JP2006101926A - Radiation detector, radiation image diagnostic device and generation method of radiation image - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide images by transmission radiations equivalent to the ones obtained in a conventional integration mode in the state of preventing artifacts due to a beam hardening phenomenon and the decline of the contrast resolution of soft tissue. <P>SOLUTION: A radiation detector is constituted of an X-ray detector 12 and a data gathering circuit 13. The X-ray detector 12 is a photon counting type detector. The radiation detector is provided with means (13(41-45)) for computing the counting data of the number of particles of the radiation classified into a plurality of energy regions on the energy spectrum of X-rays on the basis of the signals of respective gathering pixels outputted by the detector 12, means (13(46))for executing the weighting of a weight coefficient supplied for the respective energy regions to the counting data of each of the plurality of energy regions, and means (13(47)) for adding the counting data of each of the plurality of energy regions for the respective weighted gathering pixels and outputting the added data as the data for generating the radiation images for the respective gathering pixels. <P>COPYRIGHT: (C)2006,JPO&NCIPI

Description

  The present invention relates to an X-ray CT scanner or the like, a radiation image diagnostic apparatus that generates radiation images by irradiating a diagnostic object with radiation, detects transmitted radiation from the diagnostic object, and generates an image of the diagnostic object from the detection signal. In particular, it is equipped with a photon counting type (photon counting type) detector that counts the number of photons by considering the radiation transmitted from the diagnostic object as photons (particles). The present invention relates to a radiological image diagnostic apparatus and a radiographic image generation method.

  In recent years, as seen in medical image diagnosis and non-destructive inspection of structures, the importance of imaging the inside of a human body or an object using X-rays, which are a kind of radiation, has been increasing. In order to meet this requirement, it is required to image an object with a low X-ray dose, a high S / N, and a wide field of view, and research for this purpose is also actively conducted.

  This will be described by taking an X-ray CT scanner as an example. The X-ray CT scanner includes an X-ray detector, and this X-ray detector has historically started with an application of the ionization phenomenon of Xe gas. Recently, as an X-ray detector, a type of X-ray detector using a two-dimensional solid-state detector (a detector based on a combination of a scintillator and a photodiode) such as GOS, CsI, and CWO has come to be used. At the same time, in recent years, the field of view has been required to expand in the direction of the body axis, and the area of X-ray detectors has been increasing so as to support a wide field of view capable of volume scanning of the heart and lung fields.

This X-ray detector and data acquisition system (DAS) uses a scintillator, a photodiode, a preamplifier, an A / D converter, etc. to convert X-rays incident on the incident surface of the detector into a digital value. It has many kinds of circuit elements. The X-ray detector and the data acquisition system operate in an integration mode as a detection operation mode as seen in, for example, Patent Document 1 and Non-Patent Document 1. That is, after converting the total energy of incident X-rays into light, this optical signal is converted into an electrical signal, and this electrical signal is integrated for a predetermined time to be output as a detection signal.
JP 2000-131440 A “Current Status and Future of Normal Temperature Semiconductor Radiation Detectors”, Autumn Meeting Committee Meeting Session of the Atomic Energy Society of Japan, Announcement date: September 12, 1999, Presenter: Masaki Misawa, Institute of Industrial Technology, Ministry of International Trade and Industry "

  Since the above-described X-ray detector and data acquisition system operate in the integration mode, the energy of the X-ray is detected as its integration value. That is, all the X-ray energies from low energy to high energy are integrated. For this reason, since information of relatively low energy X-rays is buried, an image reconstructed using such a detection signal is caused by hardening of the X-ray quality (so-called beam hardening phenomenon). There has been a problem that image quality is deteriorated due to occurrence of artifacts or reduction in contrast resolution of soft tissue.

  The present invention has been made in view of the problems of the above-described conventional radiographic image diagnosis apparatus such as an X-ray CT scanner operating in an integration mode, and is based on the beam hardening phenomenon without increasing statistical noise. The purpose is to prevent artifacts and reduction in contrast resolution of soft tissue.

  In order to achieve the above object, a radiation detection apparatus according to the present invention has, as one aspect thereof, a plurality of collection pixels that allow radiation to enter and regards the radiation incident on each of the collection pixels as a photon. A photon counting detector that outputs an electrical signal corresponding to the energy of the particle, and the radiation classified into a plurality of energy regions on the energy spectrum of the radiation based on the signal of each collected pixel output by the detector Region-by-region data calculation means for calculating the count data of the number of particles, and a weighting factor given to each energy region in the count data of each of the plurality of energy regions for each collection pixel calculated by the region-by-region data calculation means Weighting means for weighting, and counting data for each of the plurality of energy regions for each collection pixel weighted by the weighting means. The then added together for each collection of pixels, characterized in that an adding means for outputting a radiation image generation data to obtain the added data.

  Preferably, the region-by-region data calculating means counts the signal output by the detector into the plurality of energy regions, and the signal distinguished by the discrimination means for each energy region. Counting means for passing a count value as the count data to the weighting means.

  For example, the weighting unit is a unit that applies the weighting with a larger weighting coefficient as the count data from the energy region on the high energy side or the low energy side of the plurality of energy regions for each of the collected pixels. Further, the weighting unit performs the weighting with a larger (or smaller) weighting factor as the count data from the energy regions on the high energy side and the low energy side of the plurality of energy regions for each of the collected pixels. It may be a means.

