Positron emission tomography (PET) is another nuclear medicine imaging method that has several advantages over SPECT. PET uses positron-emitting radionuclides that result in the emission of hillar pairs of 511 keV annihilation photons. Coincidence detection of the annihilation photons avoids the need for collimation and makes PET much more efficient than SPECT for detecting radioactivity. Even more important, positron-emitting radionuclides exist for oxygen, carbon, nitrogen, and fluorine, allowing a wide range of molecules to be labeled as diagnostic agents. Many of these radionuclides have short half-lives and require an on-site cyclotron. However, 18F has a long enough half-life that it can be (and is) provided regionally, and there is not a populated area of the United States where it is not available. Several others, such as 82Rb and 68Ga, are available from radionuclide generators that provide the radionuclides on demand despite their short half-life.
Coincidence detection provides spatial resolution without the need for lead collimation by taking advantage of the fact that the annihilation photons resulting from positron emission are approximately collinear. Events are only counted if two opposing detectors detect them simultaneously. The sensitive volume defined by the coincidence detectors is called answer line (LOR). Two individual detection systems are used with an additional matching module. Each individual system will generate a logic pulse when it detects an event that falls within the selected energy window. If the two logic pulses overlap in time in the coincidence module, a coincidence event is logged. PET systems use a large number (>10,000) of detectors arranged as multiple rings to form a cylinder. Since any detector can coincide with other detectors in the cylinder, the resulting LORs provide enough sampling to collect the projection information required for the tomography.
The intrinsic detection efficiency of an individual detector depends on the atomic number, density, and thickness of the detector. Ideally, the intrinsic detection efficiency should be 1, but at 511 keV that is difficult to achieve, although the intrinsic efficiency for some of the detectors is higher than 0.8. Match detection requires both listeners to register an event. Since the interactions in the two detectors are independent, the intrinsic efficiency of the coincidence depends on the product of the intrinsic efficiency in each detector. As a result, the efficiency of coincidence detection is always less than that of a single detector, and that difference is magnified for low-efficiency detectors. Due to the need for high intrinsic efficiency, scintillators are virtually the only materials currently used as detectors in PET imaging systems.
A match event is logged when there is an overlap of the individual logic outputs on the match modules. The time width of the overlap depends on the scintillation characteristics of the detectors. For current PET scanners, that width ranges from 6 to 12 ns. Although this is a very short time compared to most human activities, it is quite long compared to the distances traveled by photons traveling at the speed of light. Light travels at approximately 30 cm/ns, so a duration of 6 ns corresponds to a distance uncertainty of about 90 cm, which is the approximate diameter of the detector ring. As a result, the differential source distance between the detectors has no observable effect on the timing of coincidence events in conventional PET systems.
The arrival time of the annihilation photons is truly simultaneous only when the source is located precisely halfway between the two opposing coincidence detectors. If the source is offset from the midpoint, there will be a corresponding arrival time interval, since one annihilation photon will have a shorter distance to travel than the other. As discussed above, this time differential is too small to be useful in PET systems of conventional design. However, several of the scintillators used in PET scanners (eg, LSO, LYSO) are capable of faster response than the 6-12 ns timing discussed above. With proper electronics, the coincidence time window has been reduced to 600 ps for these detectors, resulting in a source location uncertainty of 9 cm. Even with that reduction, time-of-flight localization cannot be used to directly generate tomographic images, but can be used to regionally restrict the backprojection operation to areas where the sources are approximately located. In current implementations, the inclusion of time-of-flight information reduces noise in the reconstructed images by a factor of 2. Time-of-flight PET scanners became commercially available for a short time in the 1980s. These systems used detectors BaF2 which are very fast, but unfortunately have very low detection efficiency. As a result, these devices did not compete well with conventional BGO-based PET scanners. In 2006, a time-of-flight machine based on LYSO detectors was reintroduced and is now commercially available.
The only criteria for registering a match event is overlapping output pulses in the match module. True matches occur when a source is at the LOR defined by two detectors. It is possible that the events detected in the two source match detectors that are not on the response line happen by chance. As the count rate increases in each of the individual detectors, the probability of false matches from uncorrelated events increases. These events are called random Prayed accidental coincidences. The random match rate (R.) is directly proportional to the width of the coincidence time window
