Signal processing system of PET-CT detector and PET-CT detector

Abstract

The invention provides a signal processing system of a PET-CT (positron emission tomography-computed tomography) detector and the PET-CT detector. The signal processing system comprises a front-end analog processing module and a time-energy dual-path module, the front-end analog processing module is connected with a weight resistance network of the PET-CT detector and comprises a main channel and an auxiliary channel, the main channel is used for amplifying weak current signals, and the auxiliary channel is used for compensating common-mode noise; and the time-energy dual-path module is connected with the front-end analog processing module and an ADC circuit control chip of the PET-CT detector, and is used for separating a signal timestamp and energy information, respectively carrying out parallel processing in a time path and an energy path, and inhibiting crosstalk between the paths. Compared with a traditional single channel, the noise density is effectively reduced, and inter-path crosstalk can be suppressed.

Description

Technical Field

[0001] This invention relates to the field of medical imaging technology, specifically to a signal processing system for a PET-CT detector and a PET-CT detector. Background Technology

[0002] PET-CT integrates positron emission tomography (PET) with computed tomography (CT). PET provides detailed molecular information on the function and metabolism of lesions, while CT provides precise anatomical localization of lesions. A single imaging session can obtain tomographic images of the whole body from all directions. It is characterized by its sensitivity, accuracy, specificity, and precise localization, allowing for a clear understanding of the overall condition of the body and achieving the goal of early detection of lesions and diagnosis of diseases.

[0003] However, with the continuous development of imaging equipment and the increasing precision of image imaging technology, the amount of data is also increasing. The following problems exist in the data transmission process: 1. Limited dynamic range: The count rate of a single-channel transimpedance amplifier is prone to saturation, easily leading to overload distortion and a decrease in energy resolution; 2. Insufficient common-mode interference suppression: High power supply / ground noise reduces the accuracy of test data and affects the detection of small lesions; 3. Multi-channel crosstalk: Crosstalk between adjacent channels in a 256-channel integrated system reduces spatial resolution. Therefore, traditional PET-CT requires a host computer to optimize the images in order to generate complete images. Summary of the Invention

[0004] In view of the deficiencies in the prior art, the purpose of this invention is to provide a signal processing system for a PET-CT detector and a PET-CT detector.

[0005] A signal processing system for a PET-CT detector according to the present invention includes: a front-end analog processing module and a time-energy dual-path module;

[0006] The front-end analog processing module is connected to the weighted resistor network of the PET-CT detector and includes a main channel and an auxiliary channel. The main channel is used to amplify weak current signals, and the auxiliary channel is used to compensate for common-mode noise.

[0007] The time-energy dual-path module connects the front-end analog processing module and the ADC circuit control chip of the PET-CT detector. It is used to separate the signal timestamp and energy information and process them in parallel in the time path and energy path respectively to suppress crosstalk between paths.

[0008] Furthermore, the main channel employs a folded common-source cascode transimpedance amplifier to directly couple the SiPM anode signal of the PET-CT detector, outputting a transimpedance signal.

[0009] Furthermore, the auxiliary channel uses a fully differential instrumentation amplifier to acquire the cathode reference signal and actively cancels common-mode noise through an inverting compensation circuit.

[0010] Furthermore, the time path records the instant when the gamma photon arrives at the PET-CT detector.

[0011] Furthermore, when a gamma photon strikes the scintillation crystal of the PET-CT detector and produces a flash, the SiPM anode of the PET-CT detector outputs a current signal, the rising edge of which corresponds to the initial time of the photon's arrival.

[0012] The time path detects the SiPM anode signal through a comparator. Once the voltage of the SiPM anode signal exceeds the threshold voltage, the comparator outputs a digital edge that jumps to a higher level. The digital edge is captured by a time-to-digital converter to obtain a digitized timestamp.

[0013] Furthermore, the comparator includes:

[0014] A diode D1 and a capacitor C1 are connected in series at the non-inverting input of the comparator. A resistor R1 connects the diode D1 and the capacitor C1 to ground. The capacitor C1 stores the peak voltage of the trans-impedance signal. The inverting input of the comparator is connected to the threshold voltage Vref.

[0015] When the peak voltage is greater than the threshold voltage Vref, the timestamp of the occurrence is recorded and transmitted to the ADC circuit control chip. The ADC circuit control chip sends a control signal to the weighted resistor network to shunt the signal, and distributes the portion of the signal that exceeds the main channel current / voltage threshold to the auxiliary channel.