  In addition, the radiological image diagnostic apparatus according to the present invention typically captures a radiographic image using the radiological detection apparatus and radiographic image generation data for each collection pixel output from the adding means of the radiological detection apparatus. And an image generation means for generating. The radiation diagnostic apparatus according to the present invention includes various medical, nondestructive inspection, airport baggage inspection, internal structure analysis X-ray CT scanners, X-ray diagnostic apparatuses, various scintillation cameras, scintillation probes, gamma rays. Non-destructive inspection equipment.

  Furthermore, the radiation image generating method according to the present invention detects incident radiation as particles and detects each collected pixel, and outputs an electrical signal corresponding to the energy of the incident radiation particles. And calculating the count data of the number of particles of the radiation classified into a plurality of energy regions on the energy spectrum of the radiation based on the signal of each collected pixel, and each of the plurality of energy regions for each of the calculated collection pixels. Is weighted with a weighting factor given to each energy region, and the weighted collection data for each of the plurality of energy regions is added to each other to add the added data to the radiation image for each collection pixel. It is output as generation data, and a radiographic image is generated using the radiographic image generation data.

  According to the present invention, a signal detected for each collection pixel by using incident radiation as particles is used. Based on this signal, count data of the number of particles of the radiation classified into a plurality of energy regions on the energy spectrum of the radiation is calculated. The calculated count data for each of the plurality of energy regions for each collected pixel is weighted with a weighting factor given to each energy region. The weighted count data of the plurality of energy regions are added to each other, and the added data is output as radiation image generation data for each collected pixel. For this reason, the information of a low energy component can also be reliably taken in by changing the value of a weighting coefficient for every energy region according to the height in an energy spectrum. Therefore, by setting the weighting factor with an appropriate weighting pattern that changes according to the particle energy, it can be obtained in the conventional integration mode while preventing artifacts due to the beam hardening phenomenon and deterioration of the contrast resolution of soft tissue. It is possible to provide an image based on transmitted radiation that is equivalent to that of the previous one.

  Embodiments of a radiological image diagnostic apparatus and a radiographic image generation method according to the present invention will be described below with reference to the accompanying drawings.

(First embodiment)
A first embodiment according to the present invention will be described with reference to FIGS.

  In this embodiment, the radiological image diagnostic apparatus according to the present invention is implemented as an X-ray CT (Computed Tomography) scanner, and the radiographic image generation method according to the present invention is also implemented in this X-ray CT scanner.

  FIG. 1 shows a schematic configuration of this X-ray CT scanner.

  As shown in FIG. 1, the X-ray CT scanner 1 includes a gantry 2, a high voltage generator 3, a bed 4, and a console 5. Inside the gantry 2, an X-ray tube 11 and an X-ray detector (radiation detector) 12 as an X-ray source (radiation source) (not shown) are arranged to face each other through a cylindrical opening OP. A pair of the X-ray tube 11 and the X-ray detector 12 is disposed in the gantry 2 so as to be rotatable around the opening OP.

  A high voltage is supplied to the X-ray tube 11 from the high-voltage generator 3, so that, for example, pulsed X-rays having a continuous spectrum are exposed toward the opening OP. The top plate 4P of the bed 4 is inserted into the opening OP so as to be able to advance and retract. Since the subject P is placed on the top 4P, the X-rays emitted from the X-ray tube 11 pass through the subject P and go to the opposite side of the opening OP.

  As shown in FIG. 2, the X-ray detector 12 has a slice direction corresponding to the body axis direction of the subject P and a channel direction orthogonal to the slice direction, and a plurality of the X-ray detectors 12 along the two-dimensional plane. This is a two-dimensional detector in which a large number of pixels are arranged in a two-dimensional array. In addition, the channel direction of the X-ray detector 12 is curved in consideration of the spread angle of the X-ray beam from the X-ray tube 11 in particular. Note that the overall shape of the X-ray detector 12 is determined by the application and may be a flat plate shape.

  The X-ray detector 12 is divided into a plurality of detector blocks 12J so that the two-dimensional surface is divided into a plurality of surfaces, and the detector blocks can be detachably coupled. Further, an X-ray transmission image from the X-ray tube 11 can be obtained with a collimator (not shown) made of molybdenum or tungsten arranged in the slice direction on the front surface of the detector block 12J on the X-ray incident side. It has become.

  Each detector block 12J is made of a compound semiconductor and has a layered semiconductor cell S having a predetermined size (for example, several centimeters × several centimeters), and a radiation incident surface of the semiconductor cell S is covered with a charging electrode E1 for voltage application. The semiconductor cell S has a monolithic structure in which a surface opposite to the radiation incident surface is covered with a plurality of collecting electrodes E2 divided into a two-dimensional array (a grid pattern). The current collecting electrode E2 corresponds to each pixel. As a material of the semiconductor cell S, a cadmium telluride semiconductor (CdTe semiconductor), a cadmium zinc telluride semiconductor (CdZnTe semiconductor), a silicon semiconductor (Si semiconductor), or the like is used. A relatively high voltage of, for example, several tens to several hundreds V is applied to the charging electrode E1. As a result, due to the X-ray photons incident on the semiconductor cell S, pairs of electrons and holes are generated therein, and these electrons are collected at each of the relatively positive potential collecting electrodes E2, and this electron Is detected as a pulse signal. That is, the X-rays incident on the radiation incident surface are directly converted into electric quantity pulse signals.

  The size of the collection pixel for X-rays is determined by the size of each of the plurality of electrodes E2 divided in a grid pattern. This size is small enough to detect X-rays as photons (particles). As a result, the X-ray detector 12 capable of photon counting (photon counting) is configured, and as a whole, a predetermined number of matrix acquisition pixel channels are formed.

