JP5823208B2 - X-ray computed tomography system - Google Patents

Description

  Embodiments described herein relate generally to an X-ray computed tomography apparatus employing a photon counting method.

  An X-ray computed tomography apparatus (hereinafter referred to as an X-ray CT apparatus) rotates an X-ray source around the body axis of the subject and causes the X-ray source to irradiate the subject with X-rays. , Which reconstructs tomographic images based on projection data obtained by detecting X-rays transmitted through a subject with an X-ray detector. It plays an important role in medical practice.

  Conventionally, the number of photons of X-rays detected by an X-ray detector is counted, and a so-called photon counting method is employed that reconstructs an image related to a subject based on the count number and the energy value of each photon. A line CT apparatus is known.

  In this type of X-ray CT apparatus, if the number of X-ray photons incident on the X-ray detector increases, a pile-up phenomenon may occur in which charge pulses corresponding to the photons overlap in time. When the pile-up phenomenon occurs, the energy value of each photon cannot be accurately measured, which causes an artifact in the reconstructed image. In addition, since it is difficult to discriminate each charge pulse, there may be a case where the number of photons is counted.

  On the other hand, when the irradiation dose to the subject is reduced so as not to cause a pile-up phenomenon, photon noise appears in the reconstructed image. Therefore, in order to ensure a sufficient S / N ratio (Signal to Noise ratio), it is necessary to lengthen the imaging time, which is a burden on the subject and the operator.

JP 2006-101926 A

  The problem to be solved by the present invention is to provide an X-ray CT apparatus capable of eliminating or reducing the influence of a pile-up phenomenon on a reconstructed image and obtaining a good reconstructed image.

An X-ray computed tomography apparatus according to an embodiment detects an X-ray generation unit that generates X-rays, and X-rays that are generated by the X-ray generation unit and transmitted through a subject. An X-ray detector that outputs a corresponding signal, a first extraction unit that extracts a high-frequency component of the signal output from the X-ray detector, and a low-frequency component of the signal output from the X-ray detector A second extraction unit that performs measurement, a measurement unit that measures a peak value of the high-frequency component and a peak value of the low-frequency component corresponding to a photon of the X-ray detected by the X-ray detector, and the X per unit time A count rate calculation unit that calculates a count rate indicating the number of photons detected by a line detector; a crest value of the high frequency component and a crest value of the low frequency component measured by the measurement unit; and the count rate calculation unit. Based on the counting rate calculated by Re reconstructing a correction unit for determining the energy value of the photon, and the count unit energy value determined by the correction unit counts the predetermined condition is satisfied the photons, an image based on the count result of the counting unit And a component.

1 is a block diagram showing the overall configuration of an X-ray CT apparatus according to an embodiment. The figure which shows an example of the electric charge read-out circuit used with the conventional X-ray CT apparatus. FIG. 3 is a waveform diagram of a signal at point A or point B in FIG. 2 when no pile-up occurs. FIG. 3 is a waveform diagram of a signal at point A or point B in FIG. 2 when pileup occurs. The figure which shows the electric charge read-out circuit etc. which concern on one Embodiment. FIG. 6 is a waveform diagram of a signal at point C in FIG. 5 when a pileup occurs. FIG. 6 is a waveform diagram of a signal at point D in FIG. 5 when pileup occurs. FIG. 6 is a diagram showing a circuit that follows the subsequent stage of the charge readout circuit shown in FIG. 5. The figure which shows the relationship between the sensitivity of the low frequency component in the circuit shown in FIG.

  Hereinafter, an embodiment will be described with reference to the drawings.

  The X-ray CT system imaging system includes a rotation / rotation (rotate / rotate) type in which an X-ray tube and a detector system are integrally rotated around a subject, and a large number of detection elements in a ring shape. There are various types such as a fixed / rotation type (STATIONARY / ROTATE) type in which only the X-ray tube is rotated around the subject, and the present invention can be applied to any type. Here, a rotation / rotation type X-ray CT apparatus, which currently occupies the mainstream, will be described as an example.

  Also, in order to reconstruct an image, projection data for 360 ° around the subject is required, and projection data for 180 ° + fan angle is also required for the half-scan method. However, it is applicable to this embodiment.

