Advances in Radiotherapy Dosimetry Techniques and Pre-Treatment Verification

Abstract

In the fight against cancer, radiation therapy plays a vital role, with its two essential approaches: internal, involving the insertion or implantation of radioactive material into the patient’s body, or uptake of radiopharmaceutical, and external. Precise delivery of the appropriate radiation dosage to the tumor is critical for achieving favorable outcomes. This is where dosimetry becomes crucial—a scientific discipline that involves measuring, calculating, and evaluating radiation doses. Medical physicists utilize dosimetry to ensure the accuracy and proper calibration of machines that administer ionizing radiation, ensuring safety. This chapter provides a brief overview of advanced techniques and equipment used in dosimetry, with a primary focus on photon and electron dosimetry, the most widely employed forms of radiation for radiotherapy worldwide.

1. Introduction

Radiotherapy aims to deliver a specific amount of radiation to the tumor while minimizing damage to nearby organs at risk (OAR). Over the past decade, there has been a surge in the development of advanced radiotherapy techniques in both computation of radiation dose (inverse planning) and delivery techniques such as Intensity Modulated Radiation Therapy (IMRT) and volumetric modulated arc therapy (VMAT). The main objective of these sophisticated developments is mainly to enhance target dose uniformity and minimize high-dose regions and doses to nearby tissues and organs-at-risk (OARs). These advanced techniques are more complicated than 3D conformal therapy and require sophisticated tools and techniques for dosimetry to guarantee that the delivered dose corresponds accurately with the planned doses. Numerous guidelines and dosimeter devices, including detectors, phantoms, and electronic devices, have been designed for this purpose. This chapter will discuss the various equipment and methods used for dosimetry in photon and electron therapies, brachytherapy, neutron, and proton therapies.

2. Dosimetry of photon and electron beams

2.1 Summary of AAPM task group 51

In radiotherapy, the American Association of Physicists in Medicine Task Group 51(AAPM TG-51) [1] and its Addendum [2] are one of the main dosimetry references for external high-energy photon and electron beams. These guidelines are based on measurements conducted with ionization chambers (IC), the most common radiotherapy radiation detectors. These detectors are made up of a gas-filled chamber that has two electrodes of opposing polarity, namely, the anode and cathode. The electrodes can be in the form of either parallel plates (known as plane-parallel IC) or a cylinder (thimble) with an internal anode wire located coaxially. Ionization chambers are a type of detector that collects ion pairs from gases to measure incident radiation [3, 4]. By creating an electric field in the gas-filled cavity through a voltage potential applied between electrodes, charged particles move to electrodes of opposite polarity. This generates an ionization current, which is measured by an electrometer circuit. The current ranges from femtoamperes to picoamperes, proportional to the radiation dose, and depends on the chamber design [3]. Cylindrical chambers are frequently used in Ref. dosimetry applications of MV photon beams and electron beams (>10 MeV), while plane-parallel is designed for low energy electrons (<10 MeV) as well as in applications where precise measurement location is important such as measurement of electron percent depth dose distributions [1].

An ion chamber is calibrated in terms of absorbed dose-to-water in a standard laboratory’s reference quality Co⁶⁰ gamma-ray beam. The calibration is valid under reference conditions, which specify a field size of 10 × 10 cm² and a reference depth of 10 cm under the water surface for photon beams [5]. For electron beams, the reference depth is determined by Eq. (1)

$${\mathrm{D}}_{\mathrm{ref}}=0.{\text{6R}}_{50}-0.1\mathrm{cm}$$E1

where R₅₀ is the depth at which ionization drops to 50% of its maximum value. The reference field size is 10 × 10 cm² and 20 × 20 cm² where R₅₀ ≤ 8.5 cm and R₅₀ > 8.5 cm, respectively.

The absorbed dose in water is calculated by a general formalism for both photon and electron beams Eq. (2).

$$D_w^Q=M{N}_{D,w}^Q\left(\mathit{Gy}\right)$$E2

where M is the fully corrected reading of the electrometer, which is expressed in units of C or rdg, and can be calculated using Eq. (3)

$$M=P_{\mathit{ion}}{P}_{\mathit{TP}}{P}_{\mathit{pol}}{P}_{\mathit{\text{elec}}}{M}_{\mathit{raw}}\left(C\mathit{\text{or}}\mathit{rdg}\right)$$E3

P_ion corrects the incomplete ion collection efficiency of the detector, and its value should be ≤1.05. There are two formulas to consider when dealing with continuous beams, such as Co-60 (Eq. 4), and pulsed or pulsed-swept beams (Eq. 5):

$$P_{\mathit{ion}}=\frac{1.00-{\left(\frac{V_H}{V_L}\right)}^2}{\frac{M_{\mathit{raw}}^H}{M_{\mathit{raw}}^L}\times {\left(\frac{V_H}{V_L}\right)}^2}$$E4

and

$$P_{\mathit{ion}}=\frac{1.00-\frac{V_H}{V_L}}{\frac{M_{\mathit{raw}}^H}{M_{\mathit{raw}}^L}\times \frac{V_H}{V_L}}$$E5

V_H is the standard operating voltage for the ion chamber, and $M_{\mathit{raw}}^H$ is the raw reading at this voltage.