  In the present embodiment, the size capable of detecting X-rays as particles is “between pulse signals responding to each incident when a plurality of radiation (for example, X-rays) particles are successively incident at or near the same position. Is defined as a size that can substantially ignore the occurrence of the superposition phenomenon. When this superposition phenomenon occurs, the X-ray particle counting-down characteristic occurs in the characteristic of “number of incidents versus actual number of measurements” of the X-ray particles. For this reason, the size of the collection pixel formed in the X-ray detector 12 is set to a size that can be considered that the counting-down characteristic does not occur or does not substantially occur.

  An X-ray detector capable of photon counting is also known in Japanese Patent Laid-Open No. 2000-69369.

  Thereby, the X-rays that have passed through the subject P are counted as X-ray particles (that is, X-ray photons) by the two-dimensional X-ray detector 12 provided inside the gantry 2 at regular intervals, and are converted into photon energy. A corresponding analog detection signal is output for each pixel.

  The detection signal of each pixel output from the X-ray detector 12 is sent to a data acquisition circuit 13 provided in the gantry 2. The data collection circuit 13 and the X-ray detector 12 form a radiation detection apparatus RDS according to the present invention.

As shown in FIG. 3, the data collection circuit 13 is formed with a data collection circuit C having a plurality of channels in accordance with the semiconductor cells S constituting the collection pixels of the X-ray detector 12. Each data collection circuit C includes a charge amplifier 41 that receives an analog electric signal output from the semiconductor cell S (sensor). A waveform shaping circuit 42 and a multi-stage comparator 43 are provided at the subsequent stage of the charge amplifier 41. ₁ to 43 _n , energy region distribution circuit 44, multistage counters 45 _{1 to} 45 _n , weighting circuit 46, and addition circuit 47 are sequentially provided.

As a modification, various processing circuits for a plurality of channels constituting the data acquisition circuit 13 are formed as an ASIC layer (application-specific integrated circuit) on the back surface of the semiconductor cell (sensor) of the X-ray detector 12. But you can. That is, the waveform shaping circuit 42 forming each data acquisition channel, the multistage comparators 43 ₁ to 43 _n , the energy region distribution circuit 44, the multistage counters 45 _{1 to} 45 _n , the weighting circuit 46 that is a common circuit, and the addition The circuit 47 is formed integrally with the X-ray detector 12. Of course, the data collection circuit 13 may be configured separately from the X-ray detector 12 as in the present embodiment. As another modification, the weighting circuit addition circuit 46 and the addition circuit 47 which are common circuits may be provided in the console 5, and the functions of these circuits 46 to 48 are implemented by an arithmetic unit provided in the console 5. You may make it do.

  In each data collection channel shown in FIG. 3, the charge amplifier 41 is connected to each collector electrode E2 of the semiconductor cell S, and charges up the charges collected in response to the incidence of X-ray particles to Output as a pulse signal. The output terminal of the charge amplifier 41 is connected to a waveform shaping circuit 42 whose gain and offset can be adjusted. The waveform of the detected pulse signal is processed with a previously adjusted gain and offset to shape the waveform. The gain and offset of the waveform shaping circuit 42 are adjustment parameters that take into account the non-uniformity with respect to the charge charge-up characteristics of each collection pixel of the semiconductor cell S. By adjusting the gain and offset of the waveform shaping circuit 42 for each collection pixel channel in advance by a calibration operation, it is possible to perform waveform shaping that eliminates such non-uniformity. As a result, the waveform-shaped pulse signal output from the waveform shaping circuit 42 of each data acquisition channel has characteristics that substantially reflect the energy amount of the incident X-ray particles, and this is applied between the acquisition pixel channels. The variation is almost eliminated.

The output terminal of the waveform shaping circuit 42 is connected to the comparison input terminals of the plurality of comparators 43 ₁ to 43 _n . Reference values TH1 (to THn) having different values are respectively applied to the reference input terminals of the plurality of comparators 43 ₁ to 43 _n . As described above, the reason why one pulse signal is compared with different reference TH1 (to THn) is that the energy amount of the incident X-ray particle enters which of the energy regions set in advance divided into a plurality. To do that information. As illustrated in FIG. 4A, which value of the reference value TH1 to TH3 (when n = 3) the peak value of the pulse signal (that is, the energy amount of each incident X-ray particle) exceeds. Depending on, the energy region to be distinguished differs. When the peak value is between the reference values TH1 and TH2, the energy amount of the X-ray particles currently being measured is discriminated as entering the energy region 1, and when the peak value is between the reference values TH2 and TH3, The energy amount of the X-ray particles being measured is discriminated from being in the energy region 2, and if the reference value TH3 or more, the energy amount of the X-ray particles being currently measured is discriminated from being in the energy region 3. . The lowest reference value TH1 is set as a threshold for preventing erroneous detection that does not detect disturbance or white noise caused by circuits such as the semiconductor cell S and the charge amplifier 41. Further, the number of reference values, that is, the number of comparators is not necessarily limited to three (the number of discriminable energy regions = 3), and may be two or four, or one in some cases. It may be. In the case of one, information on whether or not X-ray particles are incident can be obtained.

  FIG. 5 shows an example of setting the reference values TH1 to THn. The waveform shown in the figure shows a continuous spectrum of the energy of X-rays emitted from a commonly used X-ray tube. The count number on the vertical axis is an amount depending on the X-ray tube current, and the energy spectrum on the horizontal axis is an amount depending on the tube voltage of the X-ray tube. For this spectrum, the first reference value TH1 is set to a value that allows the number of X-ray particles to be discriminated from the non-countable region and the lower energy region 1, and the second reference value TH2 is set to the lower energy region. 1 and medium energy region 2 are set to values that can be distinguished from each other, and third reference value TH3 is set to a value that allows medium energy region 2 and higher energy region 3 to be distinguished from each other. ing. In this embodiment, for example, by setting the three reference values TH1 to TH3 in this way, incident X-rays are detected as particles and then classified into energy regions 1 to 3, respectively. The set number and set value of the reference values TH1 to THn are arbitrary.