  In recent years, the so-called multi-tube type X-ray CT apparatus in which a plurality of pairs of X-ray tubes and X-ray detectors are mounted on a rotating ring has been commercialized, and the development of peripheral technologies has been advanced. In the present embodiment, both a conventional single-tube X-ray CT apparatus and a multi-tube X-ray CT apparatus are applicable. Here, a single tube type will be described.

  In the following description, components having substantially the same function and configuration are denoted by the same reference numerals, and redundant description will be given only when necessary.

[Overall configuration of X-ray CT apparatus]
FIG. 1 is a block diagram showing the overall configuration of an X-ray CT apparatus 1 according to this embodiment. As shown in FIG. 1, the X-ray CT apparatus 1 includes a gantry device A and a console device B.

  The gantry A exposes the subject to X-rays, detects the X-rays transmitted through the subject, and acquires projection data.

  As shown in FIG. 1, the gantry device A includes a fixed unit 10, a rotating unit 11, a bed 12, an X-ray tube 13 (X-ray generation unit), an X-ray detector 14, a data collection unit 15, and a data transmission unit 16. , The gantry bed driving unit 17, the power feeding unit 18, the high voltage generating unit 19, and the like.

  The X-ray tube 13 is a vacuum tube that generates X-rays, and is provided in the rotating unit 11. The X-ray detector 14 is a detector system that detects X-rays transmitted through the subject P, and is attached to the rotating unit 11 in a direction facing the X-ray tube 13.

  The rotating portion 11 is provided with an opening 110, and the bed 12 is disposed in the opening 110. A subject P is placed on the slide top of the bed 12. The gantry bed driving unit 17 moves the bed 12 in the body axis direction while rotating the rotating unit 11 at a high speed around a central axis parallel to the body axis direction of the subject P inserted into the opening 110. In this way, the subject P is scanned over a wide range.

  The data collection unit 15 includes a charge readout circuit and the like (see FIG. 5 and FIG. 8), which will be described later. Based on a signal output from the X-ray detector 14, the data collection unit 15 transmits the subject P and transmits the X-ray detector. 14 is counted, and the energy value of the counted photon is measured (calculated), and the count result and the measurement result of the energy value are sent to the fixed unit 10 side via the data transmission unit 16 applying optical communication. To transmit.

  The fixing unit 10 is supplied with operating power from an external power source such as a commercial AC power source. The operating power supplied to the fixed unit 10 is transmitted to each part of the rotating unit 11 via a power feeding unit 18 that is, for example, a slip ring.

  The high voltage generator 19 includes a high voltage transformer, a filament heating converter, a rectifier, a high voltage switch, and the like, and converts the operating power supplied from the power supply unit 18 into a high voltage to generate an X-ray tube 13. To supply.

Next, the console apparatus B will be described. The console device B includes a controller 21, a data storage unit 22, an image reconstruction unit 23, an image processing unit 24, an image display unit 25, an operation unit 26, a data / control bus 27, and the like.
The controller 21 includes a CPU (Central Processing Unit), a ROM (Read Only Memory), a RAM (Random Access Memory), and the like, and performs overall control regarding various processes such as photographing processing, data processing, and image processing.

  The data storage unit 22 stores various data such as data (projection data) sent from the data transmission unit 16 and reconstructed image data.

  The image reconstruction unit 23 performs reconstruction processing on the projection data based on predetermined reconstruction parameters, for example, a reconstruction area size, a reconstruction matrix size, a threshold value for extracting a region of interest, and the like, thereby reconstructed image data. Is generated.

  The image processing unit 24 performs image processing for display such as window conversion and RGB processing on the reconstructed image data generated by the image reconstructing unit 23, and outputs the processed image data to the image display unit 25. Further, the image processing unit 24 generates a so-called pseudo three-dimensional image such as a tomographic image of an arbitrary cross section, a projection image from an arbitrary direction, or a three-dimensional surface image based on an instruction from an operator, and outputs the generated image to the image display unit 25. . The output image data is displayed as an X-ray CT image on the image display unit 25.

