| The Description of Photon Scattering on the Basis of X-ray CT data G | |||
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| F. Pönisch, W. Enghardt, J. Henniger1 | |||
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The process of photon scattering considerably influences the image quality in positron emission tomography (PET): for head and neck imaging only about 20 % of the annihilation photons escape the patient without any interaction with the tissue. About 25 % of the registered true coincidences are influenced by Compton or Rayleigh scattering, which may destroy the correspondence between the source and the reconstructed radioactivity distribution, especially in highly inhomogeneous regions of the human body. In the PET application for quality assurance of carbon ion therapy [1] the scatter problem has to be considered not only with respect to the reconstruction of the source distributions from the measured data but, furthermore, to the prediction of the b+-activity distribution from the treatment plan and the time course of the irradiation [2]. Since the decision on the correctness of an irradiation is based on the comparison of this prediction with the b+-activity distribution reconstructed from the data acquired during the patient treatment, both data sets have to be processed the same way. Therefore, the first step of the prediction is a Monte Carlo calculation [2] that describes the stopping of the therapeutic ion beam in tissue, the decay of the b+-emitters, the propagation of positrons and the annihilation photons and finally the g-ray detection. This code produces a list mode data set like a measurement and thus it can be reconstructed in the same way as measured data. However, in the original Monte Carlo code photon scattering was processed in a simplified way by assuming a homogeneous scatter volume (400×200×200 mm3, r = 1.18 g/cm3) centered in the field of view (FOV) of the positron camera. Thus, the simulated data had to be reconstructed without attenuation correction. Obviously this approach is quantitatively incorrect and neglects the large tissue inhomogeneities of the head and neck region, as the typical target for carbon ion therapy at GSI. Therefore, a more comprehensive scatter description has been developed. It requires an attenuation map obtained from the X-ray computed tomograms (CT) of the patient and of the equipment for patient positioning within and nearby the camera FOV [3]. The positron and annihilation photon propagation is modelled by means of a Monte Carlo code that takes also into account multiple scattering. To reduce computing time the following approaches that do not influence the accuracy have been introduced: 1) Positron annihilation in air is neglected. 2) The points of interaction are determined by means of the Delta Scattering Photon Transport algorithm of ref. [4], leading to a run time being independent on the voxel size of the CT. 3) The processing of annihilation photons is stopped either after hitting a detector, after escaping the CT image stack or after reducing the photon energy below 250 keV (the lower energy threshold of the positron camera). 4) Due to the coincidence condition it is not necessary to follow the second photon of the annihilation pair if the first photon is not detected. The computing time is reduced by a factor of 4 in comparison with the original method. This CT based photon scatter estimation allows the reconstruction algorithm to be applied to the measured and simulated PET data sets in a similar way. This is, furthermore, the condition for including a scatter correction algorithm in the reconstruction in order to evaluate both simulated and measured data quantitatively. 1 TU Dresden, Institut für Strahlenschutzphysik
References
[1] W. Enghardt et al., Annual Report 1999, FZR-271 (1999) 89
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