[0016] Furthermore, the energy path is used to locate the impact position and energy of photons in the SiPM array of the PET-CT detector.

[0017] Furthermore, the independent voltage signals of each channel of the SiPM array enter their respective dynamic gain amplifiers. The amplification factor of the signal near the hit position decreases, while the amplification factor of the signal at the edge increases, so that all independent voltage signals reach a standardized intensity range. The voltage peak points of all independent voltage signals are then sent to the ADC for digitization.

[0018] Furthermore, each module of the signal processing system adopts at least four layers of boards, including: a high-voltage layer, an analog layer, a digital layer, and a GND layer, to isolate the analog power supply and the high-voltage power supply.

[0019] According to the present invention, a PET-CT detector includes a signal processing system for the PET-CT detector.

[0020] Compared with the prior art, the present invention has the following beneficial effects:

[0021] This invention employs transimpedance-differential dual-channel dynamic gain amplification, which effectively reduces noise density compared to traditional single-channel amplification at the same frequency, achieving noise suppression of 13dB, thus facilitating earlier detection of lesions. Simultaneously, the 13dB noise suppression equates to increased detection current sensitivity, reduced sample size requirements, and further minimization of sensor size, thereby reducing surgical incision size.

[0022] This invention employs a time-energy dual-path processing architecture, which simultaneously meets the requirements of high temporal resolution (ps level) and high energy resolution (keV level) while suppressing crosstalk between paths. Attached Figure Description

[0023] Other features, objects, and advantages of the present invention will become more apparent from the following detailed description of non-limiting embodiments with reference to the accompanying drawings:

[0024] Figure 1 This is a system architecture diagram for this application;

[0025] Figure 2 This is the schematic diagram of a comparator;

[0026] Figure 3 A schematic diagram showing a simulation comparison of single-channel and dual-channel noise density maps;

[0027] Figure 4 This is a schematic diagram of a single-channel TIA.

[0028] Figure 5 This is a schematic diagram of the differential auxiliary part in the dual-channel architecture of this application. Detailed Implementation

[0029] The present invention will now be described in detail with reference to specific embodiments. These embodiments will help those skilled in the art to further understand the present invention, but do not limit the invention in any way. It should be noted that those skilled in the art can make several changes and improvements without departing from the concept of the present invention. These all fall within the protection scope of the present invention.

[0030] like Figure 1 As shown, the present invention provides a signal processing system for a PET-CT detector, comprising: a front-end analog processing module and a time-energy dual-path module.

[0031] The front-end analog processing module is connected to the weight resistor network of the PET-CT detector, including a main channel and an auxiliary channel. The main channel is used to amplify weak current signals, and the auxiliary channel is used to compensate for common-mode noise. Main channel: A low-noise folded cascode TIA (Trans-impedance amplifier) is adopted to directly couple the SiPM anode signal. Auxiliary channel: Based on a fully differential instrumentation amplifier to collect the cathode reference signal, and actively cancel the common-mode noise through an inverting compensation circuit. The main / auxiliary channel seamless switching is achieved through a high-speed window comparator, and the dynamic range is extended to 80 dB.

[0032] The time-energy dual-path module is connected to the ADC circuit control chip of the front-end analog processing module and the PET-CT detector, and is used to separate the signal timestamp and energy information, and process them in parallel on the time path and the energy path respectively to suppress crosstalk between paths. The time path records the instant when the gamma photon arrives at the PET-CT detector. When the gamma photon hits the scintillation crystal of the PET-CT detector and generates a flash, the SiPM anode of the PET-CT detector will output a current signal, and the rising edge of the current signal corresponds to the starting moment when the photon arrives. The time path detects the SiPM anode signal through a comparator. Once the voltage of the SiPM anode signal exceeds the threshold voltage, the comparator will output a digital edge with a level jump, and the digital edge is captured through a time-to-digital converter to obtain a digitized timestamp.

[0033] As Figure 2 shown, the comparator includes: a diode D1 and a capacitor C1 connected in series to the non-inverting input terminal of the comparator. A resistor R1 is grounded between the diode D1 and the capacitor C1. The capacitor C1 stores the peak voltage of the transimpedance signal, and the inverting input terminal of the comparator is connected to the threshold voltage Vref.