V_L is reduced voltage (at least half of V_H), and $M_{\mathit{raw}}^L$ is the raw reading of the detector at the corresponding voltage.

P_TP is the temperature–pressure correction factor since calibration factors are given for standard environmental conditions of T₀ = 22°C and P₀ = 101.33 kPa (1 atmosphere).

$$P_{\mathit{TP}}=\frac{273.2+T}{273.2+22}\times \frac{101.3}{P}$$E6

P_pol corrects the polarity effects and depends on beam quality and other conditions such as cable position. To ensure an accurate calibration procedure, it’s essential to apply this correction every time.

$$P_{\mathit{pol}}=\left|\frac{M_{\mathit{raw}}^{+}-M_{\mathit{raw}}^{-}}{2M_{\mathit{raw}}}\right|$$E7

The raw readings for positive and negative normal operating voltage of the detector are represented by $M_{\mathit{raw}}^{+}$ and $M_{\mathit{raw}}^{-}$, respectively, and M_raw can be either $M_{\mathit{raw}}^{+}$ or $M_{\mathit{raw}}^{-}$.

It is recommended that the polarity correction be kept below 0.3%, although the Addendum to TG-51 permits up to 0.4%.

P_elec is the electrometer correction factor when performing separate calibrations for the electrometer and ion chamber. It is advisable to calibrate both the ion chamber and the electrometer together, in this case, P_elec = 1.00. Moreover, this factor has a unity value for cross-calibrated parallel plate chambers as it cancels out of the final equations (section X C of TG-51).

$N_{D,w}^Q$ is the absorbed dose to water calibration factor for the user’s ion chamber located in an arbitrary beam of quality Q, under reference conditions, and its value can be expressed in units of Gy/C or Gy/rdg. Typically, absorbed-dose calibration factors are obtained under reference conditions using a Co-60 beam. In this case, $N_{D,w}^Q$ is calculated for the beam of quality Q, which can be photon or electron using Eq. (8).

$$N_{D,w}^Q=k_Q{N}_{D,w}^{60\mathit{Co}}$$E8

k_Q is the quality conversion factor that converts the absorbed dose to water calibration factor for a ⁶⁰Co beam into the calibration factor for a beam of quality Q. This factor is chamber specific for photon beams and its value for most ion chambers used as reference chamber can be found in section IX.B of TG-51.

For electron beams, the quality conversion factor has two components (Eq. 9):

$$k_Q=P_{\mathit{gr}}^Q{k}_{R50}$$E9

where $P_{\mathit{gr}}^Q$ is one for plane-parallel chambers and only applies to the cylindrical ion chamber and corrects for the ionization gradient at the reference depth. To determine this factor, the user must follow the measuring procedure outlined in TG-51. $k_{R50}$ that has two components as follows (Eq. 10).

$$k_{R50}=k_{R50}^{\prime }{k}_{\mathit{\text{ecal}}}$$E10

$k_{\mathit{\text{ecal}}}$ is the photon-electron conversion factor specified for each individual ion chamber in TG-51.

$k_{R50}^{\prime }$ is the electron beam quality conversion factor and depends on the beam quality.

In an electron beam, the dose of water is given by:

$$D_w^Q=M{P}_{\mathit{gr}}^Q{k}_{R50}^{\prime }{k}_{\mathit{\text{ecal}}}{N}_{D,w}^{60\mathit{Co}}\left(\mathit{Gy}\right)$$E11

To prevent any complications with additional waterproof sleeves and potential air gaps, it is best to use a waterproof ion chamber. An additional waterproofing sleeve is necessary if the chamber is not waterproof. A waterproofing sleeve should reduce the air gaps near the chamber wall (≤ 0.2 mm) and should be made of polymethylmethacrylate (PMMA) ≤ 1 mm thick.

Clinical reference measurements must be performed in a water phantom with a minimum size of 30 × 30 × 30 cm³, and the use of non-water phantoms is not allowed. If the beam penetrates the plastic wall of the water phantom and the wall’s thickness exceeds 0.2 cm, it is necessary to adjust all depths to their water equivalent values. This is achieved by measuring from the outer surface of the wall when the phantom contains water while also considering the density of the wall.

The beam quality must be specified to determine the correct value of the quality conversion factor for electron and photon beams. Beam quality is characterized by a parameter related to the central-axis depth-dose curves for the beam.

For photon beams, it is the percentage depth dose at 10 cm depth in the water of 10x10 cm² open field size at SSD = 100 cm. If the energy of the photon beam is less than 10 MV (%dd(10) <75%), the value of %dd(10) is %dd(10)_X. However, for beam energies >10 MV and flattening filter-free (FFF) beams, due to electron contamination, the value of %dd(10)_X is obtained from %dd(10)_pb parameter when the 1 mm thick lead foil is placed at the 50 ± 5 cm from a phantom surface:

$$\%\mathit{dd}{\left(10\right)}_x=\left[0.8905+0.00150\%\mathit{dd}{\left(10\right)}_{\mathit{pb}}\right]\%\mathit{dd}{\left(10\right)}_{\mathit{pb}}$$E12

And if the lead foil is located at 30 ± 1 cm above the phantom surface, %dd(10)_x is calculated by Eq. (13):

$$\%\mathit{dd}{\left(10\right)}_x=\left[0.8116+0.00264\%\mathit{dd}{\left(10\right)}_{\mathit{pb}}\right]\%\mathit{dd}{\left(10\right)}_{\mathit{pb}}$$E13

Moreover, if the %dd (10)_pb is less than 73 and 71% where the foil is at 50 and 30 cm, respectively, then %dd(10)_x = %dd(10)_pb.

