JP2017205181A - X-ray moving body tracking device, x-ray moving body tracking method, and image processing method - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide an X-ray moving body tracking device capable of tracking a tracking target with a higher degree of precision.SOLUTION: An X-ray moving body tracking device includes a pulse voltage generation part for generating a first pulse voltage and a second pulse voltage different in magnitude from the first pulse voltage at a predetermined interval, an X-ray generation part for generating a first pulse X-ray and a second pulse X-ray based on the first pulse voltage and the second pulse voltage, an X-ray imaging part for imaging first X-ray image data corresponding to the first pulse X-ray passing through a subject, and second X-ray image data corresponding to the second pulse X-ray passing through the subject, an image processing part for generating emphasized image data for emphasizing contrast of a marker region for radiation therapy in the subject using the first X-ray image data and the second X-ray image data, and a detection part for detecting the position of the marker for radiation therapy based on the emphasized image data.SELECTED DRAWING: Figure 1

Description

  Embodiments described herein relate generally to an X-ray moving object tracking device, an X-ray moving object tracking method, and an image processing method.

  High-accuracy radiotherapy that protects normal cells and concentrates on the affected area and irradiates with a high dose has been performed clinically. For this high-accuracy radiotherapy, therapeutic beams such as heavy particle beams, proton beams, and X-rays are used. In this radiotherapy using a therapeutic beam, the energy, dose, and incident direction of the therapeutic beam are determined based on the treatment plan so that a large killing effect is produced on the tumor cells that are the target of the affected area. Yes.

  However, there are respiratory movements in the thoracic cavity sandwiching the diaphragm and organs contained in the abdominal cavity. This respiratory movement may cause these organs to move three-dimensionally, which cannot be tracked by body surface movement. For this reason, when there is an affected part target in these organs, an X-ray moving body tracking method for tracking the affected part target three-dimensionally is applied.

  In this X-ray moving body tracking method, a gating irradiation method is used in which the position of the affected part target is tracked using an orthogonal two-way X-ray fluoroscopic imaging apparatus. That is, in the gating irradiation method, the therapeutic beam is irradiated at a timing when the affected part target is located within the gate that is the irradiation range of the therapeutic beam. In this case, a marker for radiation therapy may be used to identify the position of the affected area target, and it is necessary to improve the distinguishability of the marker area for radiation therapy.

Supervised by the Radiation Therapy Subcommittee, “Practical Location Matching and Setup in Radiation Therapy”, First Edition, Japanese Society of Radiological Technology, February 20, 2015, p. 1-143

  An object of an embodiment of the present invention is to provide an X-ray moving body tracking apparatus capable of tracking a tracking target with higher accuracy.

The X-ray moving body tracking apparatus according to the present embodiment is
A pulse voltage generator for generating a first pulse voltage and a second pulse voltage different from the first pulse voltage at a predetermined interval;
An X-ray generator that generates a first pulse X-ray and a second pulse X-ray based on the first pulse voltage and the second pulse voltage;
Imaging first X-ray image data corresponding to the first pulse X-ray transmitted through the subject and second X-ray image data corresponding to the second pulse X-ray transmitted through the subject. An X-ray imaging unit,
An image processing unit that generates enhanced image data that enhances the contrast of the radiotherapy marker region in the subject using the first X-ray image data and the second X-ray image data;
A detection unit for detecting a position of the radiotherapy marker based on the enhanced image data;
It is characterized by providing.

The X-ray moving body tracking method according to the present embodiment is:
A pulse voltage generation step of generating a first pulse voltage and a second pulse voltage having a magnitude different from that of the first pulse voltage at a predetermined interval;
An X-ray generation step of generating a first pulse X-ray and a second pulse X-ray based on the first pulse voltage and the second pulse voltage;
Imaging first X-ray image data corresponding to the first pulse X-ray transmitted through the subject and second X-ray image data corresponding to the second pulse X-ray transmitted through the subject. X-ray imaging process
Using the first X-ray image data and the second X-ray image data, an image data processing step for generating enhanced image data for enhancing the contrast of the radiotherapy marker region in the subject;
A detection step of detecting a position of the radiotherapy marker based on the enhanced image data;
It is characterized by providing.

The image processing method according to the present embodiment is as follows:
First image data corresponding to the first X-ray transmitted through the subject, and second X-ray having characteristics different from those of the first X-ray, and the second X-ray transmitted through the subject An acquisition step of acquiring second image data corresponding to the line;
A calculation step of calculating a weighting coefficient based on a pixel value obtained from the bone region of the first image data and a pixel value obtained from the bone region of the second image data;
A difference processing step for performing a difference process between the first X image data and the second X image data using the weighting coefficient;
It is characterized by providing.

  According to the present embodiment, it is possible to provide an X-ray moving body tracking apparatus capable of tracking a tracking target with higher accuracy.

The block diagram explaining the whole structure of the X-ray moving body tracking device which concerns on 1st embodiment. The figure which shows the detailed structure of the high voltage pulse generation part for cathode grounding systems. The figure which shows the detailed structure of the high voltage pulse generation part for middle point grounding systems. The figure which shows the timing chart of the high voltage pulse generation part for cathode grounding systems in the 1st mode. The figure which shows the timing chart of the high voltage pulse generation part for cathode grounding systems in the 2nd mode. The figure which shows the timing chart of the high voltage pulse generation part for middle point grounding systems. The figure which shows the timing chart of the 1st X-ray tube and 1st X-ray imaging part in 1st mode. The figure which shows the timing chart of the 1st X-ray tube and 1st X-ray imaging part in 2nd mode. The figure explaining a difference process. The block diagram which shows the structure of the X-ray moving body tracking device which concerns on 2nd Embodiment. The block diagram which shows the structure of the X-ray moving body tracking device which concerns on 3rd Embodiment.

  Embodiments according to the present invention will be described below with reference to the drawings. This embodiment does not limit the present invention.

(First embodiment)
The X-ray moving body tracking apparatus according to the first embodiment irradiates pulse X-rays having a plurality of different tube voltages based on the first pulse voltage and the second pulse voltage having a voltage different from the first pulse voltage. A plurality of X-ray image data having different tube voltages is used to generate enhanced image data that enhances the contrast of the radiotherapy marker region in the subject. More details will be described below.

  Based on FIG. 1, the whole structure of the X-ray moving body tracking device 1 which concerns on this embodiment is demonstrated. FIG. 1 is a block diagram illustrating the overall configuration of the X-ray moving body tracking apparatus 1 according to the first embodiment.

  As shown in FIG. 1, the X-ray moving body tracking apparatus 1 according to the present embodiment tracks the position of an affected part target that undergoes respiratory movement. More specifically, the X-ray moving body tracking apparatus 1 includes a first high voltage pulse generation unit 100A, a second high voltage pulse generation unit 100B, a first X-ray tube unit 102A, and a second X-ray tube unit. The tube section 102B, the first X-ray tube support section 104A, the second X-ray tube support section 104B, the first collimator section 106A, the second collimator section 106B, the setting section 108, the treatment Table 110, first X-ray imaging unit 112A, second X-ray imaging unit 112B, first X-ray imaging unit 114A, second X-ray imaging unit 114B, and first 2D image data Output unit 116A, second 2D image data output unit 116B, first image processing unit 118A, second image processing unit 118B, synchronization control unit 120, three-dimensional conversion unit 122, target coordinate output Unit 124 and an irradiation permission determination unit 126. That.

