Abstract
Gamma-ray detectors based on thick monolithic scintillator crystals can achieve spatial resolutions <2 mm full-width-at-half-maximum (FWHM) and coincidence resolving times (CRTs) better than 200 ps FWHM. Moreover, they provide high sensitivity and depth-of-interaction (DOI) information. While these are excellent characteristics for clinical time-of-flight (TOF) positron emission tomography (PET), the application of monolithic scintillators has so far been hampered by the lengthy and complex procedures needed for position- and time-of-interaction estimation. Here, the algorithms previously developed in our group are revised to make the calibration and operation of a large number of monolithic scintillator detectors in a TOF-PET system practical. In particular, the k-nearest neighbor (k-NN) classification method for x,y-position estimation is accelerated with an algorithm that quickly preselects only the most useful reference events, reducing the computation time for position estimation by a factor of ~200 compared to the previously published k-NN 1D method. Also, the procedures for estimating the DOI and time of interaction are revised to enable full detector calibration by means of fan-beam or flood irradiations only. Moreover, a new technique is presented to allow the use of events in which some of the photosensor pixel values and/or timestamps are missing (e.g. due to dead time), so as to further increase system sensitivity. The accelerated methods were tested on a monolithic scintillator detector specifically developed for clinical PET applications, consisting of a 32 mm × 32 mm × 22 mm LYSO : Ce crystal coupled to a digital photon counter (DPC) array. This resulted in a spatial resolution of 1.7 mm FWHM, an average DOI resolution of 3.7 mm FWHM, and a CRT of 214 ps. Moreover, the possibility of using events missing the information of up to 16 out of 64 photosensor pixels is shown. This results in only a small deterioration of the detector performance.
Highlights
Monolithic scintillator detectors are based on a single-crystal scintillator with typical edge dimensions of 15–50 mm and a thickness of 10–25 mm, coupled to a pixelated photosensor
The accelerated methods were tested on a monolithic scintillator detector developed for clinical positron emission tomography (PET) applications, consisting of a 32 mm × 32 mm × 2 2 mm LYSO : Ce crystal coupled to a digital photon counter (DPC) array
The applicability and effectiveness of the new methods are demonstrated on a monolithic scintillator detector developed for clinical TOF-PET applications, consisting of a 32 mm × 32 mm × 22 mm LYSO : Ce crystal coupled to a digital photon counter (DPC) array
Summary
Monolithic scintillator detectors are based on a single-crystal scintillator with typical edge dimensions of 15–50 mm and a thickness of 10–25 mm, coupled to a pixelated photosensor (figure 1). The photosensor registers the light intensity on each of its pixels and acquires one or more timestamps, from which the position-, time- and energyof-interaction are derived. These detectors can achieve a spatial resolution better than 2 mm full-width-at-half- maximum (FWHM) (Li et al 2012, Cabello et al 2013, Seifert et al 2013, Borghi et al 2015) in combination with a coincidence resolving time (CRT) below 200 ps FWHM (van Dam et al 2013). Examples of statistical models are maximum-likelihood positioning (Ling et al 2007, Hunter et al 2009), neural networks (Bruyndonckx et al 2003, 2004), and k-nearest neighbor (k-NN) methods (Maas et al 2006, van Dam et al 2011a)
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