Radiant energy – Invisible radiant energy responsive electric signalling – With or including a luminophor
Reexamination Certificate
1998-11-06
2001-05-15
Hannaher, Constantine (Department: 2878)
Radiant energy
Invisible radiant energy responsive electric signalling
With or including a luminophor
C250S363040, C250S369000
Reexamination Certificate
active
06232604
ABSTRACT:
BACKGROUND OF THE INVENTION
The present invention relates to PET scanners generally and specifically to a method and apparatus for modifying time stamps of detected photons to accurately mirror temporal relationships between photons which are detected despite hardware variations.
Positrons are positively charged electrons which are emitted by radionuclides which have been prepared using a cyclotron or other device. The radionuclides most often employed in diagnostic imaging are fluorine-18 (
18
F), carbon-11 (
11
C), nitrogen-13 (
13
N), and oxygen-15 (
15
O). Radionuclides are employed as radioactive tracers called “radiopharmaceuticals” by incorporating them into substances such as glucose or carbon dioxide. One common use for radiopharmaceuticals is in the medical imaging field.
To use a radiopharmaceutical in imaging, the radiopharmaceutical is injected into a patient and accumulates in an organ, vessel or the like, which is to be imaged. It is known that specific radiopharmaceuticals become concentrated within certain organs or, in the case of a vessel, that specific radiopharmeceuticals will not be absorbed by a vessel wall. The process of concentrating often involves processes such as glucose metabolism, fatty acid metabolism and protein synthesis. Hereinafter, in the interest of simplifying this explanation, an organ to be imaged will be referred to generally as an “organ of interest” and prior art and the invention will be described with respect to a hypothetical organ of interest.
After a radiopharmaceutical becomes concentrated within an organ of interest and while the radionuclides decay, the radionuclides emit positrons. The positrons travel a very short distance before they encounter an electron and, when the positron encounters an electron, the positron is annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features which are pertinent to medical imaging and particularly to medical imaging using photon emission tomography (PET). First, each gamma ray has an energy of essentially 511 keV upon annihilation. Second, the two gamma rays are directed in substantially opposite directions.
In PET imaging, if the general locations of annihilations can be identified in three dimensions, the shape of an organ of interest can be reconstructed for observation. To detect annihilation locations, a PET scanner is employed. An exemplary PET scanner includes a plurality of detector modules and a processor which, among other things, includes coincidence detection circuitry. An exemplary detector module includes six adjacent detector units. An exemplary detector unit includes an array of crystals (e.g. 36) and a plurality of photo-multiplier tubes (PMTs). The crystal array is located adjacent to the PMT detecting surfaces. When a photon impacts a crystal, the crystal generates light which is detected by the PMTs. The PMT signal intensities are combined and the combined signal is compared to a threshold (e.g. 100 keV). When the combined signal is above the threshold, an event detection pulse (EDP) is generated which is provided to the processor coincidence circuitry. Other hardware determines which crystal generated the light (i.e. absorbed the photon).
The coincidence circuitry identifies essentially simultaneous EDP pairs which correspond to crystals which are generally on opposite sides of the imaging area. Thus, a simultaneous pulse pair indicates that an annihilation has occurred on a straight line between an associated pair of crystals. Over an acquisition period of a few minutes, millions of annihilations are recorded, each annihilation associated with a unique crystal pair. After an acquisition period, recorded annihilation data is used via any of several different well known procedures to construct a three dimensional image of the organ of interest.
Depending on its design a PET scanner may test the energy level before or after testing for coincidence timing and the coincidence timing test may be either analog or digital. In a typical analog coincidence circuit the duration of a timing signal is set to a pre-determined value (e.g. W/2 where W is a time period corresponding to a coincidence window). The timing signals from the detector units are then combined using a conventional AND logic gate which produces an output only when two timing pulses overlap (i.e. two consecutive pulses are within ±W/2).
In a typical digital coincidence circuit each timing signal is compared to a master clock signal in a time to digital converter (TDC). The time stamp digital value from the TDC corresponds to the time lapsed between the previous master clock pulse and the EDP. After each master clock cycle the time stamps of all photons detected during the completed master clock cycle (i.e. the stamps which occurred between the preceding two master clock pulses) are compared. Photons which have time stamp differences between the time stamps of smaller than ±W/2 are identified as coincidence pairs.
During an acquisition period there are several sources of annihilation detection error. Two of the more prominent sources of detection error are referred to as “dead time” and “randoms”. The phenomenon known as dead time occurs when two photons impact a single crystal at essentially the same time (e.g. within the same time period W) so that the total absorbed energy far exceeds 511 keV or so that, while a first of the photons is being processed by the detector unit, a second of the photons is ignored by the unit. In these cases, at least one of the annihilations is not recognized and an error occurs.
The phenomenon known as randoms occurs when photons from two different annihilations are detected by two crystal at essentially the same time. For example, assuming two photons are detected at the same time from two different annihilations, the coincidence circuitry cannot determine which detection correspond to a first annihilation and which detection correspond to a second annihilation.
In order to minimize the number of random coincidences and the effects of deadtime, the size of the coincidence window W should be as small as feasibly possible. Unfortunately there are three sources of timing error which limit how small window W may be. First, because photons travel at the finite speed of light, photons produced at different locations within a scanner's field of view (FOV) reach crystals at different times. For example, if an annihilation event occurs eight inches from one crystal and sixteen inches from a second crystal it will take twice as long for a photon generated thereby to reach the second crystal than for a photon to reach the first crystal.
Second, the statistical nature of the photon detection process in a detector unit and electronic noise in scanner circuits cause variations among events detected by a detector unit.
Third, because there are slight differences in detector hardware the period between an impact time and an EDP can vary between detector units. In addition, even within a detector unit, the period between impact time and a corresponding EDP can vary as a function of where across a detector surface area a photon is absorbed.
While very little can be done to compensate for different flight times and noise, the industry has devised schemes to deal with varying hardware transit times. For example, according to one scheme the transit time differences among detector units are minimized by sorting the detector units so that detectors within each scanner have similar transit times.
According to another scheme, delay lines of selected values are added to the signal paths of certain detector units to even out the transit times. One way to select delay line values is to select delay line lengths known to cause specific delays.
Another way to select delay values is to provide a separate programmable delay for each detector unit. Prior to using the scanner to generate imaging data, a calibration procedure is used to set each programmable delay to eliminate the transient time differences. Preferably, both hardware select
McDaniel David L.
Scarlett James E.
Cabou Christan G.
Gabor Otilia
General Electric Company
Hannaher Constantine
Quarles & Brady LLP
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