It is important to use an SSD of 100 cm when determining the beam quality for photon and electron beams, as both %dd(10) and R₅₀ are SSD dependent.

In the process of measurement, cylindrical and spherical chambers should be shifted to the effective point of measurement. The effective point is located upstream of the actual point of measurement. This is because the secondary electrons tend to move forward, causing a shift in the depth-dose curve toward shallower depths. Therefore, the upstream shift for cylindrical and spherical chambers is 0.6r_cav and 0.5r_cav for photon and electron beams, respectively, where r_cav is the radius of the ion chamber’s cavity. No shift is required for parallel plate chambers, as the point of measurement is the center of the front (upstream) face of the chamber air cavity.

In summary, TG-51 applies to photon beams with an energy range of Co-60 to 50 MV and electron beams with nominal energies between 4 and 50 MeV and provides the following information: (1) Prescribing a methodology for clinical reference dosimetry, (2) determining the beam quality conversion factor for both photon and electron beams, (3) measuring gradient correction factor for cylindrical chambers in electron beams, (4) measuring various correction factors to correct the raw charge reading of the ion chamber, and (5) using cylindrical and plane-parallel ICs for absolute calibration measurements.

2.2 Detectors for patient-specific quality assurance

Intensity-modulated radiotherapy (IMRT) and VMAT treatment techniques involve many challenges, particularly the dosimetry of small fields, varying dose rates, low MU segments, multileaf collimator (MLC) motions, and gantry speed variations for VMAT [6, 7]. Furthermore, dosimetry conditions for IMRT/VMAT are significantly different from the open field conditions for which the chambers are calibrated. It is highly recommended to perform patient-specific quality assurance (QA) to identify any disagreements between the dose calculated by the treatment planning system (TPS) and the actual dose delivered by the treatment machine [8]. The patient-specific QA includes several dosimetric tasks that are performed before the treatment. Several phantoms and detectors such as cylindrical ion chambers, detector arrays, Gafchromic films, and so on have been designed and implemented for this purpose. To ensure accurate pre-treatment verifications, it is important to have a comprehensive understanding of detector/phantom characteristics, including energy and dose rate dependence, collecting volume size, charge leakage, design, and materials [9, 10].

2.2.1 Ionization chambers

In pre-treatment verification or patient-specific QA of IMRT treatment plans, a phantom (either 2D or 3D) is typically used to check the dose distribution, while an ion chamber is utilized for absolute dose verifications. As the plan becomes more complex, smaller field sizes and higher MU are required to achieve a conformal dose distribution. IMRT/VMAT plans include small control points or segments with a size of less than or equal to 1x1 cm². Therefore, for absolute dose verifications, it is recommended to use ion chambers with a small volume of 0.1 cc or less [11]. Nonetheless, there are potential downsides to using small ion chambers as their sensitivity may decrease with decreasing volume [6, 9, 12]. Leybovich et al. [9] evaluated the performance of ion chambers of varying volumes such as Farmer type NE 2581 0.6 cm³ (NE Technology, Essex, UK), Capintec PR-06 0.6 cm³ (Capintec, Inc., Ramsey, NJ), PTW 0.125 cm³ (PTW-New York, Hicksville, NY), and Exradin A1 0.009 cm³ (Standard Imaging Inc., Middleton, WI) for the absolute dose verification of tomographic and step-and-shoot IMRT treatment plans. Their study revealed that the response of 0.6 cc chambers was higher than the calculated dose when the chamber was partially irradiated. Additionally, the measurements from the smaller chambers (0.125 and 0.009 cm³) required correction for charge leakage due to chamber sensitivity being proportional to volume [9]. Another study compared three different cylindrical chambers including a 0.015 cm³ PTW-Freiburg 31,006 pinpoint (PTW-Freiburg, Freiburg, Germany), a 0.13 cm³ Wellhöfer IC10 (Scanditronix Wellhöfer North America, Bartlett, TN), and a 0.69 cm³ Farmer-type (FT) IC model NE2571 (Nuclear Enterprises, Fairfield, NJ) [6]. Their study indicated small ion chambers overresponded when irradiated with a small number of monitor units (MU) because the smallest-volume chambers are more sensitive to leakage compared to large-volume ion chambers. Evaluation of 50 IMRT treatment plan quality assurance procedures indicated the Farmer chamber measurements were found to be the closest to the TPS calculated values. However, all chambers measured higher doses than those predicted by TPS [6]. Research conducted by Kumar et al. [12] assessed the response of five different IC-phantom combinations in RapidArc therapy. This study included a Medtec IMRT phantom with Exradin (A16) micro-ion chamber (0.007 cm³) and PTW pinpoint chamber (TM31014–0193) (0.015 cm³), PTW-Octavius phantom with semiflex chamber (TM31010–1571) (0.125 cm³) and PTW 2D-array 729 (T10024) (0.125 cm³), and an indigenously made circular wax phantom with a 0.61 cm³ chamber (NE 2571). All absolute dose measurements taken at the isocenter were compared to those calculated by the Eclipse TPS (Varian Medical Systems, Palo Alto, CA) version 8.6. The results showed that the leakage has a greater effect on small-volume chambers as the sensitivity of the chamber is directly related to its volume. Additionally, positioning is a very important factor when using smaller-volume chambers because they are more susceptible to errors caused by geometrical variations within the treatment fields (Figure 1) [12].

