stores, for example, detection data, projection data, reconstructed image data (photon counting CT image data), information about the subject P, imaging condition data D1, correction reference data D2, and the like. The memory 41 stores, for example, count values related to a plurality of energy bins transmitted from the gantry device 10. These data may be stored not in the memory 41 (or in addition to the memory 41) but in an external memory that the photon counting CT device 1 can communicate with. The external memory is controlled by, for example, a cloud server that manages the external memory and accepts read / write requests.

[0035] The imaging condition data D1 includes information on the imaging conditions used during the past CT imaging of the subject. The imaging condition data D1 is used, for example, to determine the range of energy bins to be used for generating a new photon-counting CT image corresponding to the past CT image among a plurality of energy bins set in the photon-counting CT apparatus 1. The imaging condition data D1 is included, for example, in the imaging examination order transmitted from a Radiology Information System (hereinafter referred to as RIS), which is a computer system for providing business support in the image diagnosis department, and is stored in the memory 41 of the console device 40. Alternatively, the imaging condition data D1 is the one in which the imaging conditions during past imaging using the photon-counting CT apparatus 1 are stored in the memory 41 of the console device 40. FIG. 3 is a diagram showing an example of the imaging condition data D1 according to the first embodiment. As shown in FIG. 3, in this imaging condition data D1, imaging conditions such as "imaging date and time", "device type (conventional CT device / photon-counting CT device)", "tube voltage", and "tube current" are linked and registered for the "subject ID" that identifies the subject.

[0036] Returning to FIG. 1, the display 42 displays various types of information. For example, the display 42 displays a medical image (CT image) generated by the processing circuit, a GUI (Graphical User Interface) image for receiving various operations by an operator such as a doctor or a technician, and the like. The display 42 is, for example, a liquid crystal display, a CRT (Cathode Ray Tube), an organic EL (Electroluminescence) display, or the like. The display 42 may be provided on the gantry device 10. The display 42 may be a desktop type, or may be a display device (for example, a tablet terminal) capable of wireless communication with the main body of the console device 40. The display 42 is an example of a "display device".

[0037] The input interface 43 receives various input operations by the operator and outputs an electrical signal indicating the content of the received input operation to the processing circuit 50. For example, the input interface 43 receives input operations such as collection conditions when collecting detection data or projection data, reconstruction conditions when reconstructing a CT image, image processing conditions when generating a post-processing image from a CT image, and setting conditions for energy bins. For example, the input interface 43 is realized by a mouse, a keyboard, a touch panel, a trackball, a switch, a button, a joystick, a camera, an infrared sensor, a microphone, etc.

[0038] The input interface 43 may be provided in the gantry device 10. Further, the input interface 43 may be realized by a display device (for example, a tablet terminal) capable of wireless communication with the main body of the console device 40. Note that in this specification, the input interface is not limited to those having physical operation components such as a mouse and a keyboard. For example, an electrical signal processing circuit that receives an electrical signal corresponding to an input operation from an external input device provided separately from the device and outputs this electrical signal to the control circuit is also included in the examples of the input interface.

[0039] The network connection circuit 44 includes, for example, a network card having a printed circuit board, or a wireless communication module, etc. The network connection circuit 44 implements an information communication protocol corresponding to the form of the network to be connected.

[0040] The processing circuit 50 controls the overall operation of the photon counting CT apparatus 1, the operation of the gantry apparatus 10, and the operation of the bed apparatus 30. The processing circuit 50 executes, for example, a system control function 51, a pre-processing function 52, a reconstruction function 53, an image processing function 54, a scan control function 55, a display control function 56, and the like. The reconstruction function 53 includes, for example, an acquisition function 53-1, a determination function 53-2, a generation function 53-3, and the like. These components are realized, for example, by a hardware processor (computer) executing a program (software) stored in the memory 41. The hardware processor means, for example, a circuit such as a CPU, a GPU (Graphics Processing Unit), an ASIC, a programmable logic device (for example, a Simple Programmable Logic Device (SPLD) or a Complex Programmable Logic Device (CPLD), a Field Programmable Gate Array (FPGA)).

