Abstract:
A radiation detector emits a signal containing pulses produced by each radiation event being detected. The pulses tend to decay and may overlap the pulse from a subsequent radiation. In order to prevent such pulse overlap a circuit is provided to clip the signal pulses. The circuit incorporates an analog delay line which produces a delayed, inverted and attenuated reflection of the original detector signal. The reflection signal is combined with the original detector signal to cancel remnants of each pulse lasting longer than the predefined delay period.

Description:
BACKGROUND OF THE INVENTION 
     The present invention relates to nuclear medicine imaging systems, such as positron emission tomography (PET) scanners; and particularly to circuits for processing signals from radiation detectors in such systems. 
     Positrons are positively charged electrons which are emitted by radionuclides that have been prepared using a cyclotron or other device. These are employed as radioactive tracers called “radiopharmaceuticals” by incorporating them into substances, such as glucose or carbon dioxide. The radiopharmaceuticals are injected into a patient and become involved in such processes as blood flow, glucose metabolism, fatty acids, and protein synthesis. 
     Positrons are emitted as the radionuclides decay. The positrons travel a very short distance before they encounter an electron, and when this occurs, they are annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features which are pertinent to PET scanners—each gamma ray has an energy of 511 keV and the two gamma rays are directed in nearly opposite directions. An image is created by determining the number of such annihilation events at each location within the field of view. 
     The PET scanner includes one or more rings of detectors which encircle the patient. Each detector includes a scintillator which converts the energy of each 511 keV photon into a flash of light that is sensed by a photomultiplier tube (PMT). Coincidence detection circuits connect to the detectors and record only those photons which are detected simultaneously by two detectors located on opposite sides of the patient. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors. Within a few minutes hundreds of million of events are recorded to indicate the number of annihilations along lines joining pairs of detectors in the ring. These numbers are employed to reconstruct an image using well known computed tomography techniques. 
     Upon stimulation by a gamma ray, the scintillators do not emit light instantaneously, instead the light is emitted with an intensity that decays exponentially with time. The energy of the gamma ray is determined by integrating the single from the light sensor over the duration of the light pulse. The duration of that light pulse limits the rate in which the gamma rays can be detected and processed. With reference to FIG. 1 consider two gamma rays that interact with the scintillator at times T 1  and T 2 . The second gamma ray strikes the scintillator at time T 2  which occurs during the light pulse produced by the first gamma ray striking the scintillator at time T 1 . Thus, when the processing circuit integrates the signal from the photomultiplier the portion of the signal from the gamma ray which occurs after time T 2  (indicated by the crosshatched area on the drawing) will be integrated along with the signal produced by the second gamma ray striking the scintillator. Thus, the signals from the two gamma rays will “pile up” and not be processed correctly as individual gamma ray invents. Thus, it will appear as though the second gamma ray occurrence has a much greater signal. 
     Several methods have been developed to produce and correct the effect of pile up on the process signal. One such method is referred to as a variable integration with pulse tail extrapolation. In this method, the integration of the first pulse is stopped when the second pulse occurs. The integration value of the first pulse and the time between the first pulse and the second pulse is used to calculate a correction to the integrated value for the first gamma ray pulse because of the shorter integration time that is used. A correction for the portion of the first pulse is included with the integral of the second pulse. However, the circuitry required for this method is often too complex and expensive for practical application on imaging systems that contain a large number of light sensors and signal processing channels. 
     Another method of reducing pile up was described by E. Tanaka et al. entitled “Variable Sampling-Time For Improving Count Rate Performance Of Scintillator Detectors”,  Nuclear Instruments And Methods 158, 1989, pp. 459-466. This method generates a delayed, inverted and attenuated reflection of the original detector signal by means of analog delay line. The attenuation of the reflected pulse is chosen such that: 
     
       
         Ar(t)=−Ao(t−td) exp(−td/Tc) 
       
     
     where Ar(t) is the amplitude of the reflected pulse, Ao(t−Td) is the amplitude of the original signal, Td is the delay between the original signal and the reflected signal, and Tc is a decay time constant of the scintillator. The original and reflected pulses then are summed to give a clipped pulse as in an output signal as shown in FIG.  2 . The output signal is given by the expression: 
     