Figure 1.

Farmer chamber.

2.2.2 Gafchromic films

As aforementioned, verifying the accuracy of IMRT/VMAT delivery can be difficult due to numerous small fields, irregularities, steep dose gradients, and sharp penumbra. Consequently, the conventional method of measuring point doses using an IC for patient-specific QA is insufficient for highly modulated treatment fields like stereotactic radiosurgery (SRS) or Stereotactic Body Radiation Therapy (SBRT) where small targets are irradiated in a single fraction using a high-intensity photon beam [13]. Radiochromic films, typically referred to as “GAFChromic,” with the added advantage of self-development have been employed for this purpose. Due to their high spatial resolution, they can provide a more accurate map of dose distribution in the sharp dose gradient regions [14]. GAFChromic™ EBT3 film is a radiochromic film produced by International Specialty Products Ashland Inc. (Covington, KY) for clinical dosimetry in 2011. The EBT3 film is very similar in dosimetric performance to previous generation EBT2 but has two major improvements: (1) the symmetric configuration of EBT3 eliminates measurement orientation effects and (2) the EBT3 prevents the formation of Newton rings that usually formed during film scanning [15, 16]. According to the manufacture documents [17], EBT3 film is made by laminating an active layer (28 μm) sandwiched between two identical polyester layers (125 μm each), which makes it more robust and allows water immersion. It is important to note that these films are sensitive to temperature and should be stored in an environment with a temperature below 25°C. It is crucial to avoid exposing them to temperatures above 50 degrees Celsius to ensure their quality and longevity [17, 18]. Several studies have been conducted on Gafchromic film characteristics, demonstrating they are energy- and dose-rate independent [19, 20, 21, 22]. In particular, Sipilä et al. [21] measured the response of EBT3 films to photon (6 MV) and electron beams (6 to 16 MeV) and showed the energy dependence of EBT3 film is uniform within 0.5% across all electron beams. Including the 6 MV and the range of electron energies, the energy dependence of the EBT3 is about 1.1%, making it suitable for dosimetry of mixed photon/electron dose distributions [21]. One of the primary uses of EBT3 is for IMRT/VMAT verification, and multiple studies have indicated that these films are effective tools for this purpose [14, 16, 23]. The gamma index is widely used to compare the 2D/3D dose distribution calculated by TPS with the measured dose. This index compares the measured dose (evaluated dose distribution) point-by-point to that calculated dose (reference dose distribution) based on a distance to agreement (DTA) and a dose difference (DD) criterion [24]. In 2020, the assessment of quality assurance of IMRT plans under a 0.35 Tesla MR guided radiotherapy using Gafchromic™ EBT3 was conducted by Gungor et al. [23]. This study included 70 patients who received treatment with the ViewRay MRIdian® Linac from September 2018 to June 2019. According to their findings, 91% of all the QA analyses demonstrated a passing rate greater than 95% even for a tolerance of 2%/2 mm (%DD/DTA). Based on their analysis, it was determined that EBT3 film dosimetry is a reliable tool for gamma analysis evaluation of treatment plans, even under the presence of a 0.35 Tesla magnetic field [23]. Nevertheless, EBT3 has an optimal range of 0.2 to 10 Gy, which limits its usability. EBT-XD Gafchromic films were developed in 2015, offering optimal performance in the 40–4000 cGy dose range to address this issue. Its design is very similar to EBT3 films, but its active layer thickness is 25 μm. These films are best suited for patient-specific QA of SRS/SBRT techniques where high doses of radiation are delivered to targets [17, 25, 26, 27].

Film dosimetry presents several advantages, notably its water equivalence and the absence of correction factors. Additionally, it offers structural flexibility, protection against water and moisture infiltration, and independence from irradiation angles, boasting high detection resolution. Nonetheless, film dosimetry comes with certain drawbacks, including variations in radiochromic film characteristics, saturation time, and reliance on lot batch consistency. Additionally, film dosimetry depends on landscape and portrait scan orientation, with limitations for peripheral scanner devices, scanning region calibration, and lateral effect. All these disadvantages introduce uncertainties in dosimetry [26, 27, 28, 29, 30, 31, 32]; therefore, one can conclude that film dosimetry may not be the primary option for conducting patient-specific QA.