Output terminals of the plurality of comparators 43 ₁ to 43 _n are connected to the energy region distribution circuit 44 as shown in FIG. The energy region distribution circuit 44 reads which comparison result is on (off) among the outputs of the plurality of comparators 43 ₁ to 43 _n , and in which region the energy of the X-ray particles currently measured is in which region. It has a function of sorting whether to enter. Specifically, the circuit 44 is formed by a combination of logic circuits, for example. For example, in the case of the pulse signal shown in FIG. 4 (a), because the peak value (energy amount) shows an energy region 2, comparators 43 _1, 43 ₂ of the output is on, as shown in FIG. 4 (b) next, and the output of comparator 43 _n are turned off. For this reason, the energy region distribution circuit 44 reads and decodes the on / off information and outputs an energy distribution signal from its output terminal. As an example, in the case of FIG. 4B, this energy distribution signal is obtained by turning on only a signal indicating “energy region 2”. As a result, distribution information indicating that the energy of the X-ray particles currently measured enters “energy region 2” is obtained.

The plurality of output terminals of the energy region distribution circuit 44 are connected to the plurality of counters 45 _{1 to} 45 _n , respectively, and the energy distribution signal is supplied to each of the plurality of counters 45 _{1 to} 45 _n . Therefore, each of the counters 45 _{1 to} 45 _n increments (counts up) the count value only when the input energy distribution signal is on, and sets the number of X-ray particles entering each energy region in charge for a certain time. It can be measured across. As an example, FIG. 4 (c), the In the situation (d), the so the energy of X-ray particle enters the energy region 2, the count value of the counter 45 ₂ which measures the energy region 2 is incremented .

In this manner, the number of X-ray particles incident on the X-ray detector 12 by the plurality of counters 45 _{1 to} 45 _n during each fixed time until resetting is set for each collection pixel and for each energy region. It is measured. This count value, that is, the count value of the number of X-ray particles, is output as count data K of a digital amount from each of the plurality of counters 45 _{1 to} 45 _n .

  The count data is then sent to a weighting circuit 46 where the count values in the energy regions are weighted as will be described later.

Specifically, the weighting circuit 46, as shown in FIG. 6, the n-number of multipliers ₅₁ 1 to 51 _n provided in accordance with the number of the counter ₄₅ 1 to 45 _n, these multipliers ₅₁ 1 ~ and a weighting coefficient generator 52 to provide a 51 _n. Each of the multipliers 51 _{1 to} 51 _n multiplies the count value output from the corresponding counter 45 ₁ (to 45 _n ) by the weighting coefficient W given separately from the weighting coefficient generator 52, thereby calculating the count value. Is weighted. As an example, the weighting coefficient generator 52 receives weighting information for each X-ray energy region sent from the controller of the console 5, and according to this information, assigns a weighting coefficient W to each of the multipliers 51 _{1 to} 51 _n. generate.

  As an example, the weighting method here (the basis thereof will be described later) emphasizes a high energy region for the purpose of suppressing artifacts due to beam hardening as shown in FIG. Increase the value); enhance the low energy region for the purpose of improving soft tissue contrast, as shown in FIG. 5B; or suppress artifacts due to beam hardening as shown in FIG. It is based on one of a plurality of modes in which both areas are emphasized together for the purpose of improving contrast of soft tissue.

The data weighted by the multipliers 51 _{1 to} 51 _n are added to each other by the adder circuit 47 and output as collected data based on photon counting for one collection pixel. As shown in FIG. 4, data collection with this weighting process is executed in parallel for the semiconductor cells (sensors) of the collected pixels, so that data can be collected in the same manner for all the collected pixels of the X-ray detector 12. ing.

(The basis for weighting and its specific method)
The weighting process performed by the weighting circuit 46 is central to the features of the present invention, and is based on the principle obtained by the inventor after research.

This principle will be described below. Assuming that the energy fluence of X-rays is Ψ (E) and the energy spectrum of X-rays is φ (E), the energy fluence is Ψ (E), and the energy spectrum is φ (E) multiplied by the energy E of each photon. And
[Equation 1]
Ψ (E) = φ (E) · E (1)
It can be expressed as. In the conventional scintillation detector, the energy is measured with the energy fluence as an integral value of Ψ (E), that is, a total value. As in this embodiment, when a photon counting type detector is employed, the energy spectrum φ (E) of the X-rays at the detector can be known, so an energy weighting function w (E) is introduced into this,
[Equation 2]
Ψ (E) ′ = φ (E) · w (E) (2)
Can be created in a pseudo manner. Therefore, equation (1) can be regarded as an example when weighting w (E) = E in equation (2).

  Therefore, the present inventor uses various functions as the energy weighting function w (E) to improve problems such as artifacts caused by beam hardening occurring in an existing X-ray CT scanner and a decrease in contrast resolution of soft tissue. I found out that I can do it.

Typically, as an energy weighting function w (E), when applying high weight to high energy (High Energy Weighted: HEW), w (E) = E ² and when applying high weight to low energy ( Low Energy Weighted: LEW) can be set to w (E) = (160−E) ² . Note that the reconstruction of a conventional X-ray CT image is energy fluence (EF), and corresponds to w (E) = E. This is equivalent to obtaining an image having the same quality as that obtained by a conventional integrating detector. The weighting factor W based on the energy weighting function w (E) is normalized in any case.