[0041] Instead of storing the program in the memory 41, it may be configured to directly incorporate the program into the circuit of the hardware processor. In this case, the hardware processor realizes its functions by reading and executing the program incorporated in the circuit. The hardware processor is not limited to being configured as a single circuit, and may be configured as one hardware processor by combining a plurality of independent circuits to realize each function. Also, a plurality of components may be integrated into one hardware processor to realize each function.

[0042] Each component of the console device 40 or the processing circuit 50 may be realized in a decentralized manner by a plurality of hardware. The processing circuit 50 may be realized not by a component of the console device 40 but by a processing device that can communicate with the console device 40. The processing device may be, for example, a workstation connected to one photon-counting CT device, or a device (such as a cloud server) that is connected to a plurality of photon-counting CT devices and collectively executes processing equivalent to that of the processing circuit 50 described below. The photon-counting CT device 1, the console device 40, the workstation, the cloud server, or a combination thereof is an example of a "medical information processing device".

[0043] The system control function 51 controls various functions of the processing circuit 50 based on the input operation received by the input interface 43. The system control function 51 performs, for example, setting of energy bins, setting of tube voltage and tube current of the X-ray tube 11, etc. The system control function 51 outputs the set conditions of the energy bins, etc. to the control device 18.

[0044] The preprocessing function 52 performs preprocessing such as offset correction processing, sensitivity correction processing between channels, and beam hardening correction on the detection data output by the DAS 16.

[0045] The reconstruction function 53 reconstructs a photon-counting CT image of the subject P based on the detection data (count value). The reconstruction function 53 calculates the X-ray absorption amounts for each of a plurality of reference substances based on the count values for a plurality of energy bins, the energy spectrum of the X-rays incident on the subject P, and a response function representing the detector response characteristics stored in the memory 41. The process of obtaining the X-ray absorption amount for each reference substance in this way is also called material discrimination. As the reference substances, any substances such as calcium, calcification, bone, fat, muscle, air, organs, lesion sites, hard tissues, soft tissues, and contrast agents can be set. The reconstruction function 53 reconstructs a photon-counting CT image representing the spatial distribution of the reference substance to be imaged among the plurality of reference substances based on the calculated X-ray absorption amounts for each of the plurality of reference substances, and stores the generated CT image data in the memory 41.

[0046] The acquisition function 53-1 acquires the imaging conditions including the tube voltage of the X-ray tube used when taking a past CT image of the subject. The acquisition function 53-1 acquires, as the imaging conditions, in addition to the tube voltage of the X-ray tube, the settings of the wedge 12 and the energy bins, etc. The acquisition function 53-1 acquires the imaging conditions (imaging condition data D1), for example, from the memory 41. The acquisition function 53-1 is an example of an "acquisition unit".

[0047] The determination function 53-2 determines, based on the acquired imaging conditions, the range of the energy bins to be used for generating a photon-counting CT image corresponding to a past CT image among the plurality of energy bins set for counting the X-ray photons detected by a new imaging of the subject using the photon-counting CT apparatus 1 for each energy band. The past CT image is an image taken by a conventional CT apparatus. The past CT image may be an integrated image taken by a photon-counting CT apparatus. The determination function 53-2 is an example of a "determination unit".

[0048] The generation function 53-3 generates a photon-counting CT image using the count value within the determined energy bin range. The generation function 53-3 is an example of a "generation unit".

[0049] The image processing function 54 converts the CT image data into three-dimensional image data or cross-sectional image data of an arbitrary cross-section by a known method based on the input operation received by the input interface 43. The conversion into three-dimensional image data may be performed by the preprocessing function 52.

[0050] The scan control function 55 controls the collection process of detection data in the gantry device 10 by instructing the X-ray high voltage device 14, the DAS 16, the control device 18, and the bed drive device 32. The scan control function 55 controls the operations of each part when collecting a positioning image and when taking an image for diagnosis.

