Abstract:
A gamma camera having a system for performing a quality control procedure with minimal to no intervention from a user of the camera. In one aspect, the gamma camera includes a relatively weak radioactive source positioned at a fixed or known location relative to the gamma camera scintillation crystal and positioned so that the entrance window side of the crystal is facing the source, wherein the photons emitted from the source have an energy that is below the energy of photons used for diagnostic imaging. The response of the gamma camera photo-multiplier tubes to the absorption events caused by the radioactive source when the camera is idle can be compared to a baseline response to determine whether one or more of the PMTs need to be adjusted.

Description:
BACKGROUND OF THE INVENTION 
     1. Field of the Invention 
     The present invention generally relates to a system and method for calibrating and tuning a gamma ray camera (“gamma camera”). 
     2. Discussion of the Background 
     Gamma cameras are primarily used by Doctors who specialize in the field of nuclear medicine. Nuclear medicine is a unique medical specialty wherein gamma cameras are used in conjunction with very low-level radioactive materials (called radionuclides or radiopharmaceuticals) to generate images of the anatomy of organs, bones or tissues of the body. Gamma cameras can also generate images that can be used to determine whether an organ is functioning properly. 
     The radionuclides or radiopharmaceuticals are introduced orally or intravenously into the body of a patient. Radiopharmaceuticals are specially formulated to collect temporarily in a certain part of the body to be studied, such as the patient&#39;s heart or brain. Once the radiopharmaceuticals reach the intended organ, they emit gamma rays that are then detected and measured by the gamma camera. The basic camera sold commercially for nuclear medical imaging is still similar to the original invention by Anger (U.S. Pat. No. 3,011,057, which is incorporated by reference herein). 
     A gamma camera includes a large area scintillation crystal, which functions as a gamma ray detector. The crystal is typically sodium iodide doped with a trace of thallium (NaI(Tl)). The crystal converts high-energy photons (e.g., gamma rays and X-rays) into visible light (i.e., lower energy photons). The crystal is positioned to receive a portion of the gamma ray emissions from the radiopharmaceuticals. 
     When a gamma ray strikes and is absorbed in the scintillation crystal, the energy of the gamma ray is converted into flashes of light (i.e., a large number of scintillation photons) that emanate from the point of the gamma ray&#39;s absorption in the scintillation crystal. A photo-multiplier tube (PMT), which is optically coupled to the scintillation crystal, detects a fraction of these scintillation photons and produces an output electronic signal (e.g., current or voltage pulse) having an amplitude that is proportional to the number of detected scintillation photons. The gamma ray camera typically has several photomultiplier tubes placed in a two dimensional array, with the signals from the different photomultiplier tubes being combined to provide an indication of the positions and energies of detected gamma rays. 
     The scintillation photons emitted from the detector crystal are typically in the visible light region of the electromagnetic spectrum (with a mean value of about 3 eV for NaI(Tl)). The scintillation photons spread out from the point of emission. A large fraction of the scintillation photons are transported from the point of emission to a light sensitive surface, called the photocathode, of the PMTs. A fraction of the scintillation photons incident on the photocathodes cause an electron to be emitted from the photocathode. 
     The electron, also called a photoelectron, is then electrostatically accelerated into an electron multiplying structure of the PMT, which causes an electrical current (or voltage) to be developed at an output of the PMT. The amplitude of the electrical signal is proportional to the number of photoelectrons generated in the PMT during the time period that scintillation photons are being emitted. Thus, after a gamma ray absorption event, the PMT outputs an electrical signal that can be used with other signals from other PMTs to determine the location of the gamma ray absorption event. 
     The number of scintillation photons producing electrical signals in each PMT is inversely related to the distance of the PMT from the point of gamma ray absorption, or event location. It is because of this relationship that the position of the event can be calculated from the signals of the PMTs surrounding the event location. 
     Ideally, the signal derived from each PMT should have exactly the same proportional relationship to the distance from the event location as for all other PMTs. The amplitudes of the signals derived from each PMT are proportional to two basic factors: 1) the number of scintillation photons detected by a PMT, and 2) the gain or amplification of the PMT. The accuracy to which the position of the event location can be calculated depends on these two factors remaining constant in time. 