       
         As(t)=Ao(T)+Ar(T) 
       
     
     since Ao(T) is given by:                Ao        (   T   )       =   0           t   &lt;   0                 Ao        (   T   )       =       Ao        (     t   =   0     )            exp        (       -   t     /   Tc     )                 t   &gt;   0               then                 As        (   t   )       =   0           t   &lt;   0                 As        (   t   )       =     Ao        (   t   )               0   &lt;   t   &lt;   Td                 As        (   t   )       =   0           t   &lt;   Td                                
     FIGS. 3 and 4 show pre-amplifier circuits that use the delay line clipping method to shorten signals from a scintillator and photomultiplier tube (PMT). In FIG. 3, current from the photomultiplier is dropped across resistor RA to generate a voltage signal which is amplified by amplifier A 1  that acts as an input buffer for the clipping circuit. The clipping circuit consists of a load resistor RB, a delay line DL and a terminating resistor RC. The value of the load resistor RL is chosen to equal the characteristic impedance of the delay line DL and the value of resistor RC is chosen so that the reflected signal has the correct amplitude. A second amplifier A 2  functions as a driver to isolate the impedance of the output cable from the clipping circuit. Resistor RD is chosen to match the characteristic impedance of the output cable. This circuit has a drawback in that it requires two high performance amplifiers. 
     As an alternative, the circuit in FIG. 4 eliminates one of the amplifiers in the previous circuit by taking advantage of the high impedance output of the photomultiplier tube. The clipping circuit again consists of a load resistor RE, a delay line DL and its terminating resistor RF. The value of load resistor RE is chosen to equal the characteristic impedance of the delay line and the value of resistor RF is chosen so that the reflective signal has the correct amplitude. The current from the photomultiplier tube is dropped across the load resistance RE and the delay line which give an equivalent resistance of one-half RE. The sole amplifier A 1  amplifies the voltage signal and acts as a driver to isolate the impedance of the output cable. Although this latter utilizes only a signal amplifier, it has the drawback that the impedance used to convert the photomultiplier signal to a voltage depends upon the impedance of the delay line. If a standard low impedance delay line is employed, the gain required in the amplifier can become very large which may cause performance and stability problems. If the gain of the amplifier is kept low, a high impedance delay line must be used which are relatively large, have limited availability and are relatively expensive. 
     SUMMARY OF THE INVENTION 
     A circuit for clipping pulses in a signal from a radiation detector such as one that is incorporated in a nuclear medicine imaging system. That circuit has a first node connected to the radiation detector for receiving the signal, a second node, and a third node to which a reference voltage level is applied. In the preferred embodiment of the present invention the third node is coupled to circuit ground. An impedance element is connected between the first and second nodes to provide the desired total load to the PMT. As used herein, an impedance element is an electrical circuit component that has more than a negligible impedance, for example a conductor would not be considered as an impedance element. 
     A signal delay element, such as an analog delay line for example, has an input connected to the second node and has an output that is coupled to the third node. In the preferred embodiment, another impedance element couples the delay line output to the third node. A load element is connected between the second node and the third node, and has an impedance that matches the impedance of the signal delay element. An amplifier includes a signal input that is connected to the first node and includes a signal output at which is produced the output signal of the circuit for clipping pulses. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a graph depicting the signals produced by two gamma rays striking the detector in the PET scanner; 
     FIG. 2 graphically illustrates processing the detector signal by the delay line pulse shortening method; 
     FIG. 3 is one embodiment of a prior art delay line circuit for performing the pulse shortening; 
     FIG. 4 is another embodiment of a prior art pulse shortening circuit; 
     FIG. 5 is a schematic diagram of a PET scanner system in which the present invention is incorporated; 
     FIG. 6 is a pictorial view of a radiation detector which forms part of the PET scanner system of FIG. 5; and 
     FIG. 7 is a schematic diagram of a pulse shortening circuit according to the present invention. 
    