[0051] The display control function 56 causes the display 42 to display a medical image (photon-counting CT image) generated by the processing circuit and a GUI image or the like that receives various operations by an operator such as a doctor or a technician. The display control function 56 causes the display 42 to display a photon-counting CT image corresponding to a past CT image of the subject generated by the generation function 53-3. The display control function 56 is an example of a "display control unit".

[0052] With the above configuration, the photon-counting CT apparatus 1 scans the subject P in scan modes such as helical scan, conventional scan, and step-and-shoot. The helical scan is a mode in which the rotating frame 17 is rotated while moving the top plate 33 to scan the subject P in a spiral shape. The conventional scan is a mode in which the rotating frame 17 is rotated with the top plate 33 stationary to scan the subject P in a circular orbit. The step-and-shoot is a mode in which the position of the top plate 33 is moved at regular intervals to perform conventional scans in a plurality of scan areas.

[0053] [Processing Flow] Next, a series of processes of the image generation process of the photon counting CT apparatus 1 will be described centering on the processes of the console apparatus 40. FIG. 4 is a flowchart showing an example of the image generation process by the photon counting CT apparatus 1 according to the first embodiment. The image generation process shown in FIG. 4 is started, for example, when an operator inputs an instruction to generate a photon counting CT image via the input interface 43 after performing a scan on the subject P or the like.

[0054] First, the acquisition function 53-1 acquires the imaging conditions used at the time of imaging the past (previous) CT image of the subject P from the imaging condition data D1 stored in the memory 41 (step S101). These imaging conditions include at least the tube voltage (kVp) of the X-ray tube used at the time of imaging the past CT image of the subject P. In the following, an example will be described in which the imaging conditions of "apparatus type (CT1 (EID))" and "tube voltage (80 kVp)" associated with the "subject ID ("0001")" are acquired from the imaging condition data D1 shown in FIG. 3.

[0055] In addition, when information in a format conforming to the DICOM (Digital Imaging and Communication in Medicine) standard (hereinafter also referred to as "DICOM tag") is attached to the past CT image of the subject P, the acquisition function 53-1 may acquire the imaging conditions from the information of this DICOM tag. Further, when information (such as a character string) indicating the imaging conditions is included in the past CT image, the acquisition function 53-1 may acquire the imaging conditions by performing image analysis on this past CT image. Further, when the past CT image cannot be acquired, the acquisition function 53-1 may acquire the tube voltage or the like input by the operator via the input interface 43 as the imaging conditions.

[0056] Next, the determination function 53-2 determines, based on the obtained imaging conditions, the range of energy bins to be used for generating a new photon-counting CT image corresponding to the past CT image of the subject P among the plurality of energy bins set in the photon-counting CT apparatus 1 (step S103).

[0057] Hereinafter, the content of the determination process (step S103) of the reconstruction conditions by the determination function 53-2 will be specifically described. FIG. 5 is a flowchart showing an example of the determination process of the reconstruction conditions by the photon-counting CT apparatus 1 according to the first embodiment. FIG. 6 is a diagram for explaining data used for reconstruction among the count values for each energy bin by the photon-counting CT apparatus 1 according to the first embodiment.

[0058] First, the determination function 53-2 determines the range of energy bins based on the obtained past imaging conditions (tube voltage at the previous imaging) (step S201). For example, assume that the tube voltage, which is the obtained past imaging condition, is "80 kVp", and count values of X-ray photons for each of six energy bins B1 to B6 (step: 20 keV) have been obtained by a new imaging of the subject P using the photon-counting CT apparatus 1. The energy band of energy bin B1 is 0 to 20 keV, the energy band of energy bin B2 is 20 to 40 keV, the energy band of energy bin B3 is 40 to 60 keV, the energy band of energy bin B4 is 60 to 80 keV, the energy band of energy bin B5 is 80 to 100 keV, and the energy band of energy bin B6 is 100 to 120 keV. In this case, as shown in FIG. 6, the determination function 53-2 determines the range of energy bins B1 to B4 (that is, the energy band 0 to 80 keV), which is close to the condition of the value "80 kVp" of the tube voltage that is the obtained past imaging condition, as the range of energy bins to be used for generating the photon-counting CT image. Next, the determination function 53-2 determines the reconstruction conditions including the conditions of the determined range of energy bins (step S203).

