KR101352771B1 - Coincidence circuit using shift register and PET data aquisition system, radiation counting system, medical diagnosis device using the same - Google Patents

Abstract

The present invention relates to a simultaneous counting circuit using a shift register, a PET data acquisition system, a radiation counting system, and a medical diagnostic device including the same. More specifically, a first data packet including a detection time of a first gamma ray signal is obtained. At least one shift register; A receiver which receives a second data packet including a detection time of a second gamma ray signal from an external data acquisition board and transmits the second data packet to the shift register; A comparator comparing the first gamma ray signal obtained from the first data packet and the second gamma ray signal obtained from the second data packet to determine whether the detection time of the first gamma ray signal and the second gamma ray signal is the same; ; A storage unit for storing a first data packet when the detection time of the first gamma ray signal and the second gamma ray signal is the same; A transmitter continuously transmitting the second data packet to another external data acquisition board disposed adjacent to the second data packet; .
With this configuration, the simultaneous counting circuit using the shift register of the present invention and the PET data acquisition system, radiation counting system and medical diagnostic apparatus including the same can detect the bottleneck of the data packet generated during transmission and storage. There is an effect that can be reduced through.

Description

Coincidence circuit using shift register and PET data aquisition system, radiation counting system, medical diagnosis device using the same}

The present invention relates to a simultaneous counting circuit using a shift register, and to a PET data acquisition system, a radiation counting system, and a medical diagnostic device including the same. In particular, a shift register capable of reducing bottlenecks generated during transmission and storage of a data packet is provided. The present invention relates to a simultaneous counting circuit, a PET data acquisition system, a radiation counting system, and a medical diagnostic apparatus including the same.

Positron emission tomography (PET) devices are commonly used to produce images of specific organs or tumors in the body, or to diagnose biochemical phenomena of active sites where metabolism occurs. Used to create an image. Such a PET device is a nuclear medical imaging apparatus for administering a radiopharmaceutical labeled with a positron emitting nuclide into a human body, and then detecting a gamma ray emitted outside the body to produce a tomographic image of the radiopharmaceutical distribution. At this time, radioisotopes such as C-11, N-13, O-15, and F-18, which are mainly used in the radiopharmaceutical, are decomposed into stable isotopes to emit positrons. In this case, the emitted positron is a particle having a positive charge with the same amount as the electron, but colliding with or surrounding the surrounding electrons to emit gamma rays at an angle of 180 degrees. Therefore, the detector of the PET device simultaneously detects the gamma rays detected in both directions to track the distribution of radiopharmaceuticals.

CT (Computed Tomography) or MRI (Magnetic Resonance Imaging), which are currently used in clinic, can provide anatomical images with high resolution. Is difficult to obtain. On the other hand, PET devices cannot provide anatomical images, such as CT or MRI, but because they can obtain biochemical information at the molecular level using various radiopharmaceuticals, metabolic, cancer and Alzheimer's disease , Parkinson's disease is widely used in the diagnosis.

In particular, such a PET device implements a co-counter to reduce the amount of data transmitted and stored. The co-counter detects and counts two or more gamma rays simultaneously within a specific time window (CTW), or generates simultaneous detection data.

Such co-counters are largely divided into co-counters using AND gates and co-counters using time stamps.

First, a co-counter using an AND gate passes a gamma ray signal detected through a detector to an AND gate to search for a case where a signal is simultaneously input. This method is useful because it is simple to implement and can be processed in real time, but as the number of output channels of the detector increases, the internal connection becomes more complicated, and the performance resolution of the individual logic gate elements degrades the time resolution of the PET device. As a result, a problem arises in that optimization is performed separately for each channel.

In contrast, a co-counter using a time stamp generates a detection time of gamma rays into a digital data packet and searches for the simultaneous detection through a logic circuit. The co-counter using this timestamp is divided into an event detection section for determining gamma-ray detection time, generating gamma-ray information as data packets, and a co-counter section for arranging different data packets to determine whether they are concurrent. In particular, the process of determining whether or not different data packets are simultaneously is performed by connecting a plurality of digital logic circuits in a tree structure, arranging the data packets in chronological order, and comparing the time information of the input data in the order of gamma ray detection at the last stage. Determine whether different data packets are concurrent. However, this process of determining whether the data packets are simultaneously requires multiplexing all the signals output from all the detector channels into a single line, which causes bottlenecks in the signal, and requires a separate parallel counting circuit including an aligner. A problem has occurred.