     Typically, a gamma camera is tuned prior to its operation so as to ensure that the camera will calculate accurately the positions of event locations anywhere within an area called the field of view (FOV). Commercial, large field of view gamma cameras have between 50 and 100 PMTs. A tuning procedure will typically require a number of steps that balance or equalize the signal amplitudes of the PMTs. The gains of the PMTs are adjusted such that the sum of the signals from all the PMTs are approximately equal in response to a fixed energy gamma event, regardless of the location of the event. 
     A known pattern of event locations are presented to the camera, usually by placing a mask of precisely spaced lines or holes over the camera crystal, so that event location calculations can be calibrated to give the known locations fixed by the positions of the holes or slits, where the gammas can pass through the mask. The exact tuning and/or calibration steps may be different among cameras produced by different manufacturers. However, once the tuning and calibration steps are complete, the image quality, which is incumbent on the camera&#39;s ability to accurately position event locations, depends on the transport of scintillation light to the PMTs and the gains of the PMTs remaining unchanged from the time when the tuning and calibration procedures were performed. 
     A number of factors can cause a change in either the gain of a PMT or the light collection properties of the camera. PMT gain is a strong function of temperature, counting rate (i.e. the number event signals per unit time), and the high voltage (HV) power supply regulation. Additionally, PMTs change their gain over time as they age. The light collection from the crystal to the photocathodes of the PMTs can change if the transmissive properties of surfaces change. For example, the PMTs are optically coupled to a glass or plastic lightpipe using either an optical grease or epoxy. If any of these materials&#39; light transmission properties change, then the transport of scintillation photons to the PMT will change. Additionally, NaI(Tl) is a hygroscopic material, and if water vapor reaches the crystal it becomes yellow and the light transmission is diminished. 
     Different manufacturers have developed and implemented different means to maintain the constancy of PMT gains. These means fall into two categories: 1) automatic (i.e. not requiring the user to initiate the process), and 2) user quality control (QC) procedures (i.e. procedures initiated by the user). Generally, a combination of both automatic and QC procedures is required. 
     One automatic system, for example, utilizes light emitting diodes (LEDs) coupled into the photomultiplier tubes to provide a light signal for calibration of each individual tube. A constant fraction of the light emitted by the LED is incident on the light sensitive photocathode of the PMT. The PMT output signal is checked against a reference that was set at the time of the last calibration. The gain of the PMT is adjusted if the measured signal has strayed from the reference. 
     This gain calibration technique depends on the light emitting diodes having a constant light output for each pulse. Light emitting diodes, however, do not have constant light output as a function of temperature, and may also vary over the lifetime of the diode. Another drawback of this technique of automatic calibration is that the light from the diode is mostly directly incident on the photocathode of the photomultiplier tube. Therefore, the transport of the light through the scintillation crystal, and associated optical elements, is not significantly sampled by the pulse of light from the diode. 
     User initiated QC procedures usually require the placement of a radioactive source to uniformly illuminate the camera. The system acquires an appropriate number of events to achieve statistically significant sampling of each event location. A computer program then analyzes the measured energies and/or image of event locations to determine whether or not the system has drifted away from the properly calibrated state. Many variations of this procedure are possible, but all require the user to position a source of radioactivity and initiate the computer controlled acquisition and analysis. Additionally, the procedures also typically require the user to remove the collimator from the camera. 
     QC procedures are cumbersome to the user. If they can be initiated at the end of the day, and complete themselves automatically, then the user&#39;s time required is minimal. However, radioactive sources that must be left out in a room overnight require institutional procedures for securing the room, logging out the source and returning it in the morning, and prohibiting access to the room by cleaning and unauthorized personnel. Performing QC procedures during working hours reduces available patient imaging time on the system and increases costs because personnel are not doing patient imaging. 
     SUMMARY OF THE INVENTION 
     The above described and other disadvantages are addressed by the present invention through the use of a system and method that is designed to calibrate and tune a gamma camera with minimal or no human intervention. The present invention provides a valuable feature for the user in that the user is assured of optimal performance of the camera without requiring laborious procedures and time that might otherwise be devoted to patient imaging. 