2.2.3 Detector arrays

Several commercial detector arrays have been developed for performing patient-specific IMRT/VMAT QA. Tables 1 and 2 provide lists of the available array detectors commonly utilized in clinical settings.

Detector Array	MapCHECK2	MapCHECK3	SRS MapCHECK	ArcCHECK
Type of Detector	SunPoint; Diode	SunPoint 2; Diode	SunPoint 2; Diode	SunPoint; Diode
Active Detector Area (mm^2)	0.8 × 0.8	0.48 × 0.48	0.48 × 0.48	0.8 × 0.8
Array Size (cm^2)	26 × 32	26 × 32	7.7 × 7.7	21 × 21
Detector Spacing	7.07 mm	7.07 mm	2.47 mm	10 mm
Number of Detectors	1527	1527	1013	1386
Update Frequency (ms)	50	50	50	50
Inherent Buildup (g/cm^2)	2	1.5	2.75	3.3
Array Geometry	Flat	Flat	Flat	Helical Grid

Table 1.

List of diode arrays manufactured by SUN NUCLEAR corporation [33].

Detector Array	OCTAVIUS detector 729	OCTAVIUS detector 1500	OCTAVIUS detector 1600	OCTAVIUS 1000 SRS
Type of Detector	vented Plane-parallel ion chamber	vented Plane-parallel ion chamber	Plane-parallel, liquid-filled ion chambers	liquid-filled ion chambers
Active Detector volume (mm^3)	5 × 5 × 3	4.4 × 4.4 × 3	2 × 2.5 × 2.5	2.3 × 2.3 × 0.5
Array Size (cm^2)	27 × 27	27 × 27	15 × 15	11 × 11
Detector Spacing center to center (mm)	10	7.1	2.5	2.5
Number of Detector	729	1405	1521	977
Repetition rate (ms)	100	100	100	100
Reference Point (mm)- Below the Surface Array	7.5	7.5	6.9	9
Array Geometry	Flat	Flat	Flat	Flat

Table 2.

List of ion chamber arrays manufactured by PTW dosimetry company [34].

The use of detector arrays reduces the amount of time required for physicists to perform pre-treatment verifications compared to using ion chamber- and film-based patient-specific QA. Many studies have examined various features of detector arrays and showed they are energy, dose rate, and angular independent plus serve as reliable dosimetry tools for pre-treatment verification of IMRT and VMAT [10, 35, 36, 37, 38, 39].

Detector arrays, however, have limited spatial resolutions, which can impact the accuracy of gamma passing rate results in small field dosimetry and detecting delivery errors. In 2018, Bruschi et al. [40] evaluated the effect of detector resolution on SBRT pre-treatment verification results. They compared three detectors (PTW OCTAVIUS 729, 1500, and 1000 SRS) in five different configurations with varying resolutions. These configurations included 729, merged 729, 1500, merged 1500, and 1000 SRS. The study evaluated 150 dose distributions in 30 plans, all of which were designed using Elekta Monaco® 5.0 TPS for a 10-MV X-ray beam on an Elekta Synergy® linac. The planned grid size was 2 × 2 × 2 mm³. Five types of error including systematic variations in collimator angle and gantry angle as well as lack of monitor units were introduced in order to establish the detection sensitivity of the three devices. This study demonstrated that the spatial resolution of the detector plays a vital role in SBRT QA. By increasing the spatial resolution of the detector, they reported that the average GPR value increases while reducing the spread of data. This happens due to small field dosimetry requiring high spatial resolution detectors. Moreover, the significance of detector spatial resolution was also dominant through an error detectability study. It was found that only the SRS array, with the highest resolution among the arrays used, was able to distinguish all the investigated errors [40]. Another study compared three detectors with different resolutions including ArcCHECK, MapCHECK2, and an electronic portal imaging device (EPID) to validate the dosimetric accuracy of the VMAT/IMRT QA [41]. First, they analyzed the performance of the three devices. Then, treatment plans from different treatment sites were assessed to evaluate the reliability of the detectors. Also, systematic variations such as MLC positioning were introduced to compare the detection sensitivity of the detectors. Then, the measurements were compared to the computed plan by TPS (Eclipse from Varian Medical Systems). The gamma analysis criteria of 3%/3, 2%/2, and 1%/1 mm with a dose threshold of 10% to remove the noise were used. Their findings indicated that both ArcCHECK and MapCHECK2 displayed the worst performance compared to EPID. This implies that low-spatial resolution can affect the gamma analysis due to under-sampling of the dose [41].

All these studies showed that the accuracy of the gamma passing rate depends on the detector resolution. Nevertheless, the gamma index method also has its sensitivity and limitations. Hence, it is crucial to carefully choose a detector with an appropriate resolution for the specific field size under investigation to prevent obtaining erroneous results.

