The LSO/APD array as a possible detector for in-beam PET in hadron therapy

Abstract

We have studied the performance of finger-like LSO:Ce (LSO) crystals coupled one by one to pixels of avalanche photodiode detector (APD) arrays during their operation in coincidence at 12 C ion beams of parameters being typical for tumor irradiations. In a first step of these experiments the parameters of the detectors and the signal processing setup have been characterized off-beam, i.e., by means of + radioactive sources ( 22 Na, 68 Ge). Afterwards, the apparatus was installed at the medical beam line of the heavy ion synchrotron (SIS) of the Gesellschaft für Schwerionenforschung (GSI) at Darmstadt, Germany. Here the + activity produced by nuclear fragmentation reactions of 12 C 200.3 AMeV heavy ion beams with polymerized methyl methacrylate phantoms were measured. Furthermore, a 68 Ge source was included into the in-beam experiment, in order to check the stability of the setup and to compare energy and time resolution before, during, and after the phantom irradiations.Additionally, it could be demonstrated by means of high resolution -ray spectroscopy that LSO is not activated by the light projectile fragments escaping the patients downbeam during therapeutic irradiations The experimental results indicate that LSO scintillator is a suitable material for in-beam positron emission tomography (PET) and, furthermore, the LSO/APD array is a feasible detector concept for in-beam monitoring of the dose application by means of PET.

First in-beam PET imaging with LSO/APD array detectors

“…This was a major achievement, since it demonstrated: 1) the in-beam imaging capability of this detector type and 2) the sufficient radiation hardness of such detectors to be operated at a radiotherapy treatment site. For completeness, we mention that an activation of LSO in therapeutical irradiations can be ruled out according to the experiments outlined in [9]. In summary, the obtained results indicate that LSO is a scintillator material suitable for in-beam PET.…”

“…Each detector consists of 32 LSO crystals (2.15 2.15 15 mm each) coupled individually to the 4 8 pixels of an S8550 Hamamatsu APDA with silicon glue [9]. For every detected coincidence, all 64 energy channels were read, together with the coincidence time between the two detectors and, in addition, one channel containing a digital signal provided by the synchrotron (high for spill on, i.e., beam extraction, s, low for spill off, i.e., pause between extractions, s).…”

“…For every detected coincidence, all 64 energy channels were read, together with the coincidence time between the two detectors and, in addition, one channel containing a digital signal provided by the synchrotron (high for spill on, i.e., beam extraction, s, low for spill off, i.e., pause between extractions, s). This synchrotron signal allows, as necessary during in-beam imaging, to select only the data acquired during the pauses between the extractions, hence reducing the background caused by random coincidences of prompt -rays from nuclear reactions as the beam penetrates the target [9], [10]. The reconstructed images presented here took into account only data acquired between beam extractions.…”

Advantages of Digitally Sampling Scintillation Pulses in Pileup Processing in PET

“…12. The coincidence timing histogram obtained with software CFD (4) and with the DiSER method (3). The FWHM of timing histogram obtained by using software CFD was 601 ps, and the FWHM obtained by using the DiSER method was 692 ps.…”

Direct time-of-flight for quantitative, real-time in-beam PET: a concept and feasibility study

“…These requirements are a high detection efficiency for annihilation radiation, a fast scintillator enabling a narrow coincidence time window for the suppression of random coincidences (Enghardt et al 2005, good energy resolution for rejecting Compton-scattered photons, insensitivity against influences from the therapy beam, hardness against hadronic radiation, magnetic field resistance in the case of in-beam PET at rotating beam deliveries, and small volume due to integration reasons and in order to provide flexibility for patient positioning. The combination of crystals of LSO (cerium-doped lutetium oxyorthosilicate, Lu 2 SiO 5 :Ce 3+ ), a fast scintillator with a fluorescence decay constant of 41 ns (Melcher and Schweitzer 1992), coupled to avalanche photodiode arrays (APDA) solves, for imaging small tumours in the head-and-neck region, all the requisites mentioned (Kapusta et al 2004, Crespo et al 2004, Crespo 2005 provided that the implementation of a closed-ring tomograph, or a dual-head tomograph with small gaps, is feasible (Crespo et al 2006a). In addition, the natural background activity of LSO, amounting to approximately 240 Bq cm −3 due to the presence of the isotope 176 Lu (Huber et al 2002), in combination with the very low β +activity densities typically measured during in-beam PET (Enghardt et al 2004a), requires a coincidence time resolution equal or better than 1 ns full width at half maximum (FWHM) at the detector level (Crespo et al 2006b).…”