As described above, a simultaneous counting circuit using a shift register, a PET data acquisition system, a radiation counting system, and a medical diagnostic apparatus including the same will be described below.

Prior art 1 relates to Korean Patent Laid-Open No. 2011-0062622 (2011.06.10), which relates to a PET device and a signal processing method of the PET device. The prior art 1 is a pre-amplification unit for receiving the detected gamma ray signal amplified so that the rise time of the output waveform is less than 100ns; An ADC for converting the signal amplified by the preamplifier into a digital signal; And a signal processor that receives a digital signal and stores a signal having an energy value in which a photoelectric effect is generated among the input digital signals, thereby reducing the non-compliance time and increasing the count rate to improve the sensitivity of the PET device. Can provide.

In addition, the prior art 2, Korean Patent Publication No. 0160088 (1999.02.18), relates to the simultaneous selection circuit of the X-ray position detector. The prior art 2 is a monostable vibrator 21 to 23 at each output terminal of the discriminators D1 to D3 into which the signals S1 to S3 generated from two neighboring anode lines 2 of a one-dimensional MWPC (Multiwire Proportional Chamber) are input. And connect the input terminal of the AND gate AD1 (AD2) and the asynchronous counter 30 to the output terminal of the monostable oscillator 21, and the AND gate AD2 (AD3) to the output terminal of the monostable oscillator 22. Input terminal and asynchronous counter 31 are connected to the output terminal of the monostable oscillator 23 and the input terminal of the AND gate AD3 and the asynchronous counter 32, respectively, and the output terminals of the AND gates AD1 to AD3. Each monostable oscillator (24 to 26) is connected to each of the simultaneous counters (27 to 29) to solve the problem of increasing the concurrency time standard as the pressure of the detector body increases, and the circuit configuration is simple. As a result, the volume and cost are reduced, and the position of the X-ray position detector There is an effect that improves performance.

In order to solve the above problems of the prior art, the present invention is a simultaneous counting circuit using a shift register for easily detecting whether or not the gamma ray signal through the data packet transmitted on the FPGA, and PET data acquisition system including the same, radiation To provide a counting system and medical diagnostic equipment.

According to an aspect of the present invention, there is provided a simultaneous counting circuit using a shift register, including: at least one shift register for acquiring a first data packet including a detection time of a first gamma ray signal; A receiver which receives a second data packet including a detection time of a second gamma ray signal from an external data acquisition board and transmits the second data packet to the shift register; A buffer to temporarily store the first data packet until the shift register receives a second data packet; A comparator comparing the first data packet and the second data packet with each other to determine whether a detection time of the first gamma ray signal and the second gamma ray signal is the same; A storage unit for storing a first data packet when the detection time of the first gamma ray signal and the second gamma ray signal is the same; And a transmitter configured to continuously transmit the second data packet to another external data acquisition board disposed adjacent to the second data packet.

More preferably, when the detection time of the first gamma ray signal and the second gamma ray signal is the same, a shift register for activating an internal enable flag may be included.

More preferably, by checking whether the enable flag is activated, and when the enable flag is activated, it may include a storage unit for storing the first data packet.

In particular, it may include a shift register disposed in the form of First In First Out (FIFO).

More preferably, it may include a comparator for determining whether a gamma ray signal is simultaneously detected for data packets input within a predetermined time every clock cycle.

More preferably, when the detection time of the first gamma ray signal and the second gamma ray signal is not the same, the first data packet may be deleted.

In particular, it may include a data packet in the form of LMF (List Mode Format).

PET data acquisition system according to another embodiment of the present invention for solving the above problems is at least one shift register for obtaining a first data packet including a detection time of the first gamma ray signal, the second from the external data acquisition board A receiver for receiving and transmitting a second data packet including a detection time of a gamma ray signal to the shift register; a buffer for temporarily storing the first data packet until the shift register receives the second data packet; A comparator comparing the first data packet and the second data packet to determine whether the detection time of the first gamma ray signal and the second gamma ray signal is the same; and the detection time of the first gamma ray signal and the second gamma ray signal are the same. A storage unit for storing the first data packet, and other external data disposed adjacent to the second data packet. And a plurality of data acquisition boards including a simultaneous counting circuit using a shift register consisting of a transmission unit continuously transmitting to the acquisition board, wherein the data acquisition boards have the same direction and are arranged in a circular state to perform serial communication. It is characterized by.