The operation unit 26 includes a keyboard, various switches, a mouse, and the like, and is used for inputting various scanning conditions such as a slice thickness and the number of slices.
The data / control bus 27 is a signal line for connecting the units and transmitting / receiving various data, control signals, address information, and the like.

[Pile-up phenomenon]
Here, the pile-up phenomenon will be described.
FIG. 2 shows an example of a charge readout circuit used in a conventional X-ray CT apparatus. The charge readout circuit includes, for example, an amplifier (AMP) 101 connected to each detection element 100 of the X-ray detector and an AD converter (ADC) 102 connected to the amplifier 101.

  The detection element 100 outputs a signal according to the charge of the incident photon, this signal is amplified to a predetermined level by the amplifier 101, and the amplified signal is converted into a digital signal indicating a time-series signal amplitude by the AD converter 102. Converted. The signal converted by the AD converter 102 is sent to a counter (not shown). This counter has a wave height discriminator that outputs a specific output pulse when a signal exceeding a predetermined amplitude is input, and by counting the number of pulses output from this wave height discriminator, X The number of photons incident on the line detector is obtained. Further, the energy value (crest value) of each photon counted by this counter is measured based on the signal output from the AD converter 102 and stored in the data storage unit together with the count number.

The waveform of the signal output from the detection element 100 (the signal at point A in the figure) or the signal output from the amplifier 101 (the signal at point B in the figure) changes with time as shown in FIGS. Show.
The waveforms shown in FIG. 3 are those when no pile-up phenomenon occurs, and the outputs corresponding to the photon charges that have passed through the subject and entered the X-ray detector appear in time series without overlapping. ing. On the other hand, the waveforms shown in FIG. 4 are obtained when a pile-up phenomenon occurs, and outputs corresponding to photon charges overlap each other, and each output has a signal value (energy value) higher than the original. Is shown. In particular, the separation between the original energy value and the actually measured value becomes prominent near the center of the superposition.

  When such a pile-up phenomenon occurs, the energy value of photons cannot be accurately measured, and there is a risk that photon counting will be missed, causing artifacts in the reconstructed image.

[Charge readout circuit according to this embodiment]
A charge readout circuit included in the X-ray CT apparatus 1 according to the present embodiment will be described.
FIG. 5 is a circuit diagram showing the X-ray detection element and the charge readout circuit according to the present embodiment. This charge readout circuit is mounted on the data collection unit 15, for example.

  An amplifier 31 is connected to each of a plurality of detection elements 30 included in the X-ray detector 14, and an output line of the amplifier 31 is separated into two paths, one being connected to a differentiation circuit 32 and the other being connected to an integration circuit 33. .

  The differentiation circuit 32 functions as a first extraction unit according to the present embodiment, and its output line is connected to the amplifier 34a, and the output line of the amplifier 34a is connected to the AD converter 35a. On the other hand, the integration circuit 33 functions as a second extraction unit according to the present embodiment, and its output line is connected to the amplifier 34b, and the output line of the amplifier 34b is connected to the AD converter 35b.

  The differentiation circuit 32 includes a capacitor 321 provided on a signal line connecting the amplifier 31 and the amplifier 34a, a resistor 322 having one end connected between the capacitor 321 and the amplifier 34a on the signal line, and the other end grounded. including. When this differentiation circuit 32 is interposed, a high frequency component is extracted from the signal output from the amplifier 31. The extracted signal indicating the high-frequency component is amplified by the amplifier 34a and converted to a digital signal indicating time-series amplitude by the AD converter 35a.

  On the other hand, the integrating circuit 33 has a resistor 331 provided on a signal line connecting the amplifier 31 and the amplifier 34b, and a capacitor having one end connected between the resistor 331 and the amplifier 34b on the signal line and the other end grounded. 332. When this integration circuit 33 is interposed, a low frequency component is extracted from the signal output from the amplifier 31. The extracted signal indicating the low frequency component is amplified by the amplifier 34b and converted into a digital signal indicating time-series amplitude by the AD converter 35b.

  When the piled up signal is output from the detection element 30, the waveform of the signal output from the amplifier 34a (the signal at point C in the figure) shows a time change as shown in FIG. 6, for example, and is output from the amplifier 34b. The waveform of the signal (the signal at point D in the figure) shows a time change as shown in FIG. 7, for example.