[0034] Function of the C1 capacitor: Store the peak voltage of the transimpedance output signal, serve as the input signal of the comparator, charge through the diode D1, and discharge slowly through the R1 to maintain the peak level.

[0035] Function of Vref: Set the threshold voltage and compare it with the peak voltage at the capacitor C1 terminal.

[0036] Comparator logic:

[0037] When the voltage at the right end of the C1 capacitor > Vref, the comparator outputs a high level (trigger threshold);

[0038] When the voltage at the right end of the C1 capacitor < Vref, the comparator outputs a low level (not triggered).

[0039] Principle: As shown in the table below, the main channel (TIA) is responsible for high-gain amplification of weak current signals, while the auxiliary channel (differential amplifier) ​​compensates for common-mode noise. When a strong signal suddenly arrives in the main channel, such as a nearby photon or multiple photons arriving simultaneously, the main channel (TIA) will become inaccurate. The dual-channel design solves the problem of sudden strong signal arrival. It uses a low-sensitivity differential amplifier that does not directly measure the strength of the signal itself, but specifically senses background fluctuations and interference (common-mode noise). Through an inverting compensation circuit, the auxiliary channel generates an inverse voltage that is equal in magnitude and opposite in direction to the background fluctuations, thus reducing noise. The high-speed comparator has a threshold voltage VREF. When a strong signal exceeds VREF, the high-speed comparator records the timestamp of the event and sends it to the FPGA chip for control signal transmission. The control signal is then transmitted to the upper-level circuit of the transimpedance differential architecture. The weighted resistor network performs current shunting control on the strong signal, diverting part of the strong signal exceeding the current / voltage threshold of the main channel to the auxiliary channel. This reduces the amplification factor of the main channel and diverts the signal to the auxiliary channel. Within less than 5ns (the switching time of the high-speed comparator), the main channel (TIA) is restored to amplify the weak current signal with high gain.

[0040]

[0041]

[0042] Table 1 Examples of Main / Support Channel Switching

[0043] The traditional single-channel TIA and the dual-channel TIA of this application were simulated and verified using software. Single-channel TIA simulation: the input current was gradually increased from 1nA to the saturation point (e.g., 500nA), and the output voltage was recorded. Dual-channel simulation: under the same input range, the output noise after compensation of the auxiliary channel was observed; a noise density map was generated and compared with the noise spectral density (1Hz–10MHz). Figure 3 This is a simulation comparison diagram of noise density maps for single-channel and dual-channel systems. By comparing the noise density maps of the two schemes, it is clear that at the same frequency, the noise of the dual-channel system is much lower than that of the single-channel system.

[0044] parameter Single-channel TIA Dual-channel architecture Maximum distortion-free input current 500nA 1μA Equivalent Input Noise (RMS) 1.2mV 0.25mV Dynamic Range (DR) (Noise Suppression) 67dB 80dB

[0045] Table 2 Simulation results for single-channel and dual-channel simulations

[0046] Taking a frequency of 1Hz as the vertical axis as an example, the standard noise density for a single channel is 99 nV / Hz per 1Hz, while the optimized noise density for a dual-channel channel is 22 nV / Hz per 1Hz. This can be calculated using the power ratio formula.

[0047]

[0048] Converted to noise suppression:

[0049] Noise suppression (dB) = 10log 10 (20.25)≈13.06dB

[0050] Clinical manifestations:

[0051] For early screening of diseases, such as myocardial ischemia screening, 13dB noise suppression can detect lesions earlier, allowing for earlier drug intervention to avoid intervention being too late after the condition worsens.

[0052] This can lower the lower limit of the detection current and voltage. For example, traditional methods have poor ability to detect weak currents, requiring only a drop of blood. A 13dB noise suppression is equivalent to an increase in current detection sensitivity, requiring only a small amount of blood. This reduction in sample size means the sensor size can be reduced. For instance, in the implantation of electrodes for Parkinson's disease, traditional head drilling requires an 8mm diameter hole. With reduced noise and increased current detection sensitivity, a 0.3mm electrode sensor can be used. For patients, this results in a smaller surgical incision and longer recovery time.

[0053] Figure 4 This is a schematic diagram of a single-channel TIA. Rf is the feedback resistor, and capacitor C3 is used to prevent high-frequency oscillation.