Simultaneous counting circuit using the shift register of the present invention, PET data acquisition system, radiation counting system and medical diagnostic apparatus including the same can reduce the bottleneck of the data packet generated during transmission and storage through the simultaneous detection of the data packet It works.

In addition, the simultaneous counting circuit using the shift register of the present invention, and the PET data acquisition system, radiation counting system, and medical diagnostic apparatus including the same can easily implement the simultaneous counting circuit on the FPGA without adding a device configuration, thereby reducing manufacturing costs. It is effective to let.

In addition, the co-counter circuit using the shift register of the present invention, and the PET data acquisition system, radiation counting system and medical diagnostic apparatus including the same, even if the number of output channels of the detector for detecting gamma rays increases, the co-counter circuit replacement or data acquisition board There is no need to perform the upgrade, it is effective to improve the user's convenience.

1 is a block diagram of a simultaneous counting circuit using a shift register according to an embodiment of the present invention.
2 is a block diagram of a PET data acquisition system according to another embodiment of the present invention.
3 is a graph showing the data reduction rate of the simultaneous counting circuit having different time windows.
4 is a graph showing the co-efficient coefficient according to the radioactivity concentration.
5 is a view showing a PET image obtained by using the co-counter circuit of the present invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT Hereinafter, the present invention will be described in detail with reference to preferred embodiments and accompanying drawings, which will be easily understood by those skilled in the art. The present invention may, however, be embodied in many different forms and should not be construed as limited to the embodiments set forth herein.

Hereinafter, a simultaneous counting circuit using a shift register according to an exemplary embodiment of the present invention will be described in detail with reference to FIG. 1.

1 is a block diagram of a simultaneous counting circuit using a shift register according to an embodiment of the present invention.

As shown in FIG. 1, the co-counter circuit 130 using the shift register of the present invention includes a shift register 131, a receiver 136, a buffer 134, a comparator 133, a storage 132, and a transmitter. (135).

The shift register 131 acquires a first data packet including a detection time of a first gamma ray signal detected through a detector (not shown) in the data acquisition board, and receives the first data packet received by the receiver 136 from an external data acquisition board. Obtain a second gamma ray signal from the two data packets.

When the detection time of the first gamma ray signal and the second gamma ray signal is the same, the shift register 131 activates an internal enable flag, that is, displays the enable flag as '1'.

 In this case, the shift register 131 is arranged in the form of FIFO (First In First Out), so that a plurality of data packets received from an external data acquisition board are processed first in the order in which they are received.

The receiver 136 receives a second data packet including a detection time of a second gamma ray signal from an external data acquisition board and transmits the second data packet to the shift register 131.

The buffer 134 temporarily stores the first data packet until the shift register 131 receives the second data packet from the external data acquisition board. That is, in order to determine whether the first gamma ray signal and the second gamma ray signal are simultaneously detected, the buffer 134 is delayed for a predetermined time so that the comparison time between the first gamma ray signal and the second gamma ray signal becomes the same. Temporary storage is desirable.

The comparator 133 compares the first gamma ray signal and the second gamma ray signal respectively obtained from the first data packet and the second data packet to determine whether the detection time of the first gamma ray signal and the second gamma ray signal is the same. To judge. In this case, the comparator 133 determines whether the gamma ray signal is simultaneously detected only for data packets input within a predetermined time for each clock period of a field-programmable gate array (FPGA) system.

When the comparator 133 determines that the detection time of the first gamma ray signal and the second gamma ray signal is the same, the storage unit 132 stores the first data packet including the first gamma ray signal.

The transmitter 135 transmits the second data packet to another external data acquisition board disposed adjacent to the data acquisition board including the simultaneous counting circuit of the present invention, and is preferably transmitted continuously in one direction.

Alternatively, when the comparator 133 determines that the detection time of the first gamma ray signal and the second gamma ray signal is not the same, the shift register 131 deletes the first data packet.

In this case, the first and second data packets to be transmitted and stored are preferably in the form of List Mode Format (LMF).

Hereinafter, the PET data acquisition system according to another embodiment of the present invention will be described in detail with reference to FIG. 2.

2 is a block diagram of a PET data acquisition system according to another embodiment of the present invention.

As shown in FIG. 2, the PET data acquisition system 100 of the present invention includes a detector 110, a data acquisition board 150 including a simultaneous counting circuit, and an image processor 170.

