First Study of a HEXITEC Detector for Secondary Particle Characterisation during Proton Beam Therapy  
[PERSON]  
1 Department of Medical Physics and Biomedical Engineering, University College London, London WC1E 68 BT, UK; [EMAIL_ADDRESS] (M.I.-P.-L.); [EMAIL_ADDRESS] (J.C.K. )1 Department of Medical Physics and Biomedical Engineering, University College London, London WC1E 68 BT, UK; [EMAIL_ADDRESS] (M.I.-P.-L.); [EMAIL_ADDRESS] (J.C.K. )1 Department of Medical Physics and Biomedical Engineering, University College London, London WC1E 68 BT, UK; maria.lara.lara.l9@ucl  
fractions (up to 35), with each fraction having a typical dose delivery of 2 Gy. Typically, the margins of error for proton beam therapy treatments are around 2.5-3% of the proton range, and range from 2 to 3 mm [4]. The uncertainty arises from a wide range of sources, such as variations in the patient's anatomy between fractions (which result in an intrinsic error in the stopping power calculations), the errors during the conversion process from CT images into stopping power maps, and even daily variations regarding the setup and beam settings [5]. Due to this, the development of new methods to perform online proton range monitoring is currently a widely studied area, as they may provide feedback that allows for the treatment to be adapted, with the ultimate goal of reducing the error margins. Positron emission tomography (PET) systems are considered to be the only clinically available method for 3D range monitoring that is non-invasive, as they measure the positron activity distribution induced by the protons interacting with the irradiated tissue. This method has been studied for particle irradiation offline/after treatment [6], but has longer acquisition times and the risk of not having updated treatment information until treatment has already been delivered. Moreover, there are intrinsic limitations to proton range verification with offline PET/CT systems that introduce errors in the range estimation. The PET signal, which comes from measurable \(\beta^{+}\) activity, is not identical to the delivered dose. Thus, verification must be obtained by finding a valid comparison between the PET signal and a reference value, which is often calculated with Monte Carlo simulations by simulating activity distributions for different tissues based on the CT information; however, the uncertainties that arise from this comparison have a non-negligible effect on the range estimation, as well as aspects like biological washout (where the radiotracer concentration decreases over time within a specific tissue of interest) and patient motion, suggesting that offline PET range verification is only feasible for a limited number of positions and tumour locations [7]. Due to this, in vivo verification studies have also been performed [8; 9; 10]. However, the activation process is not immediate: the signal delay varies from seconds to minutes due to the half-lives of the target fragments, which poses several challenges: additional geometrical constraints in the setup to avoid interference with the beam and a consequently more demanding data acquisition system to reduce the large radiation background [11]. Thus, PET systems for this application require optimisation in terms of beam-scanner arrangement, high resolution and efficiency and its combined use with other modalities [12], whilst efforts have also been made to make the acquisition quasi-prompt [13]. The constraints in the PET system as a means of verifying the proton range and monitoring the dose have led to a new focus for state-of-the-art research, based on the exploitation of information coming from secondary radiation that is produced in nuclear interactions between protons and the atomic nuclei inside the body. These particles are emitted along the beam track and they are correlated with the proton dose deposition, which suggests that secondary particle detection can be used to locate the proton range [14].