  Of course, in addition to this, it is possible to apply a strong load to medium and high energy, and it is also possible to apply a strong load to low and medium energy. Further, the weight curve can be set so as to draw a downwardly convex parabola so that both high energy and low energy are weighted more strongly than medium energy.

(Relationship with artifact removal due to beam hardening)
When X-rays pass through a medium having a large absorption coefficient, X-rays with lower energy are more likely to be absorbed by the medium. Therefore, when the transmitted X-rays are observed, relatively more X-rays on the high energy side remain. Become. This phenomenon is called radiation hardening (beam hardening), which causes artifacts. This will be explained using the spectrum shown in FIG. 5. When beam hardening occurs in the transmitted X-ray, the average energy shifts to a higher energy side than that at the time of incidence, so that it is strong against high energy in the equation (2). By adopting HEW that applies weight, artifacts due to beam hardening are reduced.

(Improvement of soft tissue contrast resolution)
The contrast in the X-ray CT image represents the difference in the linear absorption coefficient of the medium constituting the target substance. Since the value of the linear absorption coefficient is close in soft tissue, it is difficult to express the difference. However, in the case of a photon counting type detector, information on low-energy X-ray absorption can also be obtained, and this information can be used to emphasize the difference in substances. That is, in this case, the contrast of the soft tissue can be enhanced by adopting LEW that applies a strong weight to the low energy side in the equation (2).

(simulation)
Based on the above considerations, a Monte Carlo simulation was performed and compared with a conventional image. The result was that both the artifact caused by beam hardening and the contrast resolution of soft tissue were improved.

  Returning to FIG. 1, the digital data of each collected pixel output from the adder circuit 47 provided corresponding to each collected pixel of the data collecting circuit 13 is view data (collection data) that is radiation image generation data. Data) is sent to the fixed part in the gantry 2 via a slip ring or a non-contact type communication connection (not shown). This is transmitted from the fixed part of the gantry 2 to the console 5 via a communication line. The console 5 includes a controller 20 that controls all the components of the X-ray CT scanner 1 and sends the above-described weighting factor information to the weighting circuit 46. In addition to the controller 20, the console 5 includes an interface 21, a storage unit 22, a reconstruction unit 17, a display device 18, and an input device 19.

  The display 18, the input device 19, and the computer 20 form an interface with the surgeon. Based on a predetermined algorithm given by the computer 20, energy is transmitted to the surgeon as a function of a part of the interface. The desired profile of the weighting function w (E) (ie, the level of the weighting factor determined by the energy weighting function w (E)) can be determined.

  This process is executed by the computer 20 as shown in FIG. 8, for example. The computer 20 asks the operator whether or not to display a reference image for determining a desired profile of the energy weighting function w (E) via the display 18 (step S1). When the surgeon responds to the inquiry with display of a reference image (step S1, YES), the computer 20 displays the existing reference image on the display 18 (step S2). This reference image is displayed in order for the operator to visually determine whether artifacts caused by beam hardening or contrast reduction of soft tissue have occurred, and the diagnostic site to be imaged is previously imaged and stored. This applies to pre-scan images. If there is no prepared reference image (NO in step S1), the surgeon performs profile designation based on experience values and the like, and the display process in step S2 is skipped.

  Next, the computer 20 receives a pattern designation of a desired profile of the energy weighting function w (E) from the operator (step S3). Next, as shown in FIG. 7A, this pattern is a pattern A in which the weighting factor is increased as it is in a high energy region, or a pattern B in which the weighting factor is increased as it is in a low energy region as shown in FIG. 7B. Alternatively, as shown in FIG. 7C, it is determined whether the pattern C has a high weighting factor in the low energy region and the high energy region (steps S4 to S6).

  The pattern A shown in FIG. 7A is from the region on the high energy side among the energy regions (energy regions 1 to 3) in which the energy region excluding the non-measurable region is divided into three regions of low, medium, and high. It shows that the weighting coefficient W is increased as the obtained count data (number of particles measured). On the other hand, the pattern B shown in FIG. 7B is count data (number of particles measured) obtained from the low energy side region of the above three energy regions (energy regions 1 to 3). As shown, the weighting coefficient W is increased. Furthermore, the pattern B shown in FIG. 7C is the count data (particle measurement) obtained from the energy regions on the low energy side and the high energy side of the above three energy regions (energy regions 1 to 3). It is shown that the weighting coefficient W is increased as the number increases.

  For this reason, when the desired profile can be determined, the computer 20 receives adjustment information of the weighting degree from the operator for each of the patterns A, B, and C, generates a control signal corresponding to the adjustment information, and outputs the control signal to the weighting circuit. 46 weight coefficient generators 52 (see FIG. 6) are sent (steps S7 to S12). That is, even with the same pattern A, the weighting degree (normalized weight coefficient level) on the high energy region side can be adjusted (see, for example, curves CV1 and CV2 in FIG. 7A). The person can select this arbitrarily. The adjustment of the weighting degree in the same pattern can be similarly performed for the patterns B and C.

  In the console 5, the sent view data is temporarily stored in the storage unit 22. Imaging is performed by moving a pair of the X-ray tube 11 and the X-ray detector 12 around the subject while moving the top of the bed or fixing the position of the top according to a predetermined scanning method. The X-ray scan is performed on the subject by exposing the subject to X-rays at each rotational position moved by a predetermined angle during the rotation. As a result, the above-described view data based on the count data of the X-ray particles incident on each collection pixel at each rotation position is obtained, and is sequentially stored in the storage unit 22. Accordingly, the reconstruction unit 17 reads out the data of each view necessary for image reconstruction from the storage unit 22 at an appropriate timing, reconstructs this data with a desired reconstruction algorithm, and performs energy-specific X-ray particles. A CT image similar to the CT image in integral mode operation based on the number is obtained.