The detector 110 detects a pair of gamma rays emitted from a body organ of a human body, and includes a scintillator and an optical sensor.

The scintillator converts the incident gamma ray into an optical signal when the gamma rays emitted from the human body are incident, and is mainly composed of sodium iodide doped with thallium (NaI (TI)), barium fluoride (BaF2), bismuth germanate (BGO), or luteium oxyorthosilicate (LSO). Etc. are used.

The optical sensor converts and outputs an optical signal generated from the scintillator into an electrical signal, and mainly includes a semiconductor photoelectric device such as a photomultiplier tube (PMT), an avalanche photodiode (APD), a silicon photomultipliers (SiPM), and a digital silicon photomultiplers (dSiPM). Used.

The detector 110 may be a semiconductor detector such as CZT (Cadmium zinc telluride) or CdTe (Cadmium telluride), or a gas detector such as RPC (Resistive Plate Chambers).

The data acquisition board 150 includes the co-counting circuit 130 using the shift register described above with reference to FIG. 1, and the description of the co-counting circuit will be omitted below. The data acquisition board 150 is arranged in a circular state to perform serial communication, so that data packets input from other data acquisition boards disposed adjacent to each other in one direction.

That is, the detection time of the first gamma ray signal included in the first data packet acquired by the data acquisition board 150 and the second data packet received from another data acquisition board disposed adjacent to the data acquisition board are included. When the detection time of the second gamma ray signal is compared with each other, the image processor 170 reconstructs a PET image of a body organ of the human body that emits a gamma ray signal based on the first data packet.

Hereinafter, an experiment of obtaining a PET image using the co-counter circuit of the present invention will be described in more detail.

First of all, the conditions of this experiment were based on a data acquisition board consisting of a 14-bit 100MHz free-running 8-channel ADC, Virtex-4 FPGA, 128 MB SDRAM, and digital input / output peripherals. In particular, IRI (Initial Rise Interpolation) method is used to calculate the arrival time of the gamma ray signal, and the initial rise value of the gamma ray pulse is included together to calculate the time more accurate than 10 ns, which is the sampling interval of the ADC. Obtained. In addition, an AUC (Area Under Curve) was used for the energy calculation. As a result, the time required for processing one gamma ray signal was about 60 ns for the baseline calculation and 340 ns for the energy calculation of the gamma ray pulse. It can be seen that it is 400ns.

In particular, as described with reference to FIG. 2, the data packet received by the data acquisition board or received from another data acquisition board disposed adjacent thereto is in the form of List Mode FOrmat (LMF), the arrival time information of the gamma ray signal, Contains the analog channel number. For example, as the arrival time information of the gamma ray signal is 48 bits and the analog channel number is allocated to 16 bits, it can be seen that the data packet has a total size of 64 bits. Since the size of the data packet is smaller than that of a conventional 128 bit data packet including the arrival time, reaction position, energy of the gamma ray signal, and an initial rising value for using the IRI method, the size of the present invention is the same. PET image acquisition process using the counting circuit can be seen that the dependency on the communication speed is lower than the conventional.

In particular, in this experiment, the data acquisition board is a serial communication connection using unidirectional serial communication input / output terminal of 1 Gbps. In this case, the data acquisition board uses a total of three pieces, and each board has six gamma rays It receives analog signal for and processes 18 channels in total.

At this time, as the co-counter circuit in the data acquisition board is configured as a FIFO form through a shift register implemented in a logic form in the FPGA instead of using an SDRAM, the shift register may have an access time equal to the clock frequency of the FPGA system. As a result, the access time is shorter than using the SDRAM. In addition, since the data can be output from the last shift register disposed at the last stage only after the data packet is input, it is possible to easily implement the wait delay for determining whether the co-count.

In addition, the co-counter circuit determines whether a data packet having the same detection time within a predetermined Coincidence Time Window is received every clock period, and converts the internal enable flag to '1'. In addition, each time the shift register receives the data packet, the storage device transmits the data packet continuously and continuously in a unidirectional direction to another shift register arranged adjacently, and only when the value of the enable flag is '1'. Transmits the data packet obtained by itself.

At this time, the data acquisition board waits for a predetermined time the data packet obtained by the data acquisition board until the next data packet is received, in order to determine whether the gamma ray signal in the data packet is simultaneously counted. At this time, the standby delay time may be less than or equal to the time (maximum delay time) until the data packet transmitted in the network environment having a constant communication speed cycles through the communication section of the daisy chain structure, and the following equation This can be confirmed by 1.