    
     DESCRIPTION OF THE PREFERRED EMBODIMENT 
     Referring to FIG. 5, the PET scanner system includes a gantry  10  which supports a detector ring assembly  11  about a central opening, or bore  12 . A patient to be examined is positioned in front of the gantry  10  and is aligned with the central axis of the bore  12 . A motorized patient table (not shown) moves the patient into the bore  12  in response to commands received from an operator work station  15 . A gantry controller  17  is mounted within the gantry  10  and responds to commands received from the operator work station  15  through a serial communication link  18  to operate the gantry. 
     With additional reference to FIG. 6, the detector ring  11  is comprised of 336 radiation detectors  20 . Each detector  20  includes a set of scintillator crystals  21  (referred to as BGO crystals) arranged in a matrix and disposed in front of four photomultiplier tubes  22  (abbreviated PMT). Each PMT  22  produces an analog signal on one of the conductors  23  when a scintillation event occurs. A set of acquisition circuits  25  are mounted within the gantry  10  to receive these signals and produce digital signals indicating the event coordinates (x,y) and the total energy. These digital signals are sent through a cable  26  to an event locator circuit  27  housed in a separate cabinet from the gantry. Each acquisition circuit  25  also produces an event detection pulse which indicates the exact moment the scintillation event took place. 
     The event locator circuits  27  form part of a data acquisition processor  30  which periodically processes the signals produced by the acquisition circuits  25 . The processor  30  has an acquisition CPU  29  which controls communications on the serial communication link  18  and a backplane bus  31 . The event locator circuits  27  assemble the information regarding each valid event into a set of digital numbers that indicate precisely when the event took place and the position of the BGO crystal  21  which detected the radiation event. This event data packet is conveyed to a coincidence detector  32  which also is part of data acquisition processor  30 . 
     The coincidence detector  32  accepts the event data packets from the event locators  27  and determines if any two of them are in coincidence. Coincidence is determined by a number of factors. First, the time markers in each event data packet must be within 12.5 nanoseconds of each other, and second, the locations indicated by the two event data packets must lie on a straight line which passes through the field of view in the scanner bore  12 . Events which cannot be paired are discarded, but coincident event pairs are located and recorded as a coincidence data packet that is conveyed through a serial link  33  to a sorter  34 . For a detailed description of the coincidence detector  32 , reference is made to U.S. Pat. No. 5,241,181 entitled “Coincidence Detector For A PET Scanner” which is incorporated herein by reference. 
     The sorter  34  forms part of an image reconstruction processor  40 . The sorter  34  counts all events occurring along each projection ray (R,θ) and organizes them into a two dimensional sinogram array  48  which is stored in a memory module  43 . The image reconstruction processor  40  also includes an image CPU  42  that controls a backplane bus  41  and links it to the serial communication link  18 . An array processor  45  also connects to the backplane  41  and it reconstructs images from the sinogram arrays  48 . The resulting image array  46  is stored in memory module  43  and is output by the image CPU  42  to the operator work station  15 . For a detailed description of the sorter  34 , reference is made to U.S. Pat. No. 5,272,343 entitled “Sorter For Coincidence timing Calibration In A PET Scanner” which is incorporated herein by reference. 
     The operator work station  15  includes a central processing unit (CPU)  50 , a cathode ray tube (CRT) monitor  51  and a keyboard  52 . The CPU  50  connects to the serial communication link  18  and it scans the keyboard  52  for input information. Through the keyboard  52  and associated control panel switches, the operator can control the calibration of the PET scanner, its configuration, and the positioning of the patient table for a scan. Similarly, the operator can display the resulting image on the CRT monitor  51  and perform image enhancement functions using programs executed by the work station CPU  50 . 
     With reference to FIG. 7, the input of each acquisition circuit  25  includes a separate pulse clipping circuit  60  that is connected to the output of one of the photomultiplier tubes  20 . This clipping circuit  60  utilizes a signal amplifier and allows the load resistance for the photomultiplier tube to be selected independently of the resistance of the delay line. 
     The clipping circuit  60  has a first, or input, node  62  to which the output of the photomultiplier tube  20  is connected. A first resistor R 1  and a load resistor RL are connected in series between the first node  62  and circuit ground with a second node  64  formed between the two resistors R 1  and RL. A signal delay element, such as an analog delay line  66 , has an input connected to the second node  64  and has an output connected by a second resistor R 2  to the circuit ground  65 . The first resistor R 1  serves as a impedance element which has an impedance that is chosen to produce the desired output voltage level from the photomultiplier tube. The load resistor RL has a value which matches the impedance of the delay line and second resistor R 2  has a value which is chosen so that reflected signal will have the proper amplitude. 
     The first, or input, node  62  also is connected to the input of an amplifier  68  with an output that is coupled by a third resistor R 3  to the output terminal  70  of the clipping circuit  60 . The resistance of the third resistor R 3  matches the characteristic impedance of the conductor  72  connected to output terminal  70 . 
     Operation of the clipping circuit  60  is best understood by analyzing the currents, as depicted in FIG. 7, which flow in response to a gamma ray event. For such analysis the input impedance of the amplifier  68  and the impedance of the photomultiplier tube  20  are treated as being infinite. The current signal from the PMT is designated as Ipmt (t) and is dropped across the first resistor R 1 . That current then divides into I 1  (t), which flows through the load resistor RL, and I 2 (t), which flows through the delay line  66 . Because the resistance of load resistor RL matches the characteristic impedance of the delay line  66 , equal currents flow through the load resistor and the delay line with each current being equal to one-half the PMT output current Ipmt (t). 
     The current I 2 (t−td) at the output of the delay line  66  is partially transmitted through the second resistor R 2  to circuit ground  65  with that current being referred to as It(t). The remainder of the current Ir(t) is reflected back through the delay line and after passing through the delay line the reflected current Ir(t−td) is dropped across the second resistor RL. 
     It should be noted that the input of the delay line is impedance matched so that neither I 2 (t) nor Ir(t−td) generates a reflective current at the input of the delay line. The amplitude of the reflected current Ir(t) is determined by the input mismatch at the output of the delay line and is given by the expression: 
     