  That is, in the signal output from the amplifier 34 a, the output corresponding to the charge of each photon incident on the detection element 30 appears as pulses that do not overlap each other due to the action of the differentiation circuit 32.

  On the other hand, the signal output from the amplifier 34b takes a value obtained by substantially averaging the original energy value of each photon incident on the detection element 30, and reflects the degree of pileup. That is, in the signal output from the detection element 30, the peak value increases as the overlap due to pileup becomes more conspicuous.

[Counting photons and calculating energy values]
A circuit subsequent to the circuit shown in FIG. 5 and processing performed in the circuit will be described with reference to FIG.

  An output from the AD converter 35a (hereinafter referred to as a signal X) and an output from the AD converter 35b (hereinafter referred to as a signal Y) are input to an arithmetic unit 40 configured by, for example, a processor or a memory. The calculation unit 40 executes functions of the measurement unit 41, the count rate calculation unit 42, the correction unit 43, and the like by executing the computer program stored in the memory or the like by the processor.

  The measurement unit 41 measures the peak value of the signal X corresponding to the charge of photons incident on the detection element 30. Specifically, for example, the peak value of the peak of each pulse included in the waveform diagram of FIG. 6 is measured. Each time the measurement unit 41 measures the peak value corresponding to the charge of the photon based on the signal X, the measurement unit 41 outputs the measured peak value to the correction unit 43 and outputs a detection signal to the count rate calculation unit 42.

  Further, the measurement unit 41 measures the peak value of the signal Y corresponding to the charge of the photon incident on the detection element 30. Specifically, with respect to the signal X, the peak value of the signal Y is measured at the timing when the peak value of the peak of each pulse is measured. The measuring unit 41 outputs the measured peak value to the correcting unit 43 every time the peak value corresponding to the charge of the photon is measured based on the signal Y.

  The count rate calculation unit 42 calculates the number of detection signals (photon count rate) input from the measurement unit 41 per unit time immediately before, and outputs the calculated count rate to the correction unit 43.

  The correction unit 43 calculates the energy value of photons incident on the detection element 30 based on the peak value of the signal X, the peak value of the signal Y, and the count rate measured by the measurement unit 41, and counts the calculated energy value. To the unit 44. Specifically, the correction unit 43 stores in advance a calculation formula or table for calculating a correction value (original energy value of photons) using the peak value of the signal X, the peak value of the signal Y, and the count rate as parameters. The original energy value is calculated using these calculation formulas and tables. The calculation formulas and tables may have specific contents so that the correction value is uniquely determined based on the peak value and the count rate of the signals X and Y by a prior experiment or theoretically.

  The count unit 44 includes, for example, a plurality of comparators to which signals from the correction unit 43 are input, and a counter connected to each comparator. Each comparator is set with an energy region that does not overlap with each comparator. Each comparator outputs a pulse to a counter connected to the comparator when a signal in the energy range set for the comparator is input from the correction unit 43. Each counter counts pulses input to itself. The count result of each counter is read by a subsequent count number reading circuit (not shown) and sent to the data storage unit 22 together with the count result regarding the other detection elements 30 provided in the X-ray detector 14.

  In this way, projection data that takes into account the number of photons that have passed through the subject P and the energy value of each photon can be obtained. The image reconstruction unit 23 reconstructs a tomographic image of the subject P based on the projection data. In reconstruction, for example, by using the number of transmitted photons in a specific energy range, it is possible to obtain an image in which a specific substance (for example, bone or the like) is emphasized. It is also possible to obtain a color composite image in which images reconstructed using the number of transmitted photons in each energy region are combined with different colors.

[Specific example of calculation by correction unit]
A specific example of the calculation by the correction unit 43 will be described.
The high frequency component of the signal output from the detection element 30 has a value proportional to the peak value (energy of the X-ray photon) of the original signal output from the detection element 30. In particular, in the present embodiment, since the high frequency component is extracted by the differentiating circuit 32, the signal X is a differential value of the original signal waveform. This differential value is large when the peak value of the original signal is large, and small when it is small.