     Advantageously, with the present invention, analysis of PMT output pulses and calibration can be totally automatic. First, in one embodiment, the user does not need to place a radioactive source because such sources can be made part of the camera. Second, the system computer can be programmed to monitor continuously the count rate and, thereby, determine when the system is being used and when the system is idle. When the system is idle (i.e., the count rate is approximately equal to the natural background plus the contribution of the radioactive sources), the system computer can automatically monitor and record individual PMT signals. When a sufficient number of data points have been stored for each PMT, the mean amplitude and variance of each tube&#39;s response to the events can be calculated. These calculated values may be compared to baseline values (e.g., values that were calculated at the time of the last tuning and calibration of the system, providing a database for comparison) and/or to calculated values associated with neighborhing PMTs to determine whether and to what extent adjustments to the camera need to made. Further, the software may be programmed to analyze the results of the comparisons and automatically make the necessary PMT gain adjustments. 
     The above and other features and advantages of the present invention, as well as the structure and operation of preferred embodiments of the present invention, are described in detail below with reference to the accompanying drawings. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     The accompanying drawings, which are incorporated herein and form part of the specification, illustrate various embodiments of the present invention and, together with the description, further serve to explain the principles of the invention and to enable a person skilled in the pertinent art to make and use the invention. In the drawings, like reference numbers indicate identical or functionally similar elements. Additionally, the left-most digit(s) of a reference number identifies the drawing in which the reference number first appears. 
     FIG. 1 is a diagram illustrating certain components of a gamma camera according to one embodiment of the present invention. 
     FIG. 2 is flow chart illustrating a process, according to one embodiment, for tuning the gamma camera. 
     FIG. 3 is a flow chart illustrating a data gathering process according to one embodiment. 
     FIG. 4 illustrates a data structure for storing the data gathered during the data gathering process. 
     FIG. 5 is flow illustrating a data analysis process according to one embodiment. 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     While the present invention may be embodied in many different forms, there is described herein in detail an illustrative embodiment(s) with the understanding that the present disclosure is to be considered as an example of the principles of the invention and is not intended to limit the invention to the illustrated embodiment(s). 
     FIG. 1 is a diagram illustrating certain components of a gamma camera  100  according to one embodiment of the present invention. As shown in FIG. 1, gamma camera  100  includes a scintillation crystal  102  (or “detector crystal  102 ”), a number of photomultiplier tubes (PMTs)  104 ( a ) . . . ( n ), and a computer system  110  coupled to the output of each PMT  104 . Advantageously, one or more very weak radioactive sources  106 ( a ) . . . ( n ) is placed so as to be facing an entrance window side  103  of scintillation crystal  102  at fixed or known locations. Gamma camera  100  may also include a collimator  114  and a light guide  116 . In one embodiment, sources  106  are positioned between collimator  114  and crystal  102 . 
     In one embodiment, sources  106  are positioned adjacent to the entrance window side  103  of scintillation crystal  102  at fixed or known locations. In a preferred embodiment, sources  106  are permanently or detachably affixed to entrance window side  103  of scintillation crystal  102  or to another component of camera  100 , such as collimator  114 . In a particular embodiment, a user of the camera  100  need not manually position sources  106  to occupy the fixed locations. For example, the sources may be pre-positioned and affixed to a component of camera  100  as part of the manufacturing process of the camera. 
     Sources  106  are chosen to have a photon energy that is below the source energies typical of diagnostic imaging, which are typically at least 140 keV. The source activity is also chosen to be below the limits set by regulatory agencies which would require licensing and inventory control. For example, Americium-241 (Am-241) emits a 60 keV X-ray and a long half-life. For activity levels less than 10 nCi,(nanocuries) such sources do not require radioactive material licenses. 
     Each radioactivity source  106 , which is placed in a fixed location, causes scintillation photons to emanate from a small region directly “below” the source whenever an X-ray from the source  106  enters crystal  102 . The scintillation photons produced by the X-rays will produce electronic signals of small amplitude in the photomultiplier tubes  104 . Since the source activity is small, the probability of two absorption events overlapping in time is of negligible consequence. 
     The scintillation photons generated from each absorption event can be assumed to be located at a known point in the crystal  102  because each source  106  is placed in a fixed location and the range of the low energy photons (i.e., X-rays) within the scintillation crystal is short (e.g., &lt;1 mm). Additionally, the mean number of scintillation photons produced from each X-ray absorption event will be near constant. Therefore, the signals produced in nearby PMTs, resultant from a number of scintillation photons generated from a single, monoenergetic X-ray absorption and subsequently transported to the PMTs, will be random statistical variants about constant means, modified by any changes in light transport and PMT response and amplification (i.e. gain). 