  This CT image is displayed on the display 18, for example. The console 5 is provided with a post-processing processor for post-processing, and appropriate post-processing such as correction processing is performed on the view data transmitted from the data collection circuit 13, and the post-processed data is reconstructed. You may make it use for.

  As described above, according to the present embodiment, since the energy spectrum of incident photons of X-rays is artificially weighted for each energy region, the disadvantage that the degree of photon attenuation is a function of the energy value. Can be avoided or reduced. Conventionally, the energy of X-rays was measured as an integral value, so that it was not possible to cope with changes in beam quality such as beam hardening, leading to a reduction in artifact and soft tissue contrast resolution. However, X-rays are detected by the photon counting method, and the X-ray particle count data of each energy region for each collected pixel is weighted with a desired profile for each energy region, and then added to each other to obtain a view of each collected pixel. I have data. For this reason, this low energy information can be reflected in the view data of each finally collected pixel without missing the low energy information of the X-ray particles without increasing the statistical noise.

As described above, the separation of the energy regions is not necessarily limited to the three energy regions 1 to 3 (see FIG. 5), and an arbitrary number of separation regions may be set. For example, in the X-ray energy spectrum shown in FIG. 5, the remaining energy regions excluding the non-measurable region may be divided into four, five, or six, or may be subdivided into seven or more. In order to perform such further division of the energy region, a configuration corresponding to the number of divisions is adopted, that is, as a signal processing channel corresponding to each collected pixel, multistage comparators 43 ₁ to 43 _n , energy region distribution. circuit 44, a multi-stage counter ₄₅ 1 to 45 _n, the weighting circuit 46, and an adder circuit 47 may be Sonaere similarly to FIG. Specifically, as many comparators 43 as the number of divisions are prepared, and a reference value TH corresponding to the number of energy regions to be divided is given to each of these comparators 43. A counter 45 is also prepared according to the number of divisions, that is, the number of comparators 43, and is arranged between the energy region distribution circuit 44 and the weighting circuit 46. For this reason, the count data for each fixed time of each counter 45 is sent to the weighting circuit 46, and the count data is weighted using the weighting factor W for each energy region commanded from the controller 20. The weighted count data for each energy region are added to each other by the addition circuit 47 and sent to the console 5 as view data (image reconstruction data) for each collected pixel.

  Thus, as the number of energy regions to be subjected to photon counting on the energy spectrum is increased for each collected pixel, the amount of information on the amount of particle energy that can be captured in the view data increases. A CT image similar to the integration operation mode can be provided.

  The radiation detection apparatus and the radiological image diagnostic apparatus according to the present invention are not necessarily limited to the X-ray CT scanner, and may be an X-ray diagnostic apparatus that specially captures a transmission image.

(Second Embodiment)
With reference to FIGS. 9-10, embodiment of the radiographic image diagnostic apparatus and radiographic image generation method which concern on this invention is described. In the description of the second embodiment, the same or equivalent components as those used in the first embodiment described above are denoted by the same reference numerals, and the description thereof is simplified or omitted.

  In the second embodiment, radioisotope (RI) or a labeled compound thereof is selectively taken into a specific tissue or organ in a living body, and gamma rays emitted from the RI are measured from outside the body. The present invention relates to a nuclear medicine diagnostic apparatus as a radiological image diagnostic apparatus for imaging a dose distribution of RI. This nuclear medicine diagnostic apparatus is implemented as, for example, a SPECT (single photon ECT) apparatus or a positron ECT (positron ECT) apparatus. Here, the nuclear medicine diagnostic apparatus is implemented as a SPECT apparatus.

  The SPECT apparatus according to the second embodiment is characterized in that photon counting is performed by further subdividing a specific energy region among a plurality of energy regions for performing photon counting, similar to FIG. 5 described above. . If this view is changed, it can be said that the energy value shown on the horizontal axis in FIG. 5 is a technique for further subdividing a specific region of interest into photon counting.

  FIG. 9 shows an example of the signal processing circuit 113 mounted on the detector of the SPECT apparatus that implements this. The detector equipped with this signal processing circuit 113 is configured as a gamma camera. The signal processing circuit 113 includes a data acquisition circuit C of a plurality of channels corresponding to all the collection pixels provided corresponding to the semiconductor cell S for each collection pixel of the gamma ray detector 112.

As shown in the figure, each data collection circuit C includes a charge amplifier 41, a waveform shaping circuit 42, a plurality of comparators 143, an energy region distribution circuit 144, a plurality of counters 145, connected to the semiconductor cell S, for each collection pixel. A weighting circuit 146 and an addition circuit 147 are provided. Among them, a plurality of comparators 143, a first comparator 143 _1, the second comparator group 143, _second and third comparators 143 ₃ connected in parallel with each other at the output terminal of the waveform shaping circuit 42 Prepare. Further, the second comparator group 143 ₂ includes a first stage comparator 143 ₂₁ , a second stage comparator 143 ₂₂ , and a third stage comparator 143 ₂₃ connected in parallel to the output terminal of the waveform shaping circuit 42. Is provided. Thereby, the output signal of the waveform shaping circuit 42 is output in parallel to the non-inverting input terminals of all the comparators.