2.2.4 Electronic portal imaging device

EPID was originally designed for visualization and patient setup in radiotherapy applications, but later on, it was discovered the digital image produced by EPID contains dose information that can be used for conducting routine QA of Linac and treatment verification. Its popularity and acceptance have been significantly bolstered by its high sensitivity, excellent spatial resolution, and the ability to instantly capture and present the delivered dose in digital format [42, 43]. In addition, analyzing the EPID image is much faster compared to film. The EPID is attached to the gantry of Linac through a robotic arm so that it can intercept the exiting photon beam at all gantry angles to produce a digital image. The image can be acquired and displayed within seconds, which allows one to take multiple images during a treatment session. The EPID made of amorphous silicon (aSi) detector has been widely used among all available EPIDs due to faster image acquisition, high spatial resolution, high sensitivity, compact size, and stable response over time [43, 44]. Recently, Varian Medical Systems (Palo Alto, CA) has designed a new generation of digital megavolt imager called aS1200 EPID, which has a dosimetric active area of 40 × 40 cm² with 1280 × 1280-pixel arrays and a pixel size of 0.0336 cm. The device also includes extra backscatter shielding layers that minimize backscatter artifacts caused by the robotic support arm [43]. Removing the flattening filter from the beam path can increase the dose rate and central axis fluence, which can result in signal saturation [45]. Therefore, Varian Medical System has modified EPIDs for Flattening Filter Free (FFF) beams, ensuring no signal saturation occurs at any distance between the source and detector [42, 43]. This dosimetry system boasts the capability to measure dose rates up to 3200 MU/min, with an impressive acquisition rate of 25 frames per second. Furthermore, its analog-to-digital conversion bit is 16, rendering a separate digitization unit unnecessary for image preparation. Elekta company (Elekta AB, Stockholm, Sweden) has also developed the Elekta iViewGT panel, which is an aSi flat panel x-ray detector (XRD 1642 AP, Perkin Elmer Optoelectronics, Wiesbaden, Germany) and has an active imaging area of 41 × 41 cm² and a resolution of 1024 × 1024 16-bit pixels images with a pixel pitch of 400 μm and a nominal source to EPID distance of 160 cm [46].

Several studies have shown that EPID dosimetry is the most convenient tool for rapid and reliable patient-specific QA and can replace detector arrays in performing patient-specific QA of IMRT and VMAT, which can eliminate the need for phantom setup [47, 48, 49, 50, 51, 52, 53, 54]. In 2019, Torres-Xirau et al. [55] investigated the dose response of EPID in the presence of a magnetic field using Unity MR-linac (Unity, Elekta AB, Stockholm, Sweden). The study concluded that the dose modeling of the EPID in the MR-Linac does not require any additional modifications due to the presence of a magnetic field. In other words, the presence of a magnetic field does not have a significant impact on the EPID responses [55].

2.3 Brachytherapy

Computed tomography or CT is usually used for dose calculation in modern brachytherapy. Although this technique has its merits, it is not enough to accurately assess the amount of radiation dose that reaches the tumor and surrounding organs during each treatment session. That is where in-vivo dosimetry (IVD) comes in and plays a crucial role in this field. In-vivo dosimetry is absorbed dose measurements in the patients while they are being treated. It helps to identify errors caused by equipment failure, dose calculation errors, applicator positioning errors, and changes in patient anatomy. In brachytherapy, IVD serves three purposes. Firstly, it aims to identify major deviations from the treatment plan that may impact the clinical outcome of the procedure. Secondly, it records minor deviations from the plan that cross a certain threshold, making it possible to modify the plan between treatments. Lastly, it also provides a reliable estimate of the actual delivered dose, which is invaluable for maintaining accurate patient records [56]. However, IVD in brachytherapy has several challenges such as high-dose gradients and appropriate positioning of the detector, the energy and dose rate dependence of the detector, and the possibility that an extra invasive procedure may be required. Therefore, the ideal detectors used for IVD should be small in volume to be placed close to the treated target to extract meaningful information and prevent dose averaging in high-dose gradient regions. Moreover, they should have a high signal-to-noise ratio, high spatial resolution, high dynamic range, energy and dose rate independence, reproducibility, linear dose-response over a broad range of energy, directional independence, real-time readouts, and wireless and affordability [57, 58]. The common detectors for clinical settings in brachytherapy are currently thermoluminescent dosimeters (TLDs), optically stimulated luminescence dosimeters (OSLD) (Figure 2), microDiamond, plastic Scintillation Detectors (PSDs), and semiconductor dosimeters (e.g., metal oxide semiconductor field effect transistors (MOSFETs)). Small-volume IC cannot be used since it does not have enough sensitivity to measure dose at centimeter distances from a ¹⁹²Ir source [59]. To measure the depth dependence, angular dependence, and temperature dependence of the detectors, a ¹⁹²Ir high dose-rate (HDR) brachytherapy source is usually used [57, 58].

Figure 2.

Optically simulated luminescence dosimeters (OSLD).