[Equation 1]

In this case, the T _{round - trip} is a standby delay time, the S _{packet size} is the _size of one data packet, the N _channel is the number of input channels of the data acquisition board, the N _{DAQ board} is the number of data acquisition boards, and the T _{network speed} is a constant communication _speed in the network environment.

Also, the buffer of the simultaneous counting circuit stores the first data packet until the second data packet transmitted from the external data acquisition board is received. Therefore, R, the number of shift registers included in the co-counter circuit, may be calculated as shown in Equation 2 below.

&Quot; (2) "

In this case, the N _channel is the number of input channels of the data acquisition board, the T _{round - trip} is the maximum waiting delay time before the data generated from another data acquisition board is transmitted to the final data acquisition board, the T _deadtime is The gamma ray detection data in each channel of the data acquisition board is the minimum time that can be continuously generated, that is, the data acquisition dead time.

In addition, the simultaneous coefficient circuit may include a coincidence coefficient module that detects a coincidence coefficient event simultaneously detected by two different pairs of extinctions to remove noise. The chance coefficient module delays the time window, thereby performing correction according to the chance coefficient.

That is, the virtual delay time is sequentially set in each detector module of the PET device. For example, in the case of a PET device having 18 detector modules, the virtual delay time is 0 ns for the first detector and the second detector. Set 20 ns and 40 ns for detector 3, and set the last 18 detectors to 340 ns. In addition, when a single gamma ray detection event occurs, a virtual delay time is added to the gamma ray arrival time. As such, when the gamma ray detection events are rearranged based on the arrival time of the gamma ray plus the delay time, a pair of gamma rays simultaneously detected in a predetermined time window is estimated as a coincidence coefficient. These estimates are used to calculate the probability of occurrence of chance coefficients in LOR (Line Of Response) to perform the correction.

In this experiment, as shown in Figure 3, as a basic experiment for the performance evaluation of the data acquisition system of the present invention, the data reduction rate according to the radiation concentration and the time window of the co-counter was obtained.

3 is a graph showing the data reduction rate of the simultaneous counting circuit having different time windows.

First, in the case of a data acquisition system using a simultaneous counting circuit, data having a size smaller than the size of data outputted from a data acquisition system not using the simultaneous counting circuit should be output. In addition, the data reduction rate according to the change in the radioactivity concentration was measured by the data acquisition method having time windows of 30 ns, 60 ns, and 90 ns, respectively.

Prior to performing the experiment, a solution of F-18 diluted in a cylindrical phantom was added and the data reduction rate was taken for 1 to 3 minutes once every 30 minutes until the radioactivity concentration was reduced from 4.6 mCi to 0.5 mCi. The data consisted of a 128-bit LMF format containing the time of arrival, reaction location, energy, and initial rise for the IRI method for one gamma ray signal detection.

After acquiring the data using these two data acquisition methods, data corresponding to one minute after the start of photographing was cut out to calculate a data reduction rate as shown in Equation 3 below.

&Quot; (3) "

The reduction rate is a data reduction rate, the Size of output data from proposed method is the size of a data packet obtained from a data acquisition system including a co-counter circuit of the present invention, and the Size of output data from conventional method is This is the size of a data packet obtained from a data acquisition system that does not include a co-coefficient circuit.

As shown in FIG. 3, the data reduction rate when using a co-counter having a time window of 30 ns, 60 ns, 90 ns at a radiation concentration of 0.5 mCi to 4.7 mCi can be seen. As a result, when the radioactivity concentration is high, the coincidence coefficient noise increases, and thus the data reduction rate decreases as the radioactivity concentration increases.

In addition, since the coincidence coefficient noise decreases as the time window of the co-counter circuit becomes shorter, it can be seen that the data reduction rate is higher than that of the co-counter circuit with a long time window. For example, in the case of the co-efficient circuit using a time window of 90 ns, as the radiation concentration increases, the data reduction rate decreases, showing only about 50% data reduction rate at 3 mCi, In the case of 4 mCi, the data reduction rate was over 75%.

In other words, at a concentration of 1 mCi, it can be seen that the data reduction rates are 65% at 90 ns time window, 73% at 60 ns, and 81% at 30 ns, respectively.

In addition, an experiment was performed to determine the co-efficient of the radiation concentration.

4 is a graph showing the co-efficient coefficient according to the radioactivity concentration.