[0054] Figure 5 This is the schematic diagram of the differential auxiliary section in the dual-channel architecture. Capacitors C1 and C2 and resistors R5 and R6 are used to filter out high-frequency noise and retain low-frequency common-mode interference. Resistors R3 and R4 are gain setting resistors and feedback resistors from the auxiliary channel output to the main channel.

[0055] The principle of the time-energy dual-path module is to achieve high-precision event determination by separating the timestamp and energy information of the signal. Its core objective is to simultaneously meet the requirements of high time resolution (ps level) and high energy resolution (keV level) by processing the paths in parallel, while suppressing crosstalk between paths.

[0056] The clinical manifestations are as follows:

[0057] 1. Improved Image Clarity and More Precise Lesion Identification: Imaging systems can more accurately "draw" real signal points. Like drawing dots on paper, the dots are rounder, clearer, and more accurately positioned. Doctors can more clearly identify where the tumor is, its size, and its shape, with less of the surrounding normal tissue obscuring the image. This is particularly crucial for the early detection of small tumors and the precise determination of tumor boundaries (important for radiotherapy or surgery).

[0058] 2. Faster scanning, or less developer used: The same clear examination can be completed in a faster time with less reagent (radiation).

[0059] 3. Better at distinguishing adjacent lesions: Doctors can see two close lesions more clearly.

[0060] The core task of the time path is to accurately record the instant when a gamma photon arrives at the detector. When a gamma photon strikes the scintillation crystal and produces a flash, the anode of the coupled silicon photomultiplier tube (SiPM) rapidly outputs a rapidly rising current signal. The key value of this signal lies in its steep rising edge, which directly corresponds to the start of the photon's arrival. The time path utilizes a high-speed comparator. It continuously monitors the anode signal, and once the signal voltage instantaneously exceeds a set threshold voltage (dynamically adjustable to accommodate noise), the comparator completes a state transition in an extremely short time, outputting a clean, precise digital edge transitioning from low to high. Immediately afterwards, a high-precision timer using a time-to-digital converter (TDC) captures this rising edge and converts it into a digital timestamp. This timestamp has extremely high resolution, reaching the picosecond level (one trillionth of a second), essentially converting the photon's arrival time into a precise number.

[0061] The core task of the energy path is to accurately locate the impact point of photons within the detector array and measure their energy. Pixels closer to the impact point have stronger signals, while those farther away have weaker signals. The independent voltage signals from all pixels enter their respective dynamic gain amplifiers in parallel, where the amplification factor of each channel is automatically adjusted: the amplification factor for strong signals near the impact point is automatically reduced, while the amplification factor for weak signals at the edges is increased, aiming to ensure that all signals ultimately reach a standardized intensity range. After amplification, all channel signals are simultaneously fed into a peak-hold circuit group to quickly capture and lock the voltage peak point of each channel. These peak points are then synchronously fed into a multi-channel high-speed ADC (analog-to-digital converter) for digitization.

[0062] In the signal processing system of the PET-CT detector, each functional module works together in strict accordance with electromagnetic compatibility design. Since the silicon photomultiplier tube (SiPM) requires a high voltage of 29V-32V, the entire system design board uses at least 4 layers, which are divided into a high voltage layer, an analog layer, a digital layer, and a GND layer.

[0063] For example, if the front-end analog processing module is a 6-layer board, and simultaneously handles high voltage, analog voltage, and digital voltage, the following layers would be used: the top layer is the analog signal layer, dedicated to analog signal routing; the second layer is the digital power layer, providing the main power supply for the chip; the third layer is the GND layer; the fourth layer is the power layer for analog devices; the fifth layer is the GND layer, primarily used to isolate analog power and high-voltage power; and the bottom layer is the high-voltage power layer. Based on the layered anti-crosstalk power supply principle, the 6-layer board design for the front-end analog processing module is shown in the table below:

[0064]

[0065]

[0066] Table 3 Layer PCB Design Table

[0067] Medium thickness control:

[0068] Spacing between high voltage layer (layer 6) and isolation ground (layer 5): 0.4mm (enhanced withstand voltage, breakdown voltage > 4kV). Spacing between analog layer (layer 4) and main ground layer (layer 3): 0.2mm (optimized decoupling capacitor loop inductance).