  Further, the position of the affected area target is marked with a radiation therapy marker. For this reason, the X-ray moving body tracking apparatus 1 tracks the position of the radiation therapy marker as the position of the affected area target to be tracked.

  The first high voltage pulse generator 100A generates a high voltage pulse. The first high voltage pulse generator 100A can continuously generate high voltage pulses. For example, the high voltage pulse generator 100A generates the second pulse voltage at a predetermined interval after generating the first pulse voltage. The detailed configuration of the first high voltage pulse generator 100A will be described later.

  The pulse voltage generated by the first high voltage pulse generator 100A is applied to the first X-ray tube unit 102A. The first X-ray tube unit 102A generates pulse X-rays based on the applied pulse voltage. That is, the first X-ray tube unit 102A generates a first pulse X-ray and a second pulse X-ray based on the first pulse voltage and the second pulse voltage.

  More specifically, the pulse voltage is applied between the cathode and anode of the first X-ray tube portion 102A. The electrons ejected from the cathode are given energy according to the potential of the applied pulse voltage. Then, X-rays having energy corresponding to the energy of electrons colliding with the anode are generated from the anode. That is, the tube voltage and the generated X-ray energy are substantially proportional. For this reason, the first X-ray tube unit 102A generates X-rays having different energies according to the voltage value of the applied pulse voltage. The voltage applied to the X-ray tube is generally called a tube voltage. That is, the voltage value of the pulse voltage generated by the first high voltage pulse generator 100A corresponds to the tube voltage.

  Here, the X-ray absorption rate of the subject 10, that is, the patient varies depending on the energy of the X-rays. That is, when the tube voltage is increased, the energy of the generated X-rays increases, and is less likely to be attenuated when passing through the subject 10. For this reason, when the subject 10 is irradiated with X-rays having different energies, X-ray images having different intensity distributions are generated for the X-rays having different energies. Note that the two-dimensional intensity distribution of X-rays transmitted through the subject 10 is called an X-ray image or a radiation image.

  The first X-ray tube support portion 104A supports the first X-ray tube portion 102A. The first collimator unit 106A is attached to the X-ray output surface of the first X-ray tube unit 102A supported by the first X-ray tube support unit 104A, and controls the irradiation range of the first pulse X-ray. Thereby, the irradiation range of the first pulse X-ray is limited to the region of interest of the subject 10.

  The setting unit 108 sets conditions for generating the first pulse voltage and the second pulse voltage. In accordance with the setting in the setting unit 108, the first high voltage pulse generating unit 100A generates a high voltage pulse. The setting unit 108 also includes a pulse width of each of the first pulse voltage and the second pulse voltage, a voltage value of the pulse voltage, a generation interval between the first pulse voltage and the second pulse voltage, and a first pulse. A generation frequency for repeatedly generating the voltage and the second pulse voltage is set.

  Here, the pulse width is a time during which the pulse is continuously generated, and is set in a range of 0.1 milliseconds to 3 milliseconds. As the pulse width is increased, the intensity of X-rays can be increased and noise can be reduced. On the other hand, when the pulse width is increased, the affected part that moves in a respiratory manner moves, and the captured image data is blurred or misaligned. For this reason, 1 millisecond is used as the standard value here.

  The pulse voltage, that is, the tube voltage is set in the range of 60 KV to 150 KV. This voltage value is determined according to the imaging region. For example, when the first pulse voltage is set to a standard value, 80 KV is set for the chest and 130 KV is set for the lumbar spine.

  Further, as described above, the X-ray energy increases as the tube voltage is increased. For this reason, the pulse width may be set shorter as the pulse voltage is set higher. Thereby, it is possible to reduce the integrated amount of the energy of the X-rays irradiated to the subject 10.

  The generation interval is set in the range of 1 millisecond to 33 milliseconds. This generation interval is the time from when the first pulse voltage is generated until the second pulse voltage is generated. If the generation interval is shortened as much as possible in a region where respiratory movement is large, blur in image processing described later is reduced. On the other hand, the shorter the generation interval, the higher the capacity of the first high-voltage pulse generator 100A becomes necessary. For this reason, here, 3 milliseconds is used as a standard value as a time in which the influence of afterglow caused by phosphor emission decay time such as an image intensifier described later can be ignored.

  The occurrence frequency is set in the range of 1 Hz to 60 Hz. This generation frequency is a numerical value indicating how many times the first pulse voltage is generated per second. For example, if the frequency is 30 Hz, the first pulse voltage is generated 30 times per second. Similarly, the second pulse voltage is also generated 30 times per second. The imaging cycle is the reciprocal of the occurrence frequency.

  The greater the frequency of occurrence, the higher the exposure dose. On the other hand, if the occurrence frequency is reduced, there is a possibility that the acquisition interval of the position extraction of the affected part target is increased. For this reason, 30 Hz is used as a standard value here.

  The setting unit 108 may set the generation conditions according to a table in which the generation conditions are determined according to the imaging region. In this case, the setting unit 108 sets the generation condition according to the imaging region instructed by the photographer in the first high voltage pulse generation unit 100A.

  Furthermore, the setting unit 108 can set the first mode and the second mode according to the imaging region. The first mode is a mode in which the voltage values of the first pulse voltage and the second pulse voltage are different. The second mode is a mode in which the voltage values of the first pulse voltage and the second pulse voltage are equivalent.

  The treatment table 110 is placed with the subject 10 lying on his back fixed. The treatment table 110 is made of a material such as CFRP (carbon fiber reinforced plastic) that has high strength and relatively low X-ray absorption.

  The first X-ray imaging unit 112A converts the intensity distribution of the pulse X-rays irradiated through the first collimator unit 106A into image data. That is, the first X-ray imaging unit 112A includes a first X-ray imaging unit 114A and a first 2D image data output unit 116A.

  Each pixel constituting the first X-ray imaging unit 114A is converted into an electric signal having a value corresponding to the intensity of the pulse X-ray and is output. The first 2D image data output unit 116A converts the electrical signal input from the first X-ray imaging unit 114A into image data. This image data is two-dimensional.

Here, the first X-ray imaging unit 112A includes, for example, a color image intensifier (color II (registered trademark)) and a color camera. The color image intensifier uses Y ₂ O ₂ S (Eu), which is a color scintillator, or the like as an output phosphor. The fluorescence decay time of Y ₂ O ₂ S (Eu) is approximately 1 millisecond.

  Further, an image intensifier (II) may be used for the first X-ray imaging unit 112A. Alternatively, an indirect conversion FPD (FPD: Flat Panel Detector) may be used for the first X-ray imaging unit 112A. Since the indirect conversion type FPD is thinner than the color image intensifier, the X-ray moving body tracking device 1 can be downsized.

  In the image intensifier, CsI (Na) or CsI (Tl) is used as a phosphor. The fluorescence decay time, which is the time from emission of light emitted by X-ray emission to disappearance, is approximately 1050 nanoseconds for CsI (Na). Moreover, in CsI (Na), it is about 630 nanoseconds. The indirect conversion FPD also uses CsI (Na) or CsI (Tl) as a phosphor, but the charge readout speed of ASP (Amorphous Silicon Photodiode) used in the light receiving element is slow, and image data is read out. It takes several tens of milliseconds. For this reason, when an imaging interval of several milliseconds is required, use of the indirect conversion type FPD is limited at the current technical level.