[0059] That is, the determination function 53-2 determines, as the energy bin range, the range from the energy bin in the lowest energy band among the plurality of energy bins to the energy bin having an energy band corresponding to the value of the tube voltage included in the imaging conditions.

[0060] In addition, when the acquired past imaging conditions (tube voltage at the previous imaging) do not match the energy band division (boundary value between energy bins) of the energy bins of the photon counting CT apparatus 1, the determination function 53-2 determines, as the energy range, up to the energy bin division close to the acquired past imaging conditions (tube voltage at the previous imaging). For example, when the tube voltage at the previous imaging in the past imaging conditions is "95 kVp", the determination function 53-2 determines, as the energy bin range used for generating the photon counting CT image, the energy bin range (energy bins B1 to B5) from the lowest energy band energy bin to "100 keV", which is the closest to the numerical value of this tube voltage "95 kVp".

[0061] That is, the determination function 53-2 determines, as the energy bin range, the range of energy bins from the energy bin in the lowest energy band among the plurality of energy bins to the boundary value between the energy bins closest to the value of the tube voltage included in the imaging conditions.

[0062] Alternatively, the determination function 53-2 may distinguish all energy bins whose upper limit value of the energy bin is less than or equal to the tube voltage at the previous imaging under the past imaging conditions from the energy bins in the energy band including the tube voltage at the previous imaging, and determine them as the range of energy bins used for generating the photon counting CT image. For example, when the tube voltage at the previous imaging under the past imaging conditions is "95 kVp", the determination function 53-2 distinguishes energy bins B1 to B4 whose upper limit value of the energy bin is less than or equal to this tube voltage "95 kVp" from the energy bin B5 including this tube voltage "95 kVp", and determines them as the range of energy bins used for generating the photon counting CT image. In this case, in the subsequent processing, the generation function 53-3 may generate a photon counting CT image using the count values of the energy bins B1 to B4 and a part of the count value of the energy bin B5 considering the weighted average based on the ratio from the lower limit value of 80 keV in the energy band of the energy bin B5. The ratio of the part of the count value of the energy bin B5 considering the weighted average is calculated, for example, by (95 keV (tube voltage at the previous imaging under the past imaging conditions) - 80 keV (lower limit value of the energy bin B5)) / 20 keV (step) = 0.75.

[0063] That is, the determination function 53-2 determines, as the range of energy bins, the range from the energy bin in the lowest energy band among the plurality of energy bins to the upper energy bin including the value of the tube voltage included in the imaging conditions, and the generation function 53-3 adjusts the count value used for generating the photon counting CT image for the count value of the upper energy bin based on the difference between the lower limit value of the energy band of the upper energy bin and the value of the tube voltage included in the imaging conditions within the determined range of energy bins.

[0064] Alternatively, based on the tube voltage during the previous imaging, the energy bin settings of the photon counting CT apparatus 1 may be changed, a rescan may be executed, the count values in the changed energy bin settings may be collected, and a photon counting CT image corresponding to the past CT image may be generated.

[0065] Note that the determination function 53-2 may determine the energy bin range based on an instruction input by the operator via the input interface 43.

[0066] Returning to FIG. 4, the generation function 53-3 reconstructs a photon counting CT image (step S105) based on the determined reconstruction conditions (using the count values in the determined energy bin range). Here, in order to generate a photon counting CT image corresponding to a past CT image (CT generated by a conventional CT apparatus), the generation function 53-3 reconstructs a photon counting CT image (counting image) based on the combined data of the count values in the energy band of 0 to 80 keV (data obtained by adding the count values of energy bins B1 to B4). Note that the generation function 53-3 may perform an image addition average after reconstructing a plurality of photon counting CT images corresponding to each of the energy bins B1, B2, B3, and B4, and create a photon counting CT image corresponding to the energy band of 0 to 80 keV.

[0067] Next, the display control function 56 causes the photon counting CT image corresponding to the past CT image of the subject, generated by the generation function 53-3, to be displayed on the display 42 (step S107). At this time, the display control function 56 may cause the newly generated photon counting CT image and the past CT image to be displayed on the display 42 so that both can be compared. Thus, the processing of this flowchart ends.