[0069] Copper foil specifications:

[0070] Power layer (layers 2 / 4): 2oz copper thickness (to reduce IR voltage drop)

[0071] Insulating layer (5th layer): 3oz copper thickness (to improve shielding effectiveness)

[0072] In the description of this application, it should be understood that the terms "upper", "lower", "front", "back", "left", "right", "vertical", "horizontal", "top", "bottom", "inner", "outer", etc., indicate the orientation or positional relationship based on the orientation or positional relationship shown in the accompanying drawings. They are only for the convenience of describing this application and simplifying the description, and do not indicate or imply that the device or element referred to must have a specific orientation, or be constructed and operated in a specific orientation. Therefore, they should not be construed as limitations on this application.

[0073] Specific embodiments of the present invention have been described above. It should be understood that the present invention is not limited to the specific embodiments described above, and those skilled in the art can make various changes or modifications within the scope of the claims, which do not affect the essence of the present invention. Unless otherwise specified, the embodiments and features described in this application can be arbitrarily combined with each other.

Claims

1. A signal processing system for a PET-CT detector, characterized in that, include: Front-end simulation processing module, time-energy dual-path module; The front-end analog processing module is connected to the weighted resistor network of the PET-CT detector and includes a main channel and an auxiliary channel. The main channel is used to amplify weak current signals, and the auxiliary channel is used to compensate for common-mode noise. The time-energy dual-path module connects the front-end analog processing module and the ADC circuit control chip of the PET-CT detector. It is used to separate the signal timestamp and energy information and process them in parallel in the time path and energy path respectively to suppress crosstalk between paths.

2. The signal processing system for the PET-CT detector according to claim 1, characterized in that, The main channel uses a folded common-source cascode transimpedance amplifier to directly couple the SiPM anode signal of the PET-CT detector and output the transimpedance signal.

3. The signal processing system for the PET-CT detector according to claim 1, characterized in that, The auxiliary channel uses a fully differential instrumentation amplifier to acquire the cathode reference signal and actively cancels common-mode noise through an inverting compensation circuit.

4. The signal processing system for the PET-CT detector according to claim 2, characterized in that, The time path records the instant when the gamma photon arrives at the PET-CT detector.

5. The signal processing system for the PET-CT detector according to claim 4, characterized in that, When a gamma photon strikes the scintillation crystal of a PET-CT detector and produces a flash, the SiPM anode of the PET-CT detector outputs a current signal, the rising edge of which corresponds to the initial time of the photon's arrival. The time path detects the SiPM anode signal through a comparator. Once the voltage of the SiPM anode signal exceeds the threshold voltage, the comparator outputs a digital edge that jumps to a higher level. The digital edge is captured by a time-to-digital converter to obtain a digitized timestamp.

6. The signal processing system for the PET-CT detector according to claim 4, characterized in that, The comparator includes: A diode D1 and a capacitor C1 are connected in series at the non-inverting input of the comparator. A resistor R1 connects the diode D1 and the capacitor C1 to ground. The capacitor C1 stores the peak voltage of the cross-impedance signal. The inverting input of the comparator is connected to the threshold voltage Vref. When the peak voltage is greater than the threshold voltage Vref, the timestamp of the occurrence is recorded and transmitted to the ADC circuit control chip. The ADC circuit control chip sends a control signal to the weighted resistor network to shunt the signal, and distributes the portion of the signal that exceeds the main channel current / voltage threshold to the auxiliary channel.

7. The signal processing system for the PET-CT detector according to claim 2, characterized in that, The energy path is used to locate the impact point and energy of photons in the SiPM array of the PET-CT detector.

8. The signal processing system for the PET-CT detector according to claim 7, characterized in that, The independent voltage signals of each channel of the SiPM array enter their respective dynamic gain amplifiers. The amplification factor of the signal near the hit point decreases, while the amplification factor of the signal at the edge increases, so that all independent voltage signals reach a standardized intensity range. The voltage peak points of all independent voltage signals are sent to the ADC for digitization.

9. The signal processing system for the PET-CT detector according to claim 1, characterized in that, Each module of the signal processing system uses at least four layers of boards, including: a high-voltage layer, an analog layer, a digital layer, and a GND layer, to isolate the analog power supply and the high-voltage power supply.

10. A PET-CT detector, characterized in that, The signal processing system includes the PET-CT detector according to any one of claims 1-9.

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