  The first image processing unit 118A performs arithmetic processing by weighting the plurality of image data input from the first 2D image data output unit 116A. That is, the first image processing unit 118A generates enhanced image data in which the radiotherapy marker region in the subject 10 is enhanced. In addition, the first image processing unit 118A changes the image data processing method according to the mode set by the setting unit 108.

  When the first mode is set, the first image processing unit 118A performs a difference process between the first image data and the second image data. The difference process here is a process for increasing the contrast of the marker area for radiotherapy, for example. This radiotherapy marker is made of, for example, gold (Au). A detailed processing method of the difference processing will also be described later.

  When the second mode is set, the first image processing unit 118A performs an averaging process of the first image data and the second image data. In this case, noise is reduced by 1 / √n and the radiotherapy marker region is emphasized.

  The second high voltage pulse generation unit 100B has the same configuration as the first high voltage pulse generation unit 100A, and generates a third pulse voltage and a fourth pulse voltage having a magnitude different from that of the third pulse voltage. Occurs at predetermined intervals. Further, the second X-ray tube portion 102B has the same configuration as the first X-ray tube portion 102A, and the third pulse X-ray and the fourth pulse voltage are based on the third pulse voltage and the fourth pulse voltage. Pulse X-rays. The second X-ray tube support portion 104B has the same configuration as the first X-ray tube support portion 104A and supports the second X-ray tube portion 102B. Furthermore, the second collimator unit 106B has the same configuration as the first collimator unit 106A, and limits the irradiation range of the second X-rays generated by the second X-ray tube unit 102B. The second X-ray imaging unit 112B has a configuration equivalent to that of the first X-ray imaging unit 112A, and has a pixel value corresponding to the intensity distribution of the pulse X-ray irradiated through the second collimator unit 106B. Convert to image data and output. Further, the apparatus includes a second X-ray imaging unit 114B and a first 2D image data output unit 116A. The second image processing unit 118B has the same configuration as that of the first image processing unit 118A, and performs arithmetic processing by weighting a plurality of image data captured at different times.

  Two sets of X-ray fluoroscopic imaging systems including the X-ray tube sections 102A and 102B and the X-ray imaging sections 112A and 112B are arranged orthogonally via the subject 10. The X-ray tube sections 102A and 102B and the X-ray imaging sections 112A and 112B may be arranged in an upside down manner. Two sets of X-ray fluoroscopic systems are tilted by 90 ° to irradiate X-rays from the abdomen side and back side. You may comprise.

  The synchronization control unit 120 controls to synchronize the generation timing of the high voltage pulse in the first high voltage pulse generation unit 100A and the second high voltage pulse generation unit 100B. Further, the synchronization control unit 120 performs control to synchronize the imaging timing of the first X-ray imaging unit 114A and the second X-ray imaging unit 114B with the generation timing of the high voltage pulse.

  The three-dimensional conversion unit 122 synthesizes each two-dimensional image data output from the first image processing unit 118A and the second image processing unit 118B to generate three-dimensional image data. The target coordinate output unit 124 detects a radiotherapy marker region from three-dimensional image data based on each two-dimensional image data, and obtains three-dimensional coordinates. Here, the position of the radiotherapy marker is obtained as the position of the affected area target.

  In addition, the target coordinate output unit 124 obtains the first two-dimensional coordinates of the radiotherapy marker based on the two-dimensional image data output from the first image processing unit 118A, and the second image processing unit 118B. The second two-dimensional coordinates of the radiotherapy marker are obtained on the basis of the two-dimensional image data output from, and the radiotherapy marker 3 is obtained based on the first two-dimensional coordinates and the second two-dimensional coordinates. Dimensional coordinates may be obtained. For example, the target coordinate output unit 124 binarizes the two-dimensional image data output from the first image processing unit 118A, and uses the center of gravity of the high pixel region as the first two-dimensional coordinates of the radiotherapy marker. To do. Similarly, the two-dimensional image data output from the second image processing unit 118B is binarized, and the center of gravity of the high pixel region is set as the second two-dimensional coordinate of the radiotherapy marker.

  The irradiation permission determination unit 126 determines whether or not irradiation of the therapeutic beam should be permitted based on the three-dimensional coordinates of the affected area target. That is, the irradiation permission determination unit 126 determines whether or not the radiotherapy marker is located within the gate that is the irradiation range of the therapeutic beam. In the present embodiment, the high voltage pulse generators 100A and 100B constitute a voltage generator, the X-ray tube parts 102A and 102B constitute an X-ray generator, and the target coordinate output part 124 serves as a detector. It is composed.

  Next, the detailed configuration of the first high-voltage pulse generator 100A will be described based on FIGS. 2 and 3. FIG. There are mainly three types of X-ray tubes: anode grounding, cathode grounding, and midpoint grounding. Here, the first high voltage pulse generator 100A for the cathode grounding system and the first high voltage pulse generator 100A for the midpoint grounding system will be described. First, the detailed configuration of the first high voltage pulse generator 100A for the cathode grounding system will be described with reference to FIG. FIG. 2 is a diagram showing a detailed configuration of the first high-voltage pulse generator 100A for the cathode grounding method.

  As shown in FIG. 2, the first high voltage pulse generator 100A includes a control circuit 200, a DC (DC: Direct Current) power source 202, an inverter circuit 204, a booster circuit 205, an inverter circuit 206, and a booster. The circuit 207, the high voltage semiconductor switch 208, and the filament heating power supply 212 are provided.

  The control circuit 200 is connected to the setting unit 108, the synchronization control unit 120, the inverter circuit 204, the booster circuit 205, the inverter circuit 206, the booster circuit 207, the filament heating power supply 212, and the high voltage semiconductor switch 208. Has been. The control circuit 200 controls the inverter circuits 204 and 206, the booster circuits 205 and 207, the high voltage semiconductor switch 208, and the filament heating power supply 212 in accordance with the synchronization control from the synchronization control unit 120.

  The DC power source 202 is a circuit that converts an AC voltage into a DC voltage. The AC voltage is obtained from a commercial power source, for example. The DC power source 202 is, for example, an AC / DC conversion circuit, rectifies an AC voltage with a rectifier diode, and converts it into a DC voltage set using a capacitor.

  The inverter circuit 204 is connected to the control circuit 200, the DC power source 202, and the booster circuit 205. The inverter circuit 204 generates a square voltage based on the generation condition set by the setting unit 108 under the control of the control circuit 200. This rectangular voltage is generated at a predetermined cycle based on the direct current voltage supplied from the DC power source 202.

  The booster circuit 205 is connected to the control circuit 200, the inverter circuit 204, and the anode of the first X-ray tube portion 102A. The booster circuit 205 boosts the voltage input from the inverter circuit 204 under the control of the control circuit 200 and outputs the boosted voltage as a Base voltage. This Base voltage is applied to the anode of the first X-ray tube portion 102A.

  The inverter circuit 206 has the same configuration as the inverter circuit 205. That is, the inverter circuit 206 is connected to the control circuit 200, the DC power source 202, and the booster circuit 207. The inverter circuit 206 generates a square voltage based on the generation condition set by the setting unit 108 under the control of the control circuit 200. This rectangular voltage is generated at a predetermined cycle based on the direct current voltage supplied from the DC power source 202.

