## 1 Introduction  
Proton therapy has a unique feature that makes it more advantageous compared to conventional radiotherapy. This unique feature stems from its dose distribution which remains quasi-constant at a low level along the path traveled by proton and rises sharply to a maximum value at [PERSON]'s peak towards the end. Beyond [PERSON]'s peak, the dose absorbed is almost negligible, reducing the risk of damage to the healthy tissues [1, 2, 3]. To fully leverage this Bragg's peak, accurate in-vivo range verification is essential. Two main techniques are employed for this in vivo range verification which are based on the by-products of the interaction between patient and the incident irradiation. They are PET (positron emission tomography) which uses the delayed gamma [4] and prompt gamma imaging which uses prompt gamma [5]. Prompt gamma signals can easily penetrate through the tissues and emerges without much interaction with the tissues [4], hence the information of the location of its origin is not distorted, renders accurate identification of the Bragg's peak position [6, 7, 8], and thus it is commonly used. Silicon photomultipliers (SiPM) and position-sensitive multi-output detectors are commonly employed for detecting the prompt gamma signals. The detector commonly consists of a multi anode position-sensitive photomultiplier tube (PSPMT) which has \(8\times 8\) multi-anode spaced with a pixel size of 6 mm \(\times\) 6 mm per anode [7]. A scintillator crystal is used along with the detector to generate fluorescence when a prompt gamma signal is incident on it. This light is then converted into photoelectrons on the photocathode of the PSPMT and a current signal is generated. The PSPMT current signals are conventionally processed by a preamplifier and a shaping circuit, and the shaped signals are converted into digital signals using analog to digital converters. A special algorithm is used to extract the energy and position of the radiation from the digital signals [9; 10]. Conventionally, these circuits that process the PSPMT signals are built using an operational amplifier (OPA), and they must be placed in very close proximity to the detector to reduce signal loss and distortion, and hence their operation environment is radiative. In proton therapy, secondary neutrons are also generated along with the prompt gamma signals. The main source of secondary neutrons is the treatment unit and the patients themselves. The neutrons generated from the treatment unit are known as the external neutrons and are generated due to the interaction of proton with beam delivery system, whereas neutrons generated from the patient is a result of proton interaction within the patient body which are known as internal neutrons [11; 12]. These neutrons will negatively affect the health of the patient as well as the electronics present in the treatment room such as electronics/data acquisition systems, monitoring devices, robotic patient positioners and onboard imaging, etc. [13; 14]. The radiation hardness of these electronic devices is of great concern as they are operating in the radiation environment and their durability will be compromised, as can also be observed in our treatment room. The secondary and scattered radiation field present in and around the passive proton treatment nozzle was evaluated and their results showed that the dose at any point is highly dependent on the position relative to the isocenter [12]. It was found that the dose on the gantry is approximately 3 to 6 times higher than that at the gantry floor. Since the readout circuits are in close proximity to the isocenter, their reliability and performance are greatly affected by the neutrons. Several studies [15; 16; 17; 18; 19; 20; 21] have been conducted recently on the commercial OPA Integrated Circuits (ICs) under neutron radiation and they found that the characteristics of OPAs such as gain, slew rate, offset current, etc. are degraded after the radiation.