       
         Ir(t)=(R 2 −RL)/(R 2 +RL)*I 2 (t−td) =½(R 2 −RL)/(R 2 +RL)*Ipmt (t−d) 
       
     
     For t&lt;2td, the input voltage to the amplifier is given by: 
     
       
         Vamp(t)=Ipmt (t)*R 1 +I 1 (t)*RL =Ipmt(t)*(R 1 +½RL) 
       
     
     For t&gt;2td, the input voltage to the amplifier is given by: 
     
       
         Vamp(t)=Ipmt(t)*R 1 +(I 1 (t)+Ir(t−td))*RL =Ipmt(t)*R 1 +(½Ipmt(t)+½Ipmt(t−2td)*(R 2 −RL)/(R 2 +RL))*RL =Ipmt(t)*(R 1 +½RL)+½Ipmt(t−2td)*(R 2 −RL)/(R 2 +RL))*RL 
       
     
     For Ipmt(t)=Ipmt(0)*exp(−t/Tc) and t&gt;2td, 
     
       
         Vamp(t)=Ipmt(0)*exp(−t/Tc)* (R 1 +½RL+½exp(2td/Tc)*(R 2 −RL)/(R 2 +RL)*RL) 
       
     
     The circuit will clip the PMT output signal for t&gt;2td if R 2  is chosen so that: 
     
       
         R 1 +½RL+½exp(2td/Tc)*(R 2 −RL)/(R 2 +RL)*RL=0 
       
     
     That is, 
     
       
         R 2 =RL*(RL−(2R 1 +RL)*exp(−2*td/Tc))/(RL+(2R 1 +RL)*exp(−2*td/Tc)) 
       
     
     EXAMPLE 
     To better understand the advantage of the present clipping circuit  60 , consider a PET scanner in which the PMT signal from a BGO detector (time decay constant Tc=300 nanoseconds) is required to be dropped over an effective impedance of 180 ohms and clipped after 400 nanoseconds. If the prior art circuit in FIG. 4 is employed, the third resistor RF will have a value of 210 ohms and the delay line needs to have an input impedance of 360 ohms. In the clipping circuit in FIG. 7 according to the present invention, these requirements can be met using a standard 100 ohm delay line and by selecting 130 ohms for the first resistor R 1 . The second resistor R 2  can be set to zero oh ms, in other words, the output of the delay line is coupled directly to circuit ground  65 . As used herein the term “coupled directly” means a connection that has negligible impedance. However, in order to ensure that the reflected signal has the proper amplitude, an impedance element, such as a resistor, may couple the output of the delay line to circuit ground  65 . In any event, the present invention enables the use of a 100 ohm delay line, which is more readily available and less expensive than a 360 ohm delay line required by the prior art circuit.