  The booster circuit 207 has the same configuration as the booster circuit 205. That is, the control circuit 200, the inverter circuit 206, and the anode of the first X-ray tube portion 102A are connected. The booster circuit 207 boosts the voltage input from the inverter circuit 206 under the control of the control circuit 200 and outputs the boosted voltage as a Boost voltage.

  The Boost voltage is applied to the anode of the first X-ray tube section 102A via the high voltage semiconductor switch 208. The booster circuits 205 and 207 are, for example, CW (CW: Cockcroft Walton) circuits, and are configured by combining capacitors and diodes in a multistage manner.

  The high voltage semiconductor switch 208 is connected to the control circuit 200 and is disposed between the booster circuit 207 and the anode of the first X-ray tube section 102A. The high-voltage semiconductor switch 208 brings the booster circuit 206 and the anode of the first X-ray tube portion 102A into a connected state or a non-connected state according to the control of the control circuit 200. Further, the high voltage semiconductor switch 208 is switched at the above-described predetermined photographing cycle. Thereby, the Post voltage is applied to the anode of the first X-ray tube portion 102A at the predetermined imaging cycle described above.

  Furthermore, the high-voltage semiconductor switch 208 can be configured by connecting a plurality of semiconductor switching elements such as IGBT (IGBT: Insulated Gate Bipolar Transistor) and MOSFET (MOSFET: Metal-Oxide-Semiconductor Field-Effect Transistor). .

  The filament heating power supply 212 is connected to the filament in the first X-ray tube portion 102A. The filament heating power supply 212 heats the filament itself by applying a voltage of about 10 V in order to pass a specified current through the filament. The number of electrons generated from the filament is adjusted according to the current value. That is, the current flowing through the filament corresponds to the tube current of the first X-ray tube portion 102A. In the present embodiment, the Base voltage corresponds to the first pulse voltage, and the Boost voltage corresponds to the second pulse voltage.

  Next, the detailed configuration of the first high-voltage pulse generator 100A for the midpoint grounding method will be described with reference to FIG. Since the first X-ray tube portion 102A of the first high-voltage pulse generator 100A shown in FIG. 2 is a fixed cathode type, the current is limited. As a result, if the SID (SID: Source Image receptor Distance) is long, the necessary tube current may not be obtained. For this reason, when a large tube current is required, an anode rotation type X-ray tube is used. Many of these anode-rotating X-ray tubes are of the midpoint grounding system. That is, the first high voltage pulse generator 100A of the midpoint grounding method is used when a large tube current is required.

  FIG. 3 is a diagram showing a detailed configuration of the first high-voltage pulse generator 100A for the midpoint grounding method. As shown in FIG. 3, the first high voltage pulse generator 100A includes a control circuit 214, a DC power source 202, an inverter circuit 204, a booster circuit 205, an inverter circuit 206, a booster circuit 207, A high voltage semiconductor switch 208, a filament heating power source 212, and a high voltage semiconductor switch 222 are provided. The same components as those in FIG. As shown in FIG. 3, a negative (−) high voltage is also applied to the cathode, which is different from the first high voltage pulse generator 100A shown in FIG.

  The control circuit 214 is connected to the synchronization control unit 120, inverter circuits 204 and 206, booster circuits 205 and 207, high voltage semiconductor switches 208 and 222, and a filament heating power supply 212. The control circuit 214 controls the inverter circuits 204 and 206, the booster circuits 205 and 207, the high voltage semiconductor switches 208 and 222, and the filament heating power supply 212 in accordance with the synchronization control of the synchronization control unit 120.

  The inverter circuits 204 and 206 are connected to the control circuit 214 and the DC power source 202. The inverter circuits 204 and 206 generate a square voltage under the control of the control circuit 214. This rectangular voltage is generated at the above-described predetermined period based on the DC voltage supplied from the DC power source 202.

  The booster circuit 205 is connected to the inverter circuit 204, the control circuit 214, and the cathode of the first X-ray tube unit 102A. The booster circuit 205 boosts the square voltage supplied from the inverter circuit 204 via the high-voltage semiconductor switch 222 in accordance with the control of the control circuit 214, and generates a negative (−) Base voltage and a negative (−) Boost. Output as voltage. The negative (−) Base voltage and the negative (−) Boost voltage are applied to the cathode of the first X-ray tube section 102A.

  The booster circuit 207 is connected to the inverter circuit 206, the control circuit 214, and the anode of the first X-ray tube unit 102A. The booster circuit 207 boosts the square voltage supplied from the inverter circuit 206 via the high-voltage semiconductor switch 208 in accordance with the control of the control circuit 214, so that the positive (+) Base voltage and the positive (+) Boost voltage are boosted. Output as voltage. The positive (+) Base voltage and the positive (+) Boost voltage are applied to the anode of the first X-ray tube portion 102A.

  The high voltage semiconductor switch 208 is connected to the control circuit 214 and is disposed between the inverter circuit 206 and the booster circuit 207. The high voltage semiconductor switch 208 brings the inverter circuit 206 and the booster circuit 207 into a connected state or a disconnected state according to the control of the control circuit 214.

  The high voltage semiconductor switch 222 is connected to the control circuit 214 and is disposed between the inverter circuit 204 and the booster circuit 205. The high voltage semiconductor switch 222 puts the inverter circuit 204 and the booster circuit 205 into a connected state or a disconnected state according to the control of the control circuit 214.

  The above is the description of the configuration of the X-ray moving body tracking apparatus 1 according to the present embodiment. Next, based on FIG. 4 and FIG. 5, the operation for each mode in the first high-voltage pulse generator 100A of the cathode grounding method. An example will be described. As described above, the setting unit 108 can set the first mode and the second mode.

  First, an operation example of the first high-voltage pulse generator 100A for the cathode grounding method in the first mode will be described with reference to FIG. FIG. 4 is a timing chart of the first high-voltage pulse generator 100A for the cathode grounding method in the first mode. (a) The figure has shown the high voltage pulse which the booster circuit 207 generate | occur | produces. Here, the horizontal axis represents elapsed time. The vertical axis indicates the voltage value of the generated pulse, that is, the tube voltage.

FIG. 4B shows the timing when the high voltage semiconductor switch 208 is connected. During the period indicated as ON, the high-voltage semiconductor switch 208 is connected.

  (c) The figure shows the timing when the filament heating power supply 212 sends a specified current to the filament. Current is flowing during the period indicated as ON.

  FIG. 4D shows the timing at which the first X-ray tube unit 102 generates pulse X-rays. During the period indicated as high voltage L, high voltage L pulse X-rays corresponding to the Base voltage are generated. Further, during the period indicated as high voltage H, high voltage H pulse X-rays corresponding to the Post voltage are generated.

  As shown in FIG. 4 (c), the filament heating power supply 212 passes a specified current through the filament for a pulse width of 1 ms. At this timing, the Base voltage is applied to the anode of the first X-ray tube section 102A, and a high-voltage L pulse X-ray is generated.