Lithium fluoride (LiF), which is an alkali halide, is commonly used to construct TLDs. There are various physical forms of TLDs such as powder, cubical or cylindrical chips, rods, and pellets [60]. LiF TLD rods have been widely utilized for brachytherapy dosimetry due to easy insertion into catheters [58]. TLDs are the recommended detectors when measuring the AAPM TG-43 dosimetry parameters regarding low-energy photon-emitting brachytherapy sources. Because the validity of the absolute and relative dosimetry results of other detectors has not been persuasively demonstrated [57]. According to TG-43, 1 × 1 × 1 cm³ LiF TLD known as TLD 100 is a valid detector for absolute and relative dose measurements because its total combined uncertainty is low (7 to 9 percent) [57]. TLD-100 chips are LiF crystals doped with titanium and magnesium in order to enhance the number of traps and luminescence centers [60]. TLDs are small, portable, have no bias needed to function, are affordable, cover a wide dose range, and have a small dose rate, energy, and temperature dependency [58]. However, there are some concerns associated with the use of TLDs, primarily the fact that they do not provide real-time measurements, which can lead to inaccuracies in dose measurements [59]. Also, they require a time-consuming preparation and readout process. In 2013, Horowitz and Moscovitch [61] showed that applying different readout parameters can impact the TLD response. Moreover, the positioning of TLD (LiF) for measurement [60, 62] and intrinsic energy dependence between ¹⁹²Ir and megavoltage calibration energy can be another source of uncertainty [62]. In recent years, OSLDs have been utilized more frequently as a potential alternative to TLDs. OSLDs are more affordable, smaller in size, and easier for patients’ application than TLDs. OSLDs are usually placed in a plastic case (10 × 10 × 2 mm³) infused with carbon-doped aluminum oxide (Al₂O₃:C). The detector is disk-shaped, 0.2 mm thick, and approximately 5 mm in diameter, with a mass density of 1.03 g/cm³. The information is captured in an Al₂O₃:C and released by laser stimulation [63], which is much faster, more precise, and more reproducible compared to heating required for reading the TLDs [64]. The OSLD is optically bleached and can be reused multiple times until it reaches a cumulative dose of 10 Gy. The sensitivity decreases considerably once the radiation dose goes beyond 10 Gy [65]. The compact size of the OSLD makes it easy to fit into small and narrow spaces without visible volume-averaging effects.

Another detector that has gained popularity as an online in-vivo dosimeter is MOSFET. The Best® Medical Canada offers 3 different MOSFET dosimeters as standard MOSFET, microMOSFET, and Linear 5ive Array (Figure 3) with the active region of 0.2 × 0.2 mm and width of 2.5, 1, and 1.5 mm respectively. The linear five array is the recommended dosimeter for high-dose-rate and low-dose-rate brachytherapy by the manufacturer [66]. MOSFET is a semiconductor with three terminals: gate, drain, and source. They can be either n- or p-channel, but the p-channel MOSFETs are more common. In the p-channel, the source and drain are made of a p-type semiconductor and isolated from the gate by a layer of silicon dioxide (SiO₂). The gate is constructed using an n-type semiconductor. The sensitivity of a MOSFET detector’s volume is determined by the SiO₂ insulator that captures electron-hole pairs during irradiation. The thickness of the silicon dioxide insulator can be varied between 0.1 to 1 μm, which affects the MOSFET’s sensitivity [67]. Using a MOSFET dosimeter has several benefits, including its ability to accurately measure radiation doses in areas with steep dose gradients and cases of electronic disequilibrium due to its small detector size, lightweight design that poses no harm to patients, low power requirements, user-friendly operation, and real-time online readout [68, 69, 70, 71]. The threshold voltage of its gate varies in proportion to the amount of absorption of radiation. The “threshold voltage” required for these devices is very low, and there is no risk of electrical shock [69]. The readings obtained from MOSFET indicate consistent results, displaying high dose linearity and insignificant fading [70, 71]. Zilio et al. [72] conducted a comparison between the dosimetry result of MOSFET and Monte Carlo calculations and confirmed the accuracy of the MOSFET dosimeter as an absolute dosimeter for brachytherapy. However, they are not water equivalent, have a limited lifetime, and do exhibit slight angular and depth dependence [69, 70]; therefore, care should be taken, and corrections must be applied for clinical applications.

Figure 3.

MOSFET 5 Array. Best™ medical Canada [66]. (with permission).

Plastic Scintillation Detectors (PSDs) have also been used for IVD in brachytherapy [73, 74]. The PSDs are made up of plastic scintillating material with organic scintillating molecules dissolved in a polymerized solvent. This material emits light when exposed to ionizing radiation, and the amount of light produced is directly proportional to the radiation dose. The light is collected by optical fibers and converted into an electric charge that can be read by an electrometer [75]. In other words, the PSDs have four main components including a scintillating fiber that is a sensitive volume, a light pipe that is used to transmit scintillation photons, a photodetector, and an electrometer that is used for readout. PSDs have the advantage of being water equivalent, which makes them superior to inorganic detectors like MOSFETs or OSLDs [75, 76]. They are also flexible and waterproof and have high spatial resolution due to their small size, energy, angular independence, and dose linearity response [74, 75, 76, 77]. Typically, PSDs are paired with photomultiplier tubes to provide real-time readouts [75].