     Because the PMT output signal caused by one of the sources  106  should be a random statistical variant about a constant mean, absent changes in light collection and absent changes in the PMT itself, a process  200 , which is illustrated in FIG. 2, can be used to determine whether such changes have occurred and can be used tune PMTs  104  to compensate for the changes. Process  200  assumes a single source  106 , but multiple sources may be used. 
     Process  200  begins in step  202 , where PMTs  104  are tuned and calibrated to desired settings. Step  204  requires, for each PMT  104 , monitoring the output of the PMT for at least a certain period of time (e.g., one or more hours, but usually several hours) and, for each absorption event caused by source  106  during that period of time, recording the response of the PMT to the absorption event (i.e., recording the amplitude of the PMT output signal in response to the absorption event) so that a mean signal amplitude and a variance of the signal amplitude distribution for the PMT  104  can be calculated. 
     In step  206 , after a large number of events have been recorded for each PMT  104 , then, for each PMT  104 , a mean signal amplitude and a variance of the signal amplitude distribution are calculated based on the data recorded in step  204 . The mean signal amplitudes and signal amplitude distribution variances calculated in step  206  are the baseline means and variance values for PMTs  104 . In step  208 , the mean signal amplitudes and signal amplitude distribution variances calculated in step  206  may be stored in data storage unit  112 . 
     In step  210 , camera  100  may be used one or more times to image one or more patients. Like step  204 , step  212  requires, for each PMT  104 , monitoring the output of the PMT for at least a certain period of time and, for each absorption event caused by source  106  during that period of time, recording the response of the PMT to the absorption event. In step  214 , for each PMT  104 , mean and variance values are calculated based on the data acquired in step  212 . In step  216 , the mean signal amplitudes and signal amplitude distribution variances calculated in step  214  may be stored in data storage unit  112 . 
     In step  218 , for each PMT  104 , the mean and variance calculated in step  214  for the PMT are compared to the baseline mean and variance for the PMT, respectively, to determine whether there has been a change in mean amplitude or variance for the PMT. Additionally, in step  220 , the data collected in data storage unit  112  may be analyzed to determine if there are any data trends that may indicate changes in light collection efficiency. After step  220 , control may pass back to step  210  or proceed to step  222 . In step  222 , one or more PMTs  104  are re-tuned, if necessary, to compensate for changes, if any, in light collection efficiency and/or in PMT Gain. After step  222 , control may pass back to step  204  or  210 . 
     A change in the mean amplitude of a particular PMT  104  may be indicative of a gain change in the PMT or change in light collection efficiency. Increases in mean amplitude are almost certainly indicative of a gain change, as it unlikely that light collection could increase. Global changes in mean amplitudes of the PMT&#39;s could be due to high voltage supply drift. Changes in an individual PMT where the amplitudes of the individual PMT&#39;s neighboring PMTs remain constant, would indicate a gain change of a single a single PMT (i.e., the individual PMT). 
     Changes in light collection efficiency are likely to affect more than one PMT. Changes in light collection efficiency are also likely to occur slowly, over a long period of time (weeks to months). Measured changes in the PMTs&#39; responses are not likely to all be the same, but they will trend the same in time, so they should be recognizable as changes in light collection. 
     The variance of the amplitude distribution may be an indicator of light collection changes. Since the number of scintillation photons created by a single absorption event is a Poisson process, then the number of scintillation photons reaching a particular photocathode is also Poisson. If light collection does not change, then the variance in the distribution of acquired signal amplitudes should be predictable from random counting statistics. For example, if the mean number of scintillation photons reaching a photocathode is 100, then the standard deviation should be 10 (which is the square root of the variance which is equal to the mean for a Poisson distribution with a mean greater than 20) (Ref. G. F. Knoll, Radiation Detection and Measurement, 2 nd  Edition, John Wiley and Sons, 1989, pp. 74-75). Since the gain of a PMT has an extremely good signal to noise ratio, a gain change will shift the mean of the distribution of amplitudes, but should not change the standard deviation of the distribution, as measured relative to the mean. If, however, the number of scintillation photons reaching the PMT changes (i.e., there is a change in light collection efficiency), then the standard deviation of the distribution will change, as measured relative to the mean value. Changes in light collection of greater than 10% should be recognizable when signal distributions of statistical precision of about 1% are acquired and analyzed. 