  Subsequently, when 3 ms elapses after the Base voltage is applied, the high-voltage semiconductor switch 208 connects the booster circuit 207 and the anode of the first X-ray tube unit 102A for 1 ms. As a result, the booster circuit 207 applies a Boost voltage having a pulse width of 1 ms to the anode of the first X-ray tube unit 102A. During this pulse width of 1 ms, a prescribed current is passed through the filament heating power supply 212 filament, and high-voltage H pulse X-rays are generated.

  When another 28 ms elapses, the booster circuit 205 applies a Base voltage having a pulse width of 1 ms to the anode of the first X-ray tube portion 102A. In this way, high-pressure L and H pulse X-rays are repeatedly generated from the first X-ray tube portion 102A at intervals of 33 ms.

  Next, an operation example of the first high-voltage pulse generator 100A for the cathode grounding method in the second mode will be described with reference to FIG. FIG. 5 is a diagram illustrating a timing chart of the first high-voltage pulse generator 100A for the cathode grounding method in the second mode. The operation example shown in FIG. 5 is different from the operation example shown in FIG. 4 in that the Base voltage and the Boost voltage are equivalent values. In the second mode, the booster circuit 205 applies a Base voltage having a pulse width of 1 ms to the anode of the first X-ray tube unit 102A. When another 3 ms elapses, the booster circuit 207 applies a Boost voltage having a pulse width of 1 ms, which is a voltage equivalent to the Base voltage, to the anode of the first X-ray tube portion 102A. Thereby, pulsed X-rays of high pressure L are emitted at intervals of 3 ms. The other operations are the same as those in FIG.

  Next, an example of the operation in the first mode in the first high-voltage pulse generator 100A for the midpoint grounding method will be described with reference to FIG. FIG. 6 is a diagram illustrating a timing chart of the first high-voltage pulse generator 100A for the midpoint grounding method.

  (A) The figure has shown the positive (+) Base voltage and positive (+) Boost voltage which the booster circuit 207 generate | occur | produces. Here, the horizontal axis represents elapsed time. The vertical axis indicates the voltage value of the generated pulse. Similarly, FIG. 5B shows the negative (−) Base voltage and the negative (−) Boost voltage generated by the booster circuit 205.

  (C) The figure shows the timing at which the high voltage semiconductor switch 208 is connected. During the period indicated as ON, the high-voltage semiconductor switch 208 is connected. Similarly, (d) shows the timing at which the high-voltage semiconductor switch 222 is connected. During the period indicated as ON, the high-voltage semiconductor switch 222 is connected.

  (E) The figure has shown the timing which the filament heating power supply 212 sends a regular electric current through a filament. Current is flowing during the period indicated as ON.

  (F) The figure shows the timing at which the first X-ray tube unit 102 generates X-rays. During the period indicated as high voltage L, high voltage L pulse X-rays corresponding to the Base voltage are generated. Further, during the period indicated as high voltage H, high voltage H pulse X-rays corresponding to the Post voltage are generated.

  As shown in FIG. 6, the booster circuits 205 and 207 apply a Base voltage having a pulse width of 1 ms to the cathode and the anode of the first X-ray tube section 102A, respectively. Further, at this timing, the high voltage semiconductor switches 208 and 222 are connected, and the filament heating power supply 212 causes a specified current to flow through the filament. Further, the first X-ray tube unit 102 generates a high-voltage L pulse X-ray corresponding to the Base voltage during a period in which the X-ray is indicated as a high voltage L.

  Subsequently, when 3 ms elapses after the Base voltage is applied, the booster circuits 204 and 207 apply the Boost voltage having a pulse width of 1 ms to the cathode and the anode of the first X-ray tube unit 102A, respectively. Further, at this timing, the high voltage semiconductor switches 208 and 222 are connected, and the filament heating power supply 212 causes a specified current to flow through the filament. Further, the first X-ray tube unit 102 generates a high-voltage H pulse X-ray corresponding to the Post voltage during a period when the X-ray is indicated as a high voltage H.

  When 28 ms elapses, the booster circuits 205 and 207 apply a Base voltage having a pulse width of 1 ms to the first X-ray tube unit 102A. In this way, the Base voltage and the Boost voltage are repeatedly applied to the first X-ray tube unit 102A at intervals of 33 ms, and high-voltage L and H pulse X-rays are generated.

Next, an example of the operation of the first X-ray tube unit 102A and the X-ray imaging units 114A and 114B will be described with reference to FIGS. Here, the case where the high voltage pulse generated by the cathode-grounded first high voltage pulse generator 100A shown in FIG. 2 is applied to the anode of the first X-ray tube unit 102A will be described. First, FIG. Based on this, an operation example of the first X-ray tube unit 102A and the first X-ray imaging unit 114A in the first mode will be described. FIG. 7 is a timing chart of the first X-ray tube unit 102A and the first X-ray imaging unit 114A in the first mode. Here, a case where the Base voltage and the Boost voltage illustrated in FIG. 4 are applied to the anode of the first X-ray tube portion 102A will be described. The upper diagram shows pulse X-rays generated by the first X-ray tube unit 102A. Here, the horizontal axis represents elapsed time. The vertical axis indicates the magnitude of the tube voltage applied to generate the generated pulse X-ray.

  The lower diagram shows the timing at which the first X-ray imaging unit 114A takes an image. The first image data is captured during the period indicated as the first image data. Similarly, the second image data is captured during the period indicated as the second image data.

  As shown in FIG. 7, when the Base voltage is applied, the first X-ray tube unit 102A generates a first pulse X-ray (pulse X-ray with a high voltage L) having a pulse width of 1 ms. Subsequently, when 3 ms elapses after the first pulse X-ray is generated, the Boost voltage is applied. As a result, a second pulse X-ray (pulse X-ray of high voltage H) having a pulse width of 1 ms is generated. Further, when 28 ms elapses after the second pulse X-ray is generated, the first pulse X-ray is generated again.

  The first X-ray imaging unit 114A captures the first image data and the second image data in synchronization with the pulse X-ray generated by the first X-ray tube unit 102A. In this way, the first pulse X-ray and the second pulse X-ray are repeatedly generated at an interval of 33 ms.

  Next, based on FIG. 8, an operation example of the first X-ray tube unit 102A and the first X-ray imaging unit 114A in the second mode will be described. FIG. 8 is a diagram illustrating a timing chart of the first X-ray tube unit 102A and the first X-ray imaging unit 114A in the second mode. Here, a case where the Base voltage and the Boost voltage exemplified in FIG. 5 are applied to the anode of the first X-ray tube portion 102A will be described.

  Similar to FIG. 7, the upper diagram shows the pulse X-rays generated by the first X-ray tube portion 102A. The lower diagram shows the timing at which the first X-ray imaging unit 114A takes an image. The first pulse X-ray (the high-voltage L pulse X-ray) and the second pulse X-ray (the high-voltage H pulse X-ray) have the same size, which is different from the operation example of FIG.

  As shown in FIG. 8, when 28 ms elapses after generating the first pulse X-ray and the second pulse X-ray, the first pulse X-ray is generated again. The first X-ray imaging unit 114A captures the first image data and the second image data in synchronization with the pulse X-ray generated by the first X-ray tube unit 102A. In this way, the first pulse X-ray and the second pulse X-ray are repeatedly generated at an interval of 33 ms.

  When a color image intensifier is used, the pulse width of the pulse X-ray may be set to 1 millisecond, which is the same as the fluorescence decay time of the output phosphor. Alternatively, the pulse width of the pulse X-ray may be set to 2 to 3 milliseconds, which is twice or more the fluorescence decay time of the output phosphor.