This relationship is expressed as the following formula (1).

Here, “a” is a coefficient, and “h _n ” is the peak value of the signal X corresponding to the charge of the nth photon appearing in the waveform of the high frequency component.

Further, the low frequency component of the signal output from the detection element 30 can be considered as a function of the peak value of the signal until the time t _{n when} the original signal was measured and the input time of the signal, It is expressed as the following formula (2).

  Here, “b” is a coefficient, and the remaining part is the peak value of the signal Y. “Τ” is a time constant related to the attenuation of the output waveform from the detection element 30 and is determined based on the resistance value of the feedback system of the differentiation circuit 32 and the integration circuit 33, the capacitance of the capacitor, and the like.

Based on the above relationship, the energy value En of the original signal is expressed as the following mathematical formula (3).

  The energy value En is expressed in this way because it is difficult to completely eliminate the effect of pile-up even though it is a high-frequency component. Therefore, a low-frequency component capable of measuring the degree of pile-up is mainly used. Based on the need for correction of high frequency components. This relationship is schematically shown in FIG. The actual shape of the graph depends on, for example, the performance of the detection element 30, the responsiveness of the integrating circuits 33, 34a, and 34b, the capacitances of the capacitors 321 and 332 constituting the amplifier 31 and the differentiating circuit 32, and the values of the resistors 322 and 331. Different.

  The coefficients a and b can be calculated using, for example, a single energy RI (Radio Isotope). That is, if the RI is arranged in the opening 110 so that the distance from the X-ray detector 14 is constant, the counting rate is kept constant, and the energy value En and the peak values of the signals X and Y are known. Become. Therefore, by changing the distance between the X-ray detector 14 and the RI, the count rate is changed, and the relationship between the coefficients a and b at each count rate (for example, the value of the coefficient b when the count a is 1). ) Can be obtained.

  Then, when actually capturing a tomographic image of the subject P, the correction unit 43 calculates the crest value of the signal X and the crest value of the signal Y measured by the measurement unit 41 and the count rate calculation unit 42, for example. The original energy value of the photon may be calculated by substituting the coefficients a and b corresponding to the counted rate into Equation (3). Alternatively, based on Equation (3), a table in which the peak values of the signals X and Y and the original energy values of the photons are associated with each count rate is created and stored in the memory of the correction unit 43 in advance. The original energy value of the photon may be obtained from the table using the peak values of the signals X and Y measured by the unit 41 and the count rate calculated by the count rate calculation unit 42.

  On the other hand, in an X-ray CT apparatus using an X-ray tube, the energy of photons of X-rays is not single but has a continuous energy distribution determined by the tube voltage. Accordingly, continuous energy is obtained by using several types of RI having different energies, obtaining the coefficients a and b as described above for the photons of the respective energies, and interpolating the obtained coefficients between the energies. You may create the correction table or numerical formula which simulates distribution. Note that the accuracy of calculation of the original energy value of transmitted photons can be improved by assuming the shape of the energy distribution by a rough distribution of high-frequency components that are temporally close to the signal at the time of correction.

  It is also possible to obtain the coefficients a and b using the X-ray tube 13. The energy distribution of X-rays radiated from the X-ray tube 13 can be kept constant by fixing the tube voltage. Therefore, the coefficients a and b corresponding to the count rate are obtained by changing the tube current with the tube voltage fixed and changing the count rate while keeping the energy distribution constant. Can do. Further, when the coefficients a and b are obtained in this way, for example, by arranging a water phantom in the opening 110, correction data can be collected with an energy distribution close to that at the time of actual imaging of the subject, The correction accuracy can be improved.

  According to the configuration according to the present embodiment described above, the original energy value of the X-ray photons incident on the X-ray detector 14 can be accurately grasped even when the pile-up phenomenon occurs. In addition, since the original energy value of the photons can be accurately grasped in this way, it is possible to prevent the photons from being counted off. Therefore, even when the pile-up phenomenon occurs, the adverse effect on the reconstructed image is eliminated or reduced, and a good reconstructed image can be obtained.

  In addition, since it is not necessary to reduce the irradiation dose to the subject so as not to cause a pile-up phenomenon, it is possible to perform imaging in a short time while ensuring a sufficient S / N ratio, and to the subject and the operator. The burden of is reduced.