First, in order to confirm the co-efficient coefficient according to the radioactivity concentration, an experiment was conducted to compare the data acquisition system according to the present invention with the data acquisition system of the existing tree structure using 18 analog signals output from the brain imaging PET. First, the conventional data acquisition system and the data acquisition system of the present invention depend on the use of the simultaneous counting circuit.

As shown in Fig. 4, the dots indicated by the circle indicate the co-efficient coefficient of the data acquisition system without using the co-counter circuit, and the dots indicated by the triangles indicate the co-count of the data acquisition system using the co-counter circuit according to the present invention. Indicates.

First, in order to measure the difference in the co-efficient of the data acquisition system according to the use of the co-coefficient circuit, in this experiment, the F-18-diluted solution was filled in a cylindrical phantom, and the data was once every 30 minutes according to the decrease in the radioactivity concentration. Obtained. The reordering program was used to calculate the co-efficient and chance coefficients from the acquired data. At this time, the energy window applied to the rearrangement program was about 350 keV to 750 keV, the time window was 20 ns, and the delay time of the delayed window used for calculating the chance coefficient was set to 20 ns.

In the present experiment, if the co-counter circuit using the shift register has a loss in the co-coefficient due to an algorithm error, the co-count will decrease. However, as shown in FIG. 4, if there is no error in the co-coefficient algorithm, the co-counter It can be seen that the coefficient of coherence is similar to that of a data acquisition system that does not use.

This co-count represents the count rate calculated by the rearrangement program for data obtained once every 30 minutes from 5.2 mCi to 0.5 mCi. When the data is obtained, there is a slight difference in the radioactivity concentration, but the method using the co-counter circuit of the present invention shows that there is no count loss, as the count rate is similar to that of the conventional method.

Hereinafter, the reconstruction of the PET image will be described in detail with reference to FIG. 5.

5 is a view showing a PET image obtained by using the co-counter circuit of the present invention.

In the present invention, after taking a Hot-rod model using a brain imaging PET, data was obtained. If there is no abnormality in the operation of the co-counter circuit, tomographic images similar to those of the system using the co-counter circuit and the system without the co-counter circuit are taken.

Figure 5 (a) shows a hot-rod model for the verification of the present experiment, Figure 5 (b) is obtained from a data acquisition system using the image obtained from the data acquisition system using the co-coefficient circuit of the present invention FIG. 5C is a reconstructed image obtained from a data acquisition system that does not use a conventional simultaneous counting circuit.

As shown in Figure 5 (a), the internal structure diameter of the hot-rod model is 6.5 mm, 5.5 mm, 4.5 mm, 3.5 mm, 2.5 mm rod size is divided into five compartments, respectively, the inner diameter Its size consists of a 75 mm cylinder. At this time, a solution diluted with 1 mCi of isotope F-18 (FDG) was filled in the cylinder and data was obtained for 10 minutes. This experiment was divided into a data acquisition system with a co-counter circuit and a data acquisition system without a co-counter circuit. The energy window with the data rearrangement is from 350 keV to 750 keV and the time window is 20 ns. The image reconstruction method used 2D FBP (filtered backprojection), and did not undergo normalization and chance coefficient correction.

5 (b) to 5 (c), even though the normalization and chance coefficient correction is not performed, both images can be seen through the naked eye 6.5 mm, 5.5 mm, 4.5 mm rod, In addition, the 3.5mm rod also contains some noise but can be distinguished.

In addition, the number of shift registers used in the simultaneous counting circuit of the present invention can be determined.

Previously, the maximum delay time of the data packet through Equation 1 is taken as an example. In the system that uses three 6-channel data acquisition boards at a communication speed of 1 Gbps and sends 64 bit data packets per packet, the packet is sent to the final data acquisition board. The maximum delay time before this transmission is 768 ns, as shown in Equation 4 below.

&Quot; (4) "

In addition, when the maximum delay time 768ns and the system dead time 400ns are applied to Equation 2 above, the number of shift registers R required for the configuration of the simultaneous counting circuit is 12 as shown in Equation 5 below. Able to know.

&Quot; (5) "

In addition, in order to correct according to the chance coefficient event, the time information input to the chance coefficient module should add the delay time of the delayed window to the arrival time of the gamma ray signal. Therefore, the maximum transfer time for determining the number of shift registers required for the coincidence coefficient module is 360 ns maximum delay time of the delayed window used for calculating the coincidence coefficient (18 analog channels × Delayed window 20 ns) as shown in Equation 6 below. You can see that we add 1) to 1128 ns.