  Next, the difference processing in the first image processing unit 118A will be described based on FIG. FIG. 9 is a diagram for explaining the difference processing. Here, for example, as shown in FIG. 7, low tube voltage image data (L) imaged with a first pulse X-ray (high-voltage L pulse X-ray) and a second pulse X-ray (high-voltage H pulse X). A difference processing example using the high tube voltage image data (H) imaged by (line) will be described. The diagram shown in FIG. 9A shows representative pixel values for each region of the image data (L). The horizontal axis represents the area. The vertical axis represents the pixel value. Here, au represents a radiotherapy marker, for example, a radiotherapy marker region formed of gold, b represents a bone region, and S represents a skin region. Similarly, the diagram shown in FIG. 9B shows representative pixel values for each region of the image data (H). The figure shown in FIG. 9C shows representative values of pixel values for each region in the difference image data (D) after the difference processing.

  As described above, when different pulse voltages are applied to the first X-ray tube portion 102A, X-rays having different characteristics are generated. For this reason, pixel value distributions of image data captured with X-rays having different characteristics are also different. As a result, the ratio between the representative pixel value of the radiotherapy marker region in FIG. 9A and the representative pixel value of the bone region becomes the representative pixel value of the radiotherapy marker region in FIG. The ratio with the representative pixel value of the partial area is different. For example, the ratio in FIG. 9A is 300/120 = 2.5, and the ratio in FIG. 9B is 200/96 = 2.08.

  The difference processing here uses such characteristics to obtain difference image data (D) in which the radiotherapy marker region is emphasized. This difference processing is used, for example, when the radiotherapy marker region is imaged so as to overlap the bone region.

  Such a difference process is arithmetically processed according to, for example, an expression D (x, y) = (L (x, y) −K1 × H (x, y)) × K2. Here, the image data (L) is indicated by L (x, y), the image data (H) is indicated by H (x, y), and the difference image data (D) is indicated by D (x, y). Further, (x, y) represents the coordinates of the image data, K1 represents a weighting coefficient, and K2 represents a coefficient for setting the pixel value range of the difference image data (D) to a predetermined range. In some cases, this equation is simplified and expressed by the equation D = (L−K1 × H) × K2.

  More specifically, first, when the image data (H) is subtracted from the image data (L), a weighting coefficient K1 for obtaining a representative pixel value of 0 in the bone region is obtained. Here, the representative pixel value of the bone region of the image data (L) is 120, and the representative pixel value of the bone region of the image data (H) is 96. Therefore, by dividing 120 by 96, the weighting coefficient K1 = 1.25 can be obtained. In addition, as the representative pixel value of the bone region here, for example, an average value of pixel values in the bone region extracted by region extraction is used.

  Next, the weighting coefficient K1 = 1.25 is multiplied by the image data (H) and subtracted from the image data (L). That is, a weighted difference process for setting the representative pixel value of the bone region of the image data (L) and the representative pixel value of the bone region of the image data (H) to 0 is performed on the image data (L) and the image data (H). To do.

  Next, a coefficient K2 is obtained for setting the pixel value range of the difference image data (D) obtained by the difference process to a predetermined range. For example, when the pixel value range of the difference image data (D) is 0 to 50, the coefficient K2 for obtaining the pixel value range 0 to 300 is obtained. In this case, the maximum value 300 in the pixel value range 0 to 300 is divided by the maximum value 50 of the difference image data (D) to obtain a coefficient K2 = 6.0.

  Then, the difference image data (D) is multiplied by the coefficient K2 by the difference process. Thereby, the calculation which makes the pixel width of difference image data (D) the predetermined pixel value range is performed. Note that the difference image data (D) may be multiplied by a coefficient K2. Alternatively, the difference image data (D) may not be multiplied by the coefficient coefficient K2. By multiplying the coefficient K2, the contrast of the radiotherapy marker region is enhanced. On the other hand, even when the coefficient coefficient K2 is not multiplied, the contrast of the radiotherapy marker region is relatively improved. Here, both the image data when multiplied by the coefficient coefficient K2 and the image data when not multiplied by the coefficient coefficient K2 are referred to as difference image data (D). In this way, the calculation of D = (L−1.25 × H) × 6.0 is performed here.

  Thereby, difference image data (D) can be obtained. As shown in FIG. 9A, generally, since the radiotherapy marker region is a metal, it does not transmit X-rays and has a higher pixel value than the bone region. For this reason, as shown in FIG. 9C, the representative pixel value of the radiotherapy marker region does not become 0 by this difference processing. On the other hand, the representative pixel value in the bone region is zero.

  As can be seen from these facts, the contrast of the bone region is reduced more than before the difference processing. As a result, the contrast of the radiotherapy marker region is relatively improved, so that the discrimination of the radiotherapy marker region is improved.

  Further, the image data here becomes a high pixel as the X-ray irradiation intensity decreases. On the other hand, it is possible to perform the same processing on the image data that is in the negative / positive inversion state, that is, the image data that has a lower pixel value as the X-ray irradiation intensity decreases.

  As described above, by using the representative pixel value of the bone region, an image in which the radiotherapy marker is emphasized can be obtained. Note that the representative pixel value in the bone region may be obtained by histogram analysis of the pixel value in the subject region.

  Next, an example of difference processing for each color image data when a color image intensifier is used for the imaging units 112A and 112B will be described. First, the X-ray sensitivity for each color image data in the color image intensifier will be described.

When the output phosphor Y ₂ O ₂ S (Eu) used in the color image intensifier is irradiated with X-rays having the same energy and intensity, the red region emits the strongest light, and then the light is emitted in the order of green and blue. become weak. The phosphor Y ₂ O ₂ S (Eu) can adjust the light emission amount of the red region, that is, the R component, the green region, that is, the G component, and the blue region, that is, the B component, by changing the concentration of the Eu component. For this reason, the components of the phosphor Y ₂ O ₂ S (Eu) are adjusted in accordance with the sensitivity characteristics of the R, G, and B pixels of the CCD imaging device used in the color image intensifier.

That is, the light emission of the output phosphor Y ₂ O ₂ S (Eu) is captured by a color camera having an R pixel that captures red, a G pixel that captures green, and a B pixel that captures blue. As a result, the signals are converted into R, G, and B signals having different sensitivities in accordance with the intensity of the incident X-rays. That is, for X-rays, the R pixel has the highest sensitivity, and the sensitivity decreases in the order of the G pixel and the B pixel. In other words, red image data captured by the R pixel has the highest sensitivity to X-rays, and the sensitivity decreases in the order of green image data captured by the G pixel and blue image data captured by the B pixel. Here, high sensitivity means that the pixel value becomes high for the same X-ray intensity. Or it means that the S / N ratio becomes high for the same X-ray intensity.

  In this way, three types of image data with different sensitivities are obtained. Note that by combining these image data, it is possible to obtain image data with an expanded dynamic range.

  Next, a difference processing example for each color image data will be described. Monochrome cameras have only one characteristic curve as shades of white to black. On the other hand, in the case of a color camera, it is possible to give each of R, G, and B of the three primary colors light and dark characteristic curves. For this reason, it is possible to obtain image data after differential processing having different characteristics for each of R image data obtained by R pixels, G image data obtained by G pixels, and B image data obtained by B pixels. It is.