[0068] According to the first embodiment described above, it is possible to compare medical images taken under different imaging conditions in a CT apparatus at the same energy level. Further, by enabling such comparison of medical images, it is possible to assist diagnosis by doctors and the like. Also, in a new imaging using a photon-counting CT apparatus, data in a wide energy range (for example, 0 to 120 keV) can be acquired, and thus it is also possible to generate a spectral image unique to the photon-counting CT apparatus.

[0069] <Second Embodiment> Next, the second embodiment will be described. The intensity and energy of the X-rays (X-ray energy spectrum) output from the X-ray tube 11 change depending on the tube voltage of the X-ray tube 11. For this reason, when a new photon-counting CT image is generated based only on the difference in tube voltage without considering the difference in such X-ray energy spectra, the appearance will be different from that of a past CT image (the reconstruction conditions are different), and it may be difficult to compare the two. Therefore, in the photon-counting CT apparatus 1 of the second embodiment, a photon-counting CT image is reconstructed in consideration of the X-ray energy spectrum output from the X-ray tube 11 or the X-ray energy spectrum incident on the X-ray detector 15. In the following description, the same components and functions as those in the first embodiment are denoted by the same reference numerals as those in the first embodiment, and detailed description thereof is omitted.

[0070] FIG. 7 is a flowchart showing an example of a determination process (step S103) of reconstruction conditions by the photon counting CT apparatus 1 according to the second embodiment. First, the determination function 53-2 determines the range of energy bins based on the acquired past imaging conditions (tube voltage at the previous imaging) (step S301). For example, assume that the tube voltage is "80 kVp" as the past imaging condition, and count values for each of six energy bins B1 to B6 (step: 20 keV) are obtained by a new imaging of the subject P using the photon counting CT apparatus 1. In this case, the determination function 53-2 determines the range of energy bins B1 to B4 (that is, the energy band 0 to 80 keV), which is close to the condition of the acquired past imaging condition of the tube voltage (80 kVp), as the range of energy bins to be used for generating the photon counting CT image.

[0071] Next, for each energy bin included in the determined energy bin range, the determination function 53-2 determines a correction factor (weighting factor) in consideration of the X-ray energy spectrum (step S303). FIG. 8 is a graph showing the relationship between the X-ray energy and the X-ray intensity for each tube voltage according to the second embodiment. As shown in FIG. 8, the intensity of the X-rays output by the X-ray tube 11 at a tube voltage of 120 kVp (energy spectrum ES1) is higher overall than the intensity of the X-rays output by the X-ray tube 11 at a tube voltage of 80 kVp (energy spectrum ES2), but the difference in intensity varies depending on the energy value. Specifically, as the energy value increases, the difference in intensity between the two becomes even greater. Therefore, a correction factor is determined for each energy bin to reduce the influence of such a difference in the X-ray energy spectrum (the difference in intensity between the two). For example, for the energy band (20 to 40 keV) corresponding to the energy bin B2, the ratio of the X-ray intensity of the energy spectrum ES2 to the X-ray intensity of the energy spectrum ES1 at the average energy (30 keV) is determined as the correction factor. Reference data for determining such a correction factor is stored in the memory 41 as correction reference data D2. Note that the correction factor may not be determined each time, but may be calculated in advance and stored in the memory 41 as correction reference data D2. Further, the determination function 53-2 may determine (adjust) the correction factor based on an instruction input by the operator via the input interface 43.

[0072] Next, the determination function 53-2 determines reconstruction conditions including the determined energy bin range and the correction factor for each energy bin (step S305). Thereafter, the generation function 53-3 reconstructs a photon-counting CT image based on the reconstruction conditions determined by the determination function 53-2 (using the count value and the correction factor for each determined energy bin). For example, for the energy bin B2, the generation function 53-3 does not use the count value as it is, but uses a value obtained by multiplying the count value by the correction factor (using the reduced count value) to generate a photon-counting CT image.