&Quot; (6) "

Substituting this maximum delay time into Equation 2, it can be seen that R, which is the number of shift registers required for the coincidence coefficient module, becomes 17 (16.92), as shown in Equation 7 below.

[Equation 7]

In addition, the co-counter circuit as described above can be implemented not only in the PET data acquisition system but also in the radiation counting system or the medical diagnostic apparatus by including the co-counter circuit described in the present invention.

Simultaneous counting circuit using the shift register of the present invention, PET data acquisition system, radiation counting system and medical diagnostic apparatus including the same can reduce the bottleneck of the data packet generated during transmission and storage through the simultaneous detection of the data packet It works.

In addition, the simultaneous counting circuit using the shift register of the present invention, and the PET data acquisition system, radiation counting system, and medical diagnostic apparatus including the same can easily implement the simultaneous counting circuit on the FPGA without adding a device configuration, thereby reducing manufacturing costs. It is effective to let.

In addition, the simultaneous counting circuit using the shift register and the PET data acquisition system, the radiation counting system, and the medical diagnostic apparatus including the same simplify the internal configuration regardless of the increase in the number of output channels of the detector for detecting gamma rays. Since the optimization process for each channel is not performed due to the difference in performance, the time required for simultaneous coefficient detection can be reduced.

While the present invention has been described in connection with what is presently considered to be practical exemplary embodiments, it is to be understood that the invention is not limited to the disclosed embodiments, but, on the contrary, Do.

131: shift register 132: storage unit
133: comparator 134: buffer
135: transmitter 136: receiver

Claims (10)

At least one shift register for acquiring a first data packet including a detection time of the first gamma ray signal;
A receiver which receives a second data packet including a detection time of a second gamma ray signal from an external data acquisition board and transmits the second data packet to the shift register;
A buffer to temporarily store the first data packet until the shift register receives a second data packet;
A comparator comparing the first gamma ray signal obtained from the first data packet and the second gamma ray signal obtained from the second data packet to determine whether the detection time of the first gamma ray signal and the second gamma ray signal is the same; ;
A storage unit for storing a first data packet when the detection time of the first gamma ray signal and the second gamma ray signal is the same; And
A transmitter continuously transmitting the second data packet to another external data acquisition board disposed adjacent to the second data packet;
Simultaneous counting circuit using a shift register comprising a.
The method of claim 1,
The shift register
And the internal enable flag is activated when the detection time of the first gamma ray signal and the second gamma ray signal is the same.
3. The method of claim 2,
The storage unit
Checking whether the enable flag is activated, and when the enable flag is activated, storing the first data packet.
The method of claim 1,
The shift register
A co-counter circuit using a shift register, characterized in that arranged in the first in first out (FIFO) form.
The method of claim 1,
The comparator
A co-counter circuit using a shift register, characterized in that for determining the simultaneous detection of the gamma ray signal for the data packets input within a predetermined time for each clock period.
The method of claim 1,
And the first data packet is deleted if the detection time of the first gamma ray signal and the second gamma ray signal is not the same.
The method of claim 1,
The first and second data packets
A simultaneous counting circuit using a shift register, characterized in that it is in the form of a List Mode Format (LMF).
A radiation counting system comprising the simultaneous counting circuit according to any one of claims 1 to 7.
A medical diagnostic device comprising the simultaneous counting circuit according to any one of claims 1 to 7.
At least one shift register for acquiring a first data packet including a detection time of a first gamma ray signal, and a second data packet including a detection time of a second gamma ray signal from an external data acquisition board is received and transmitted to the shift register. A receiving unit, a buffer for temporarily storing the first data packet until the shift register receives the second data packet, a first gamma ray signal obtained from the first data packet, and a second data packet obtained from the second data packet. A comparator comparing two gamma-ray signals with each other to determine whether the detection time of the first gamma-ray signal and the second gamma-ray signal are the same; and if the detection time of the first gamma-ray signal and the second gamma-ray signal is the same, A storage unit for storing the second data packet continuously to another external data acquisition board disposed adjacent to the second data packet; And a plurality of data acquisition boards including a simultaneous counting circuit using a shift register formed by a transmitting unit.
The data acquisition board is
PET data acquisition system having the same direction and arranged in a circular state to perform serial communication.

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