  Difference processing using high-sensitivity R image data is performed on an area where the X-ray dose transmitted through the X-ray in the subject is small. For example, a difference process using R image data is performed on a region where a radiotherapy marker region overlaps a bone part. Thereby, it is possible to improve the contrast of the radiotherapy marker region as compared with the case of using monochrome image data.

  On the other hand, a difference process using low-sensitivity B image data is performed on an area where the X-ray dose transmitted through the subject is large. For example, a difference process using B image data is performed on an area where the radiotherapy marker area overlaps the skin area. Thereby, it is possible to improve the contrast of the radiotherapy marker region as compared with the case of using monochrome image data.

  Further, for example, a difference process using G image data is performed on a region where the radiotherapy marker region overlaps the organ part. Thereby, it is possible to improve the contrast of the radiotherapy marker region as compared with the case of using monochrome image data. Thus, the image data used for the difference process is changed according to the X-ray absorption amount of the region in the subject where the radiation treatment marker region is arranged. Thereby, it is possible to obtain image data by further improving the contrast of the radiotherapy marker region.

  As can be seen from these facts, when treating an affected part target that undergoes respiratory movement, enhanced image data after differential processing that suppresses motion artifacts is obtained by imaging with X-rays of different tube voltages within a predetermined time. be able to. For this reason, it is possible to improve the contrast of the radiotherapy marker region, which has been difficult to identify conventionally. Thereby, the X-ray moving body tracking apparatus 1 which improved the identification accuracy of the marker for radiotherapy can be implement | achieved.

  According to this embodiment, based on the first pulse voltage and the second pulse voltage having a magnitude different from that of the first pulse voltage, a plurality of pulse X-rays having different tube voltages are supplied to the X-ray tube portions 102A and 102B. The image processing units 118A and 118B are caused to generate enhanced image data that enhances the contrast of the radiotherapy marker region in the subject 10 using a plurality of X-ray image data that are irradiated and have different tube voltages. For this reason, the radiotherapy marker region can be detected more accurately. Thereby, the position of the tracking target in the subject 10 can be detected with higher accuracy.

(Second Embodiment)
The X-ray moving body tracking apparatus according to the second embodiment attempts to reduce the integrated amount of X-rays irradiated to the subject by changing the imaging cycle based on the respiratory waveform. Differences from the first embodiment will be described below.

  Next, an X-ray moving body tracking apparatus according to the second embodiment will be described with reference to FIG. FIG. 10 is a block diagram showing the configuration of the X-ray moving body tracking apparatus according to the second embodiment. As shown in FIG. 10, the respiration monitor unit 300, the respiration waveform generation unit 302, and the imaging cycle setting unit 304 are provided, which is different from the X-ray moving body tracking apparatus according to the first embodiment. The same code | symbol is attached | subjected to the structure same as the X-ray moving body tracking device 1 which concerns on 1st Embodiment, and the overlapping description is abbreviate | omitted.

  The respiration monitor unit 300 acquires a measurement signal related to the respiration waveform from the subject 10. This measurement signal indicates, for example, the height of the abdomen. The respiration monitor unit 300 can use a non-contact sensor and a contact sensor. For example, as the non-contact type sensor, an infrared type, an ultrasonic type, a radio wave type, a laser type, or the like can be used for the respiration monitor unit 300.

  The respiration waveform generation unit 302 generates a respiration waveform based on the abdominal height information acquired by the respiration monitor unit 300. The height of the abdomen measured by the respiration monitor unit 300 and the time-series change in the amount of air in the lung have a high correlation. For this reason, the respiration waveform generation unit 302 generates a respiration waveform using the relationship between the time-series change in the abdominal height and the time-series change in the air volume in the lung field obtained in advance. For example, the respiration waveform generation unit 302 generates in advance a function that uses the height of the abdomen during the inspiration period as an input value and uses the air volume in the lung field as an output value. Similarly, a function having the abdominal height during the expiration period as an input value and the air volume in the lung field as an output value is generated in advance. Thereby, the respiration waveform generation unit 302 converts the height of the abdomen into the amount of air in the lung field using these functions, and generates a respiration waveform.

  The imaging cycle setting unit 304 changes the imaging cycle based on the respiratory waveform generated by the respiratory waveform generation unit 302. The imaging cycle setting unit 304 generates an analysis signal when the subject 10 is in a predetermined breathing state. For example, the analysis signal is generated when the value of the respiratory waveform is equal to or less than a predetermined threshold value.

  Then, the shooting cycle setting unit 304 sets the shooting cycle based on the analysis signal. That is, during the period in which the analysis signal is generated, since the radiotherapy marker is in the gate or close to the gate, the imaging cycle for irradiating pulsed X-rays is set to the first imaging cycle. For example, the first imaging cycle is 33 milliseconds.

  On the other hand, when no analysis signal is input, the radiotherapy marker is located away from the gate, so the imaging cycle is set to the second imaging cycle. This second imaging cycle is set longer than the first imaging cycle. For example, the second imaging cycle is 100 milliseconds.

  Thus, the imaging cycle is lengthened at the position where the radiotherapy marker is away from the gate. That is, since the irradiation frequency of X-ray pulses is reduced, the integrated amount of X-rays irradiated to the subject 10 can be reduced.

  According to this embodiment, the imaging cycle setting unit 304 changes the imaging cycle based on the respiratory waveform generated by the respiratory waveform generation unit 302. Thereby, the integrated amount of X-rays irradiated to the subject 10 can be reduced.

(Third embodiment)
In the second embodiment described above, the three-dimensional conversion unit 122 detects the three-dimensional position of the tracking target based on the plurality of image data generated by the image processing units 118A and 118B. In the embodiment, the two-dimensional position of the tracking target is to be detected based on one image data. Hereinafter, a different part from 2nd Embodiment mentioned above is demonstrated.

  FIG. 11 is a block diagram illustrating the configuration of the X-ray moving body tracking apparatus 1 according to the third embodiment. As shown in FIG. 11, the X-ray moving body tracking apparatus according to this embodiment includes an X-ray tube support section 104A and an X-ray imaging section 114A as compared with the X-ray moving body tracking apparatus according to the second embodiment. Only a pair of photographing systems is provided, and the display unit 400 is further provided. The same code | symbol is attached | subjected to the structure same as the X-ray moving body tracking device 1 which concerns on 1st Embodiment, and the overlapping description is abbreviate | omitted.

In the present embodiment, the target coordinate output unit 124 detects the position of the additional radiation treatment marker in the subject 10 based on the two-dimensional enhanced image data output from the image processing unit 118A.

  The display unit 400 sequentially displays the two-dimensional difference image data output from the image processing unit 118A, and displays a display form indicating the position of the tracking target based on the two-dimensional position sequentially obtained by the target coordinate output unit 124. . This display part is comprised with the monitor, for example. Reference numeral 402 denotes an image based on the two-dimensional difference image data output from the image processing unit 118A. Reference numeral 404 denotes an example of a display form that indicates the position of the tracking target based on the two-dimensional position obtained by the target coordinate output unit 124. Here, a range indicated by a dotted line indicates a gate that is an irradiation range of the therapeutic beam, and a shadow in the gate indicates a tumor shadow.