[0073] That is, the determination function 53-2 determines a correction coefficient based on the difference between the energy spectrum of X-rays output by the X-ray tube at the tube voltage of the X-ray tube used during past CT image acquisition and the energy spectrum of X-rays output by the X-ray tube at the tube voltage during new acquisition using the photon counting CT apparatus. The generation function 53-3 corrects the count values included in the determined energy bin range using the correction coefficient, and generates a photon counting CT image using the corrected count values.

[0074] Still, the intensity of X-rays (X-ray energy spectrum) also changes depending on the tube current of the X-ray tube 11. Therefore, when the tube current at the time of taking a past CT image differs from the tube current at the time of a new imaging, a correction factor may be determined in consideration of the influence of the energy spectrum resulting from this difference in tube current. FIG. 9 is a graph showing the relationship between the X-ray energy and the X-ray intensity for each tube current according to the second embodiment. As shown in FIG. 9, the intensity of the X-rays (energy spectrum ES3) output by the X-ray tube 11 at a tube current of 600 mA is overall higher than the intensity of the X-rays (energy spectrum ES4) output by the X-ray tube 11 at a tube current of 300 mA, but the difference in intensity changes depending on the energy value. Therefore, a correction factor is determined to reduce the influence of such a difference in the X-ray energy spectrum (the difference in intensity between the two). For example, for the energy band (20 to 40 keV) corresponding to the energy bin B2, the ratio of the X-ray intensity of the energy spectrum ES4 to the X-ray intensity of the energy spectrum ES3 at the average energy (30 keV) is determined as the correction factor. Here, when the tube current at the time of taking a past CT image is "300 mA" and the tube current at the time of a new imaging is "600 mA", the generation function 53-3 generates a photon-counting CT image based on the reconstruction conditions determined by the determination function 53-2 (using the determined count value for each energy bin and the correction factor). For example, for the energy bin B2, the generation function 53-3 does not use the count value as it is, but uses a value obtained by multiplying the count value by the correction factor calculated based on the tube current (using the reduced count value) to generate a photon-counting CT image.

[0075] That is, the determination function 53-2 determines a correction coefficient based on the difference between the energy spectrum of X-rays that pass through the subject and enter the X-ray detector at the tube voltage of the X-ray tube used during past CT image acquisition, and the energy spectrum of X-rays that enter the X-ray detector at the tube voltage during new acquisition using a photon-counting CT apparatus. The generation function 53-3 corrects the count values included in the determined energy bin range using the correction coefficient, and generates a photon-counting CT image using the corrected count values.

[0076] Note that instead of the spectrum of X-rays output from the X-ray tube 11, spectrum data of X-rays that pass through the subject and enter the X-ray detector 15 may be acquired in advance as correction reference data D2. For example, the correction reference data D2 may be collected with a phantom placed on the bed apparatus 30. The determination function 53-2 may determine a correction coefficient for each of the energy bins included in the determined energy bin range using the correction reference data D2 of the energy spectrum of X-rays that enter the X-ray detector 15.

[0077] Note that by adjusting the tube current of the X-ray tube 11, the energy spectra of X-rays output from the X-ray tube 11 or the energy spectra of X-rays that enter the X-ray detector 15 at different tube voltages may be made as uniform as possible.

[0078] According to the second embodiment described above, it is possible to compare medical images taken under different imaging conditions in the CT apparatus at the same energy level. Further, by enabling such comparison of medical images, it is possible to assist diagnosis by doctors and the like. Also, in a new imaging using a photon counting CT apparatus, data in a wide energy range (for example, 0 to 120 keV) can be acquired, so it is also possible to generate a spectral image unique to the photon counting CT apparatus. Further, by reconstructing the photon counting CT image in consideration of the energy spectrum of the X-rays output from the X-ray tube or the energy spectrum of the X-rays incident on the X-ray detector, it becomes possible to compare medical images with more unified imaging conditions.

[0079] Although several embodiments have been described, these embodiments are presented as examples and are not intended to limit the scope of the invention. These embodiments can be implemented in various other forms, and various omissions, replacements, and changes can be made without departing from the gist of the invention. These embodiments and their modifications are included in the scope and gist of the invention, and are also included in the invention described in the claims and its equivalent scope.