MicroDiamond™ detector or mDD (e.g., Type 60,019 PTW-Freiburg, Germany) is another detector for IVD in brachytherapy that is a synthetic single-crystal diamond detector, with no bias voltage for operating [78]. This innovative synthetic diamond detector offers a perfect combination of the benefits found in natural diamond detectors and silicon diode detectors [79]. MicroDiamond detector is waterproof and has a spatial resolution of 4 mm², an active volume of 0.004 mm³, and a thickness of 0.001 mm [79]. The effective point of measurement is 1 mm away from its tip and has a nominal response of 1 nC/Gy for Co-60. It exhibits a directional response in the water of less than ±0.9% for radial incidence and less than ±1% for axial incidence of ±40°, making it a reliable tool for precise measurements [78]. The mDD was found to have the smallest energy-dependence among diodes and ICs [80]; much better water equivalence than p-type silicon diodes, ICs, and natural diamond detectors (PTW-60003-Freiburg, Germany) [78]; and also a lower absorbed-dose energy dependence compared to Si diodes and TLDs [81]. Moreover, the microDiamond detector displayed no dose-rate dependence effects observed in previous-generation natural diamond detectors [78].

It is worth noting that all detectors mentioned above are also utilized for external dosimetry of photon/electron and proton beams with energies falling within the radiotherapy range in external beam dosimetry.

2.4 Radiopharmaceutical

Radiopharmaceuticals are commonly used for both diagnostic imaging and radiation therapy. While they have the advantage of killing cancer cells, there are some levels of limitations to the use of these drugs due to damaging normal cells as well. Radiation dosages are optimized based on research studies performed on animals and clinical trials on human subjects before approval for clinical applications. Appropriate dosages are based on the careful study of pharmacokinetics, the physical characteristics of the radionuclide, the metabolism of the subject, and the pharmacodynamics of the radiopharmaceutical in animal and human subjects. The dosage is normally adjusted based on the weight of the patient or their total body surface area, with lower dosages recommended for children [82].

2.4.1 Radiopharmaceuticals approved by food and drug administration

Among the currently approved therapeutic radiopharmaceuticals by the Food and Drug Administration (FDA) commonly used ones’ are Sodium iodide-131, Iobenguane iodine-131, Radium-223 dichloride, Yttrium-90 ibritumomab tiuxetan, Yttrium-90 microspheres, Strontium-89 chloride, Samarium-153 lexidronam, Lutetium-177 DOTATATE, and Lutetium-177 vipivotide tetraxetan [83].

Sodium iodine-131 has a half-life of approximately 8 days and decays by beta-particle emission (mean energy of 192 keV) to a stable ¹³¹Xe. Sodium iodide-131 is used to treat thyroid carcinoma and hyperthyroidism (an overactive thyroid). It is available in liquid or capsule form and is absorbed mainly by the thyroid gland. The radiation from radioactive iodine damages the thyroid gland, reducing its activity to normal levels. Radioiodide is used in larger doses after thyroid cancer surgery to destroy any remaining diseased thyroid tissue or to destroy thyroid cancer that has spread to other tissues. Small doses of radioactivity help physicians to determine whether the thyroid gland is working properly or locate tumors caused by certain types of thyroid cancers [84].

Iobenguane-iodine-131 or AZERDA® (Progenics Pharmaceuticals) is a radioactive agent used for the treatment of pheochromocytoma or paraganglioma. Adults and children older than 12 years old with iobenguane scan positive, unresectable, locally advanced or metastatic pheochromocytoma or paraganglioma who require systemic anticancer therapy are common candidates for iobenguane iodine-131 therapy [85].

Radium-223 dichloride, sold under the brand name Xofigo® (Bayer), is used to treat prostate cancer that has spread only to the bony anatomy and is symptomatic and no longer responds to hormonal or surgical treatment that lowers testosterone. Radium-223 dichloride is primarily an alpha particle emitter with a half-life of 11.4 days. The energy range of alpha particles emitted from Ra-223 and its daughter is from 5.0 to 7.7 MeV [86].

Yttrium ibritumomab tiuxetan (⁹⁰Y-IT) or ZEVAKIN® contains a radioactive substance called Yttrium-90 (⁹⁰Y) to treat certain types of B-cell non-Hodgkin lymphoma (NHL) or previously untreated follicular NHL who get a partial or complete response to first-line chemotherapy. Yttrium-90 emits beta particles (2.28 MeV) with a physical half-life of 64.1 hours (2.67 days) [83, 87].

Yttrium-90 microspheres are radioactive particles employed for treating patients with unresectable hepatocellular carcinoma (HCC) or liver cancer. In the United States, the Yttrium-90 Microspheres that are commercially available are TheraSphere® (glass microspheres, Boston Scientific Corporation), and SIR-Spheres® (resin microspheres, Sirtex Medical Ltd) [83].

Samarium-153 lexidronam contains samarium-153, which emits medium-energy beta particles and an imageable gamma photon with a physical half-life of 46.3 hours and maximum energy of 0.808 MeV. Radioactive samarium is used to alleviate bone pain in certain types of cancer by emitting radiation in the affected area.

Strontium-89 chloride (the common name of Metastron) is a radioisotope agent that may be used to diagnose some diseases by studying the function of the body’s organs or be used to treat certain diseases. When used in therapy, it is injected into patients intravenously to deliver radiation to cancer sites and ultimately decreases bone pain. The radioactive strontium is taken up in the bone cancer area