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Positron Emission Tomography (PET) – Nuclear Medicine Imaging Technology
Positron emission tomography (PET) is an alternative to nuclear medicine imaging that has several advantages over SPECT. PET uses positron-emitting radionuclides that result in the emission of collinear pairs of 511-keV nuclear photons. Collision detection of nuclear photons eliminates the need for collimation and makes PET more efficient than SPECT for detecting radioactive decay. Even more important, there are positron-emitting radionuclides for oxygen, carbon, nitrogen, and fluorine, which allow many 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) supplied locally, and there is no saturated area in the United States where it is not available. Many others such as 82Rb and 68Ga come from radionuclide producers who supply radionuclides on demand despite their short half-lives.
Plain detection provides spatial resolution without the need for lead accumulation by taking advantage of the fact that nuclear photons resulting from positron emission are approximately colinear. Events are only counted if they are seen simultaneously by two opposing detectors. The level of suspicion defined by ordinary investigators is called a line of results (LOR). Two single elephant systems are used with an additional ordinary module. Each system will generate a logic pulse when they detect an event that falls in the selected power window. If two logic pulses overlap in time in the coincidence module, a coincidence event is recorded. PET systems use a large number (>10,000) of detectors arranged as multiple rings to form a cylinder. Since any detector can be co-located with other detectors in the cylinder, the resulting LORs provide sufficient sampling to obtain the spectral information required for tomography.
The intrinsic detection efficiency for a single detector depends on the atomic number, density, and thickness of the detector. Ideally, the internal detection efficiency should be 1, but at 511 keV that is difficult to achieve, although the internal efficiency for some detectors is greater than 0.8. Collision detection requires that both detectors register an event. Since the interactions in the two detectors are independent, the expected efficiency depends on the internal efficiency product of each detector. As a result, the random detection efficiency is often less than that for a single detector, and that difference increases for low-efficiency detectors. Because of the need for higher efficiency, scintillators are the only devices currently used as detectors in PET imaging systems.
A common event is recorded when there is an overlap of the logical results of the units in common modules. The duration of the overlap depends on the scintillation characteristics of the detectors. For current PET scanners, the range is from 6 to 12 ns. Although that’s a very short time compared to most human actions, it’s fairly long compared to the distances photons travel at the speed of light. Light travels at about 30 cm/ns so that a time duration of 6 ns corresponds to a distance uncertainty of about 90 cm, which is the approximate detector diameter. As a result, the spatial difference of the source between the detectors has no noticeable effect on the timing of spontaneous events in conventional PET systems.
The arrival time of the nuclear photons is truly simultaneous only when the source is exactly halfway between two opposing collision detectors. If the source is displaced from the center point, there will be a corresponding arrival time interval since one nuclear photon will have a shorter distance to travel than the other. As discussed above, this time difference is too small to be useful in custom designed PET systems. However, many of the scintillators used in PET tomographs (for example, LSO, LYSO) have the ability to have a faster response than the 6 to 12 ns time discussed above. With the appropriate electronics, the ordinary time window is reduced to 600 ps for these detectors, yielding a source localization uncertainty of 9 cm. Even with that limitation, time-of-flight classification cannot be used to generate tomographic images directly, but it can be used to restrict the activity of the area to areas where resources are close. In current implementations, the inclusion of time-of-flight information reduces the noise in the reconstructed images by a factor of 2. Time-of-flight PET tomographs were actually commercially available for some time in the 1980s. these systems use BaF2 detectors which are very fast, but unfortunately have very low detection efficiency. As a result, these devices do not compete well with conventional BGO-based PET printers. In 2006, a time-of-flight device based on LYSO detectors was rebuilt and is now commercially available.
The only source for recording an ordinary event is the overlap of the outputs of the ordinary module. True collisions occur when a source is on the LOR defined by two detectors. It is possible that events detected in two collision detectors from non-linear sources may occur by chance. As the reading rate in each of the single detectors increases, the likelihood of false positives resulting from unrelated events increases. These events are called randomly or accidental coincidences. Accident rate (R) is directly proportional to the size of the random time window
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