TOF-PET
Time-of-Flight Positron Emission Tomography
Radiomolecular imaging is a branch of nuclear medicine aimed at visualizing physiological processes in-vivo using radionuclides. A radioactive isotope is used to label part of the molecules of a radiopharmaceutical, or tracer, which is designed to target a particular feature of interest within the subject. By imaging the distribution of this tracer it is possible to obtain information about molecular processes non-invasively. The tracer distribution is measured by detecting the gamma rays emitted by the radionuclide. Positron emission tomography (PET) is a molecular imaging technique that makes use of tracers labeled with positron-emitting isotopes. PET is usually integrated with CT in a so-called PET/CT scanner to simultaneously obtain functional and anatomical information. Currently, the most common application of PET is to diagnose tumors and to find cancer metastasis. However, PET is also employed for treatment response monitoring, for diagnosing brain diseases, to assess cardiac viability, and as a research tool to e.g. study brain or heart function or to support drug development.
The positrons emitted by the administered radiotracer annihilate with electrons in the body almost instantaneously. The PET imaging technique is based on the coincident detection of the two 511 keV gamma rays that are emitted in opposite directions as a result of this process. The two gamma-ray interaction points define a so-called line-of-response (LOR) on which the annihilation must have taken place (Figure 1). The combined information of many millions of LORs measured during a PET acquisition is used to produce a 3D image of the estimated tracer distribution using analytical or probabilistic image reconstruction methods. If the time difference between the two moments of interaction can be measured with sufficient precision (≤ 500 ps) this so called time-of-flight (TOF) information can be used to estimate the segment of the LOR on which the annihilation occurred. This helps to improve the signal-to-noise ratio of the image. Thus, PET image quality is determined largely by the performance of the detectors used to measure the position and time of interaction of the gamma rays. Current PET detectors are based on scintillation crystals that convert the energy of the gamma photons into tiny flashes of light and photomultiplier tubes (PMTs) that convert these optical signals into electronic pulses.
Figure 1 Imaging principle of PET: (a) After annihilation of a positron and an electron, two 511 keV photons are emitted in (almost) opposite directions; (b) When two interactions are simultaneously detected within a ring of detectors surrounding the patient, it is assumed that an annihilation occurred on the so-called line-of-response (LOR) connecting the two interactions. By recording many LORs the activity distribution can be tomographically reconstructed.
Within Radiation and Isotopes for Health (RIH) we study new gamma-ray detectors for clinical PET/CT and PET/MRI scanners. In particular we investigate detectors based on monolithic scintillator crystals and a novel type of solid state photosensor called silicon photomultiplier (SiPM) (Figure 2). This approach may overcome inherent limitations present in the currently used detectors based on pixelated crystals. To this end, we focus on testing new crystal materials and photosensor technologies and on developing new algorithms to optimally estimate the time and position of interaction of gamma-rays inside the crystal [1-3]. In the laboratory we have already been able to simultaneously achieve spatial resolutions close to ~1 mm FWHM and coincidence resolving times (CRTs) significantly better than 200 ps FWHM with monolithic crystals of sufficient thickness (~20 mm) for the efficient detection of 511 keV gamma rays. In addition, we can also obtain depth-of-interaction (DOI) information and good energy resolution. These results are a significant improvement compared to the spatial resolutions of ~4 mm (without any DOI information) and the CRTs of ~ 500-600 ps FWHM which are currently achieved with pixelated crystal detectors in commercial clinical PET scanners.
Besides further improving the performance of the monolithic scintillator detector, we aim to experimentally emulate realistic PET scanner acquisitions in the near future to demonstrate that this new detector can be a competitive technology for future PET systems.
Figure 2 Photographs of a gamma-ray detector based on a LYSO monolithic scintillator crystal (32 mm x 32 mm x 22 mm) and a fully digital implementation of the silicon photomultiplier, the digital SiPM or dSiPM. The side surfaces have been covered with reflective foil, while the top surface is still uncovered.
References
[1] S. Seifert, G. v. d. Lei, H. T. v. Dam, and D. R. Schaart, "First characterization of a digital SiPM based time-of-flight PET detector with 1 mm spatial resolution," Physics in Medicine and Biology, vol. 58, p. 3061, 2013.
[2] H. T. van Dam, G. Borghi, S. Seifert, and D. R. Schaart, "Sub-200 ps CRT in monolithic scintillator PET detectors using digital SiPM arrays and maximum likelihood interaction time estimation," Physics in Medicine and Biology, vol. 58, p. 3243, 2013.
[3] H. T. van Dam, S. Seifert, R. Vinke, P. Dendooven, H. Löhner, F. J. Beekman, et al., "Improved Nearest Neighbor Methods for Gamma Photon Interaction Position Determination in Monolithic Scintillator PET Detectors," Nuclear Science, IEEE Transactions on, vol. 58, pp. 2139-2147, 2011.