  Thus, since the display form indicating the position of the tracking target is displayed together with the emphasized image data based on the two-dimensional position obtained by the target coordinate output unit 124, other isolated shadows and the additional radiation therapy marker that is the tracking target are displayed. Identification can be performed easily.

  As described above, according to the X-ray moving body tracking apparatus according to the present embodiment, the target coordinate output unit 124 detects the two-dimensional position of the additional radiation treatment marker based on the enhanced image data generated by the image processing unit 118A. It was decided to. For this reason, the position of the marker for additional radiation therapy can be acquired more accurately.

  Although several embodiments have been described above, these embodiments are presented as examples only and are not intended to limit the scope of the invention. The novel apparatus, methods and systems described herein can be implemented in various other forms. In addition, various omissions, substitutions, and changes can be made to the forms of the devices, methods, and systems described in the present specification without departing from the spirit of the invention. The appended claims and their equivalents are intended to include such forms and modifications as fall within the scope and spirit of the invention.

1: X-ray moving body tracking apparatus, 100A: first high-voltage pulse generator, 100B: second high-voltage pulse generator, 108: setting unit, 112A: first X-ray imaging unit, 114B: second X-ray imaging unit, 118A: first image processing unit, 118B: second image processing unit, 122: three-dimensional conversion unit, 124: target coordinate output unit, 126: irradiation permission determination unit, 204: inverter circuit, 205 : Booster circuit, 206: Inverter circuit, 207: Booster circuit, 300: Respiration monitor unit, 302: Respiration waveform generation unit, 304: Imaging cycle setting unit, 400: Display unit

Claims (15)

A voltage generator for generating a first pulse voltage and a second pulse voltage different from the first pulse voltage at a predetermined interval;
An X-ray generator that generates a first pulse X-ray and a second pulse X-ray based on the first pulse voltage and the second pulse voltage;
Imaging first X-ray image data corresponding to the first pulse X-ray transmitted through the subject and second X-ray image data corresponding to the second pulse X-ray transmitted through the subject. An X-ray imaging unit,
An image processing unit that generates enhanced image data in which a radiotherapy marker region in the subject is enhanced using the first X-ray image data and the second X-ray image data;
A detection unit for detecting a position of the radiotherapy marker based on the enhanced image data;
An X-ray moving body tracking device comprising: The pulse voltage generator includes an inverter circuit that converts a supplied voltage into a square voltage;
A voltage boosting circuit that boosts a voltage input from the inverter circuit and outputs the boosted voltage as the first pulse voltage and the second pulse voltage;
The X-ray moving body tracking device according to claim 1, wherein the first pulse voltage and the second pulse voltage are applied to at least one of an anode and a cathode of the X-ray generation unit.   The pulse width of the pulse voltage having the higher voltage out of the first pulse voltage and the second pulse voltage is shorter than the pulse width of the pulse voltage having the lower voltage. 2. The X-ray moving body tracking device according to 2.   4. The pulse width of each of the first pulse voltage and the second pulse voltage is in a range of 0.1 milliseconds to 3 milliseconds. 5. X-ray moving body tracking device. The pulse voltage generator repeatedly generates each of the first pulse voltage and the second pulse voltage at a predetermined generation frequency,
The X-ray moving body tracking device according to any one of claims 1 to 4, wherein the occurrence frequency is in a range of 1 Hz to 60 Hz.   6. The time until the second pulse voltage is generated after the first pulse voltage is generated is set in a range of 1 to 33 milliseconds. The X-ray moving body tracking device according to one item. A setting unit configured to set a voltage value and a pulse width of each of the first pulse voltage and the second pulse voltage according to an imaging region of the subject;
The first mode in which the magnitudes of the first pulse voltage and the second pulse voltage are made different according to the imaging region, and the first pulse voltage and the second pulse voltage have the same magnitude. Set one of the second modes to
When the first mode is set, the image processing unit performs weighted difference processing between the first X image data and the second X image data, and when the second mode is set. The X-ray moving object tracking according to any one of claims 1 to 6, further comprising a setting unit that performs an averaging process of the first X image data and the second X image data. apparatus. The X imaging unit includes a color image intensifier,
A color camera unit that captures emission image data of the color image intensifier;
The X-ray moving body tracking device according to claim 1, further comprising:   The image processing unit performs weighted difference processing based on a pixel value obtained from the bone region of the first X image data and a pixel value obtained from the bone region of the second X image data. The X-ray moving body tracking apparatus according to claim 1, wherein the tracking is performed on the first X image data and the second X image data. The first X-ray image data includes red image data, green image data, and blue image data, and the second X-ray image data includes red image data, green image data, and blue image data. And
The X-ray moving body tracking device according to claim 9, wherein the image processing unit performs weighted difference processing between image data corresponding to a color selected from red, green, and blue. A second voltage generator for generating a third pulse voltage and a fourth pulse voltage having a magnitude different from the third pulse voltage at a predetermined interval;
A second X-ray generator that generates a third pulse X-ray and a fourth pulse X-ray based on the third pulse voltage and the fourth pulse voltage;
The third X-ray image data corresponding to the third X-ray transmitted through the subject and the fourth X-ray image data corresponding to the fourth X-ray transmitted through the subject are captured. Two X-ray imaging units;
A second image processing unit that generates enhanced image data that enhances the contrast of the radiotherapy marker region in the subject using the third X-ray image data and the fourth X-ray image data. When,
Based on the image data obtained by the image processing unit and the image data obtained by the second image processing unit, a three-dimensional conversion unit that three-dimensionalizes the image data;
A detection unit that detects a three-dimensional position of the radiotherapy marker region based on the three-dimensional image data;
The X-ray moving body tracking device according to claim 1, further comprising:   The X-ray moving body tracking apparatus according to claim 1, further comprising a display unit that displays image data obtained by the image processing unit together with a mark indicating a radiation beam irradiation position. A pulse voltage generation step of generating a first pulse voltage and a second pulse voltage having a magnitude different from that of the first pulse voltage at a predetermined interval;
An X-ray generation step of generating a first pulse X-ray and a second pulse X-ray based on the first pulse voltage and the second pulse voltage;
Imaging first X-ray image data corresponding to the first pulse X-ray transmitted through the subject and second X-ray image data corresponding to the second pulse X-ray transmitted through the subject. X-ray imaging process
Using the first X-ray image data and the second X-ray image data, an image data processing step for generating enhanced image data for enhancing the contrast of the radiotherapy marker region in the subject;
A detection step of detecting a position of the radiotherapy marker based on the enhanced image data;
An X-ray moving object tracking method comprising: First image data corresponding to the first X-ray transmitted through the subject, and second X-ray having characteristics different from those of the first X-ray, and the second X-ray transmitted through the subject An acquisition step of acquiring second image data corresponding to the line;
A calculation step of calculating a weighting coefficient based on a pixel value obtained from the bone region of the first image data and a pixel value obtained from the bone region of the second image data;
A difference processing step for performing a difference process between the first X image data and the second X image data using the weighting coefficient;
An image processing method comprising: Each of the first image data and the second image data is any of image data captured by an image sensor having R, G, and B pixels,
The first image data is any one of first red image data, first green image data, and first blue image data, and the second image data is second red image data, 15. The image processing method according to claim 14, wherein the image processing method is one of second green image data and second blue image data, and is image data of a color corresponding to the first image data. .

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