Explanation of Reference Numerals

[0080] 1...Photon counting CT apparatus, 10...Gantry apparatus, 11...X-ray tube, 12...Wedge, 13...Collimator, 14...X-ray high voltage apparatus, 15...X-ray detector, 16...Data acquisition system, 17...Rotating frame, 18...Control apparatus, 20...Detector module, 30...Bed apparatus, 31...Base, 32...Bed drive apparatus, 33...Top plate, 34...Support frame, 40...Console apparatus, 41...Memory, 42...Display, 43...Input interface, 44...Network connection circuit, 50...Processing circuit, 51...System control function, 52...Preprocessing function, 53...Reconstruction function, 53-1...Acquisition function, 53-2...Determination function, 53-3...Generation function, 54...Image processing function, 55...Scan control function, 56...Display control function

Claims

1. An acquisition unit that acquires imaging conditions including the tube voltage of an X-ray tube used during past CT imaging of a subject; Among a plurality of energy bins set to count X-ray photons detected by a new imaging using a photon counting CT apparatus for the subject for each energy band, a determination unit that determines a range of energy bins to be used for generating a photon counting CT image corresponding to the past CT image based on the acquired imaging conditions; A generation unit that generates the photon counting CT image using the count values of the X-ray photons in the determined range of energy bins; A medical image processing apparatus comprising:

2. The determination unit determines, as the range of the energy bins, a range from the energy bin in the lowest energy band among the plurality of energy bins to the energy bin having an energy band corresponding to the value of the tube voltage included in the imaging conditions. The medical image processing apparatus according to Claim 1.

3. The determination unit determines, as the range of the energy bins, a range of energy bins from the energy bin in the lowest energy band among the plurality of energy bins to the boundary value between the energy bins closest to the value of the tube voltage included in the imaging conditions. The medical image processing apparatus according to Claim 1.

4. The determination unit determines, as the range of the energy bins, a range from the energy bin in the lowest energy band among the plurality of energy bins to the upper energy bin including the value of the tube voltage included in the imaging conditions in the energy band, and The generation unit adjusts the count value used for generating the photon counting CT image based on the difference between the lower limit value of the energy band of the upper energy bin and the value of the tube voltage included in the imaging conditions for the count value of the upper energy bin among the count values in the determined range of energy bins. The medical image processing apparatus according to Claim 1.

5. The determination unit determines a correction coefficient based on the difference between the X-ray energy spectrum output by the X-ray tube at the tube voltage of the X-ray tube used during the past CT imaging and the X-ray energy spectrum output by the X-ray tube at the tube voltage during a new imaging using the photon counting CT apparatus. The generation unit corrects the count value included in the determined energy bin range using the correction coefficient, and generates the photon counting CT image using the corrected count value. The medical image processing apparatus according to claim 1.

6. The determination unit determines a correction coefficient based on a difference between an energy spectrum of X-rays that pass through the subject and enter an X-ray detector at the tube voltage of the X-ray tube used during imaging of the past CT image, and an energy spectrum of X-rays that enter the X-ray detector at the tube voltage during new imaging using the photon counting CT apparatus. The generation unit corrects the count value included in the determined energy bin range using the correction coefficient, and generates the photon counting CT image using the corrected count value. The medical image processing apparatus according to claim 1.

7. The apparatus further includes a display control unit that causes a display device to display both the generated photon counting CT image and the past CT image in a comparable manner. The medical image processing apparatus according to any one of claims 1 to 6.

8. A computer acquires imaging conditions including the tube voltage of the X-ray tube used during imaging of a past CT image of a subject, determines a range of energy bins to be used for generating a photon counting CT image corresponding to the past CT image from among a plurality of energy bins set for counting X-ray photons detected by a new imaging using a photon counting CT apparatus for the subject, based on the acquired imaging conditions, generates the photon counting CT image using the count value of the X-ray photons in the determined energy bin range. A medical image processing method.

Images