Abstract
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Objectives: Breast cancer is the most common cancer in women worldwide, with 1.3 million cases diagnosed per year. The current standard of care in breast cancer management has challenges. Mammograms are well known for having low specificity and often being inconclusive for patients with dense breasts, leading to unnecessary surgical procedures and patient trauma. The conventional PET systems have neither the photon sensitivity nor the spatial resolution required to affect earlier stages of breast cancer management. In recent years, there have been several groups and companies that developed high-performance dedicated positron emission mammography (PEM) systems that show promise for more sensitive cancer detection than standard clinical cameras while also providing better specificity than traditional anatomic imaging modalities such as x-ray mammography [1-5]. In this study, we used a prototype PET setup based on 3-D position-sensitive CZT detectors to demonstrate the combined PET-Compton data acquisition and its application in breast cancer imaging. The CZT detector used in the system not only provides an excellent intrinsic spatial resolution of around 0.5 mm in X-, Y-, and Z- directions, but also offers an excellent energy resolution of around 1% at 511 keV and the ability to detect and localize multiple interactions. This makes it capable of forming images with single photons through Compton scattering, and also use the Compton kinematics on coincidence photons to help reject scattered or chance coincidence in PET data. The objective of this study is to use PET and Compton data to jointly reconstruct the distribution of positron emitters in the breast and to enhance the image quality in regions close to the chest wall and even inside the chest wall through Compton imaging.
Methods: The proof-of-concept Compton-PET scanner is shown in Fig. 1(A). The system consists of four CZT detector panels, and each panel consists of four (2×2 ) CZT detectors of 2.2×2.2×1.0 cm3 in size. The detector offers a subpixel resolution of <0.5 mm FWHM in all three dimensions with CZT crystal of up to 1.5 cm thickness, and an energy resolution of 5.4 keV FWHM at 511 keV [6]. In the prototype PET system, the distance between opposite PET panels is designed to be 10 cm. The output of all CZT detectors are saved in list-mode with the position, energy and timing information of each individual gamma ray interaction. The coincidence pairs and or the single-photon Compton events are determined in post-processing through the time stamp of individual detected events. We computed a list-mode system response function (SRF) that represents the probability of each source voxel contributing to the given event, and the SRF is used by a list-mode penalized maximum likelihood algorithm for reconstructing the joint-PET-Compton images.
Results: Fig. 1(B) shows the energy spectrum measured with the CZT detector in this Compton-PET prototype. Fig. 1(C) shows the PET image of a resolution phantom acquired with the Compton-PET prototype, where the hot-rods of 0.75 mm diameter can be resolved. Fig. 1(D&E) illustrates a simple back-projection of seven single photons detected on the CZT detector through Compton interactions from a Na-22 point source placed at 5 cm from the CZT detector.
Conclusions: The joint-PET-Compton imaging reported in this paper not only offers very high sensitivity and an ultrahigh-resolution in conventional PET acquisition, but it also allows simultaneous detection of single (or coincidence) photons through Compton interactions. This offers the unique ability to enhance the PET images in regions close to the chest wall and provide imaging information, using Compton data alone, for areas inside the chest wall. The joint-PET-Compton image acquisition would, therefore, help to address one of the major limitations in current positron emission mammography (PEM) instrumentations.
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