Among the inverting input terminals of the comparators 143 _{1 to} 143 _3, the inverting input terminal of the first comparator 143 ₁ is given a reference value TH1 for energy region discrimination, and the second comparator group 143 ₂ Reference values TH2, TH21, and TH22 are respectively supplied to the inverting input terminals of the first stage comparator 143 ₂₁ , the second stage comparator 143 ₂₂ , and the third stage comparator 143 ₂₃ of the first stage comparator 143 ₂₁ . reference value TH3 is applied to the inverting input of 143 _3. These reference values are set to have a magnitude relationship of TH1 <TH2 <TH21 <TH22 <TH3. From these reference values TH1, TH2, TH3, as shown in FIG. A non-measurable region, an energy region 1, an energy region 2, and an energy region 3 are set in order. Further, by setting the reference values TH21 and TH22 between the reference values TH1 and TH3, the intermediate energy region 2 is further divided into _three sub energy regions 2 ₁ , 2 ₂ , 2 ₃ in order from the low energy side. It has become so. This intermediate energy region 2 is normally set so that a photo peak PK of the energy spectrum of gamma rays incident on the position of each collection pixel of the detector 113 enters.

The results compared by the comparators 143 _{1 to} 143 ₃ are sent to the energy domain distribution circuit 144, respectively. For this reason, by this distribution circuit 144, as in the first embodiment, the energy region 1 (˜3) (and sub-energy) of the gamma rays (assumed to be particles) incident on each collection pixel of the detector 113 are displayed. It is finally determined whether it is classified into the region 2 ₁ (˜2 ₃ ).

A plurality of counters 145, energy regions 1,3 and an intermediate low energy side and the high energy side sub energy region _{2 1,} 2 2, _{2 3} of each of the plurality of counters ₁₄₅ corresponding to each other (145 1, _{145 21} , 145 ₂₂ , 145 ₂₃ , 145 ₃ ). For this reason, each counter 145 receives a signal output from the energy region distribution circuit 145 indicating that it has entered the energy region in charge of gamma rays regarded as particles, and counts up when this signal is on. . Count data of the plurality of counters 145 is sent to the weighting circuit 146.

In this weighting circuit 146, as shown in FIG. 10B, a weighting coefficient W is commanded from the controller 20 according to each of the energy regions 1 and 3 and the intermediate sub-energy regions 2 ₁ , 2 ₂ , and 2 _3. Has been. The weighting coefficient W is composed of five values W1, W2, W3, W4, and W5 assigned to each of these areas. In this embodiment, the weighting coefficient W is set to have a relationship of W5 ≦ W1 <W2 <W3 <W4. The energy region 1 of low energy side is assigned a weighting factor W1, sub energy region ₂ 1 of the first stage to 3-stage intermediate energy regions _2, 2 2, _{2 3} are weighting coefficients W2~W4 this The weighting factor W5 is assigned to the energy region 3 on the high energy side. Among these, the weighting coefficient W5 assigned to the energy region 3 on the high energy side is set to W5 = 0 in the present embodiment, but may be W5≈0. W1 = 0 may also be used.

  As a result, the weighting circuit 146 performs weighting by multiplying the count data (count value) of the number of gamma-ray particles by the given weighting count W for each energy region. The weighted count data is sent to the adder circuit 147 and added to the adder circuit 147. As a result, data for generating a SPECT image of each collected pixel is obtained, and this data is sent to the console 5 for generation of a SPECT image.

  Configurations and operations other than those described above are the same as or equivalent to those described in the first embodiment.

  For this reason, according to the SPECT apparatus according to the second embodiment, the detector 113 may be affected by the temperature and the magnetic field, and the detection characteristics may drift. However, the influence of such a drift is minimized. To the limit.

  That is, when such a drift phenomenon occurs, as indicated by the arrow D in FIG. 10A, the peak of the count value (photo peak) moves in the horizontal axis direction, and the change in the count value curve due to this movement amount is ignored. There are things you can't do. This also causes artifacts in the SPECT image, and in the case of an image including many scattered rays, it is easily affected by such a drift phenomenon.

In contrast, according to the present embodiment, the intermediate energy regions 2 will enter the photo peak (sub energy region 2 _{1 (~ 2))} performs photon counting and weighted more finely divided, and The energy region 3 on the high energy side is weighted with the weighting coefficient = 0 that has been dropped sharply.

  Thereby, the following advantages are obtained. In nuclear medicine, scattered radiation also includes some positional information. For this reason, the weighting coefficients W1 to W4, which are set in the intermediate energy region 2 and the energy region 1 on the low energy side, are finely changed in stages, and the information is taken in, and the weighting set in the energy region 3 on the high energy side With the coefficient W5 = 0, the coefficient value curve moved to the high energy side due to the drift phenomenon can be corrected, and an effect similar to that obtained by removing the scattered radiation can be obtained. Further, this makes it possible to reduce artifacts due to drift of the SPECT image, while avoiding statistical noise increase as much as possible, and to provide a higher-quality and stable SPECT image.

  In the case of a nuclear medicine diagnostic apparatus, since the energy value of radiation from RI can usually be measured as a continuous value, the weighting coefficient to be multiplied by this value can be continuously changed. In this case, since the energy value itself output as a digital value can be interpreted as a result of comparison in each window, it is included in the scope of the present invention.

  Note that the radiation detector and the radiation imaging system according to the present invention are not limited to the configurations described in the above-described embodiments and the configurations of modifications thereof, and those skilled in the art will understand the gist described in the claims. Changes and modifications can be made as appropriate without departing from the scope. For example, as described above, the X-ray detector is not limited to the collection pixel having a two-dimensional array, and the collection pixel is formed by stacking a one-dimensional array detector or a two-dimensional array detector. A detector having a three-dimensional array may be used.