     Referring now to FIG. 3, FIG. 3 is a flow chart illustrating a data gathering process  300 , according to one embodiment of the invention, for gathering data that is used in tuning and calibrating gamma camera  100 . Process  300  can be used to implement steps  204  and  212 . Process  300  assumes that there may be more than one source  106 . 
     Data gathering process  300  begins in step  302  where computer system  110  determines whether camera  100  is idle. Computer system  110  can determine whether camera  100  is idle by determining the count rate (e.g., the number of absorption events seen in a given period of time). The count rate is determined by monitoring the output of PMTs  104 . If the determined count rate is approximately equal to the count rate expected from natural background radiation plus the contribution of the sources  106 , then system  110  determines that camera  100  is idle. When camera  100  is determined to be idle, control passes to step  304 , otherwise control passes back to step  302 . 
     In step  304 , system  110  monitors the outputs of PMTs  104  and waits for an absorption event. System  110  can be programmed to determine when an absorption event occurs because the output of each PMT in a group of PMTs that is located in a neighborhood surrounding the area where the event took place change at or about the same time as a result of the event. When an absorption event occurs, for each affected PMT  104 , system  110  measures the magnitude of the PMT&#39;s output signal produced by the event (step  306 ). In step  308 , system  110  determines the source  106  that caused the event by determining the location of the absorption event. Because the sources  106  are in a known, fixed location, there is a direct correlation between the location of an absorption event and the source  106  that produced the event. 
     In step  310 , system  110  records the amplitude measurements taken in step  306 . That is, for example, for each measurement, system  110  stores in data storage unit  112  a value corresponding to the measured amplitude and associates the value with the PMT  104  from which the measurement was taken and the source  106  that was determined in step  308 . After step  310 , control passes back to step  302 . 
     FIG. 4 illustrates a possible schema for storing the acquired data. As shown in FIG. 4, a data table  402  is provided for each PMT  104 . From FIG. 4 one can determine, for each PMT, the magnitude of the output from the PMT from each recorded event from each source  106 . For example, one can see that the magnitude of the output from PMT  104 ( 1 ) caused by the second event from source  106 ( 2 ) has a value of 11. 
     Once data gathering process  300  has recorded a sufficient of amount of data (usually it takes at least a couple of hours to gather a sufficient amount of data), the process may end and a data analysis process  500  (see FIG. 5) may begin. FIG. 5 is a flow chart illustrating one embodiment of data analysis process  500 . Process  500  may be used to implement steps  214 - 218  of process  200 . 
     Data analysis process  500  begins in step  501 , where system  110  selects one of the PMTs  104 . In step  504 , system  110  selects one of the sources  106 . In step  506 , system  110  reads the recorded amplitude values associated with the selected PMT and the selected source. In step  508 , system  110  uses the values read in step  508  to calculate a mean value and a standard deviation. In step  510 , system  110  records in data storage unit  112  the values calculated in step  508 . In step  512 , system  110  determines whether there is another source  106  to select. If there is, system  110  selects the source and control passes back to step  506 , otherwise control passes to step  514 . In step  514 , system  110  determines whether there is another PMT  104  to select. If there is, system  110  selects another PMT and control passes back to step  504 , otherwise control passes to step  516 . In step  516 , system  110  compares the values recorded in step  510  to baseline measurements for the purpose of determining whether a change in gain and/or light collection efficiency has occurred. 
     It should be apparent to a designer of gamma camera tuning and calibration procedures that the data provided from the accumulation of PMT responses to the sources  106  is a sensitive measure of light collection and PMT gains. The fact that the sources  106  are low energy provide a particular advantage because licensing and inventory control are not required. Additionally, The fact that the sources may remain in the camera even when the camera is being used to image a patient is another advantage because this relieves the user from having to place the sources manually each time the user desires to initiate quality control procedures. Further, the user need not manually initiate the quality control procedures as the system  110  can be programmed when the camera is idle and automatically initiate the quality control procedures. These and other advantages provide significant improvement over existing procedures and methods. 
     While the invention has been described in detail above, the invention is not intended to be limited to the specific embodiments as described. It is evident that those skilled in the art may now make numerous uses and modifications of and departures from the specific embodiments described herein without departing from the inventive concepts.