  More specifically, the original energy value of the photon is obtained in consideration of the count rate of photons in addition to the high-frequency component and the low-frequency peak value extracted from the output from the detection element 30. By considering the counting rate in this way, the original energy value of photons can be obtained more accurately.

[Modification]
The configuration disclosed in the above embodiment can be embodied by appropriately modifying each component in the implementation stage.

  For example, in the above embodiment, the high frequency component is extracted from the signal output from the detection element 30 by the differentiation circuit 32 and the low frequency component is extracted by the integration circuit 33. However, the high-frequency component and the low-frequency component may be extracted using other methods without being limited to the differentiation circuit and the integration circuit.

  Further, the circuit configuration disclosed in the above embodiment may be appropriately modified and implemented. For example, if it is not necessary to amplify a signal output from the differentiation circuit 32 or a signal output from the integration circuit 33, the amplifiers 34a and 34b may not be provided. Furthermore, one readout circuit (differential circuit 32, integration circuit 33, etc.) is connected to the plurality of detection elements 30 via a switching circuit, and the detection element 30 from which charges are read is appropriately changed by the switching circuit. The calculation unit 40 may execute processing such as energy value calculation for each detection element 30.

  In the above embodiment, the measurement unit 41 and the correction unit 43 are realized by the processor of the calculation unit 40. However, all or a part of the measurement unit 41, the count rate calculation unit 42, and the correction unit 43 may be realized as a circuit using various electronic components.

  Although several embodiments of the present invention have been described, these embodiments are presented by way of example and are not intended to limit the scope of the invention. These novel embodiments can be implemented in various other forms, and various omissions, replacements, and changes can be made without departing from the scope of the invention. These embodiments and modifications thereof are included in the scope and gist of the invention, and are included in the invention described in the claims and the equivalents thereof.

  DESCRIPTION OF SYMBOLS 1 ... X-ray CT apparatus, 13 ... X-ray tube, 14 ... X-ray detector, 15 ... Data acquisition part, 22 ... Data storage part, 23 ... Image reconstruction part, 25 ... Image display part, 30 ... Detection element 32 ... Differentiating circuit 33 ... Integration circuit 31, 34a, 34b ... Amplifier, 35a, 35b ... AD converter, 40 ... Operation part, 41 ... Measurement part, 42 ... Count rate calculation part, 43 ... Correction part, 44 ... Count part

Claims (4)

An X-ray generator for generating X-rays;
An X-ray detector that detects X-rays generated by the X-ray generator and transmitted through the subject, and outputs a signal corresponding to the detected X-rays;
A first extraction unit for extracting a high-frequency component of a signal output from the X-ray detector;
A second extraction unit for extracting a low-frequency component of a signal output from the X-ray detector;
A measurement unit that measures the peak value of the high-frequency component and the peak value of the low-frequency component corresponding to photons of the X-ray detected by the X-ray detector;
A count rate calculating unit that calculates a count rate indicating the number of photons detected by the X-ray detector per unit time;
A correction unit for obtaining an energy value of the photon based on the crest value of the high frequency component and the crest value of the low frequency component measured by the measurement unit, and the count rate calculated by the count rate calculation unit ;
A counting unit that counts the photons in which the energy value obtained by the correcting unit satisfies a predetermined condition;
A reconstruction unit for reconstructing an image based on the count result of the counting unit,
An X-ray computed tomography apparatus comprising:   The first extraction unit is a differentiation circuit connected to the X-ray detector, and the second extraction unit is an integration circuit connected to the X-ray detector. The X-ray computed tomography apparatus described. The correction unit multiplies the peak value of the high frequency component by a first coefficient corresponding to the count rate, and multiplies the peak value of the low frequency component by a second coefficient corresponding to the count rate. value and by adding the, X-rays computed tomography system according to claim 1, characterized in that to calculate the energy value. The correction unit obtains an energy value of the photon using a table that defines a photon energy value for a combination of a peak value of the high frequency component and a peak value of the low frequency component for each count rate. The X-ray computed tomography apparatus according to claim 1 .

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