BRIEF DESCRIPTION OF THE DRAWINGS Explanatory drawing which shows schematic structure of the X-ray CT scanner as a radiographic image diagnostic apparatus concerning the 1st Embodiment of this invention. The perspective view which shows the outline of the X-ray detector mounted in the X-ray CT scanner. The block diagram which shows the electrical schematic structure of the X-ray CT scanner centering on an X-ray detector and a data acquisition circuit. The timing chart explaining from the energy discrimination | determination of incident X-ray particle to the counting operation | movement. The graph which shows the relationship between the energy spectrum of the X-rays exposed from a normal X-ray tube, and a threshold value. The block diagram which shows the outline of the weighting circuit mounted in a data collection circuit. The graph explaining various energy weighting functions according to a pattern. The flowchart explaining the outline | summary of the process for setting a weighting coefficient based on the dialogue with a user. The block diagram which shows the electrical schematic structure centering on the gamma-ray detector and data collection circuit of a SPECT apparatus (nuclear medicine diagnostic apparatus) as a radiographic image diagnostic apparatus concerning the 2nd Embodiment of this invention. The figure explaining the weighting process of the count value curve by which the energy area | region and weighting coefficient which are set with a SPECT apparatus, and the photon count of the gamma ray were carried out.

Explanation of symbols

1 X-ray CT scanner (radiological image diagnostic equipment)
RDS radiation detector 2 gantry 3 high voltage generator 4 bed P subject 11 X-ray tube (radiation source)
5 Console 12 X-ray detector (radiation detector)
13 Data collection circuit 20 Controller 22 Storage unit 23 Reconfiguration unit 41 Charge amplifier 42 Waveform shaping circuit 43 Comparator 44 Energy region distribution circuit 45 Counter 46 Weighting circuit 47 Addition circuit 112 Gamma ray detector 113 Data collection circuit 143 Comparator 144 Energy Area distribution circuit 145 Counter 146 Weighting circuit 147 Addition circuit S Semiconductor cell E1 Charge electrode E2 Current collection electrode

Claims (13)

A photon counting detector that has a plurality of collection pixels that allow radiation to enter and outputs the electrical signal corresponding to the energy of the particles by regarding the radiation incident on each of the collection pixels as a photon;
Area-specific data calculation means for calculating count data of the number of particles of the radiation classified into a plurality of energy areas on the energy spectrum of the radiation based on the signal of each collected pixel output from the detector;
Weighting means for applying the weighting coefficient given to each energy region to the count data of each of the plurality of energy regions for each collected pixel calculated by the region-specific data calculating unit;
And adding means for adding the count data of each of the plurality of energy regions for each collection pixel weighted by the weighting means to each other and outputting the added data as radiation image generation data for each collection pixel. Radiation detection device. The area-specific data calculation means discriminates a signal output from the detector into the plurality of energy areas, and counts the signal discriminated by the discrimination means for each energy area, and calculates the count value. The radiation detection apparatus according to claim 1, further comprising: counting means for passing the counting data to the weighting means. The weighting means is means for applying the weighting with a larger weighting coefficient as the count data obtained from an energy area on the higher energy side of the plurality of energy areas for each of the collected pixels. The radiation detection apparatus according to 1 or 2. The weighting means is means for applying the weighting with a larger weighting coefficient as the count data obtained from the energy region on the lower energy side of the plurality of energy regions for each of the collected pixels. The radiation detection apparatus according to 1 or 2. The weighting means is a means for applying the weighting with a larger weighting factor as the count data obtained from the high energy side energy region and the low energy side energy region of the plurality of energy regions for each of the collected pixels. The radiation detection apparatus according to claim 1 or 2. The weighting unit is a unit that applies the weighting with a smaller weighting coefficient as the count data obtained from the energy regions on the high energy side and the low energy side of the plurality of energy regions for each of the collected pixels. The radiation detection apparatus according to claim 1 or 2. The radiation according to any one of claims 1 to 6, wherein the detector is a one-dimensional detector or a two-dimensional detector in which the collection pixels are arranged in a one-dimensional shape or a two-dimensional shape. Detection device. The radiation detector according to any one of claims 1 to 7, wherein the detector is an X-ray detector that makes an X-ray as the radiation incident. An image generation means for generating a radiation image using the radiation detection apparatus according to any one of claims 1 to 8 and the radiation image generation data for each collection pixel output from the addition means of the radiation detection apparatus. And a radiological image diagnostic apparatus. The radiographic image diagnosis according to claim 9, wherein the image generation unit is a unit that generates a tomographic image of a subject scanned with the radiation by subjecting the radiographic image generation data to reconstruction processing of a desired algorithm. apparatus. The radiation detector according to any one of claims 1 to 7, wherein the detector is a gamma ray detector that makes gamma rays as the radiation incident. The radiographic image diagnosis apparatus according to claim 1, further comprising coefficient setting means for setting the weighting coefficient used by the weighting means to a value in accordance with a desired profile. The incident radiation is detected for each collection pixel, the incident radiation is regarded as particles, and an electrical signal corresponding to the energy of the particles is output,
Calculate the count data of the number of particles of the radiation classified into a plurality of energy regions on the energy spectrum of the radiation based on the output signal of each collected pixel,
Applying a weighting factor given to each energy region to the calculated count data of each of the plurality of energy regions for each collected pixel,
The weighted collection pixel count data for each of the plurality of energy regions is added to each other and the added data is output as radiation image generation data for each collection pixel.
A radiographic image generation method, wherein a radiographic image is generated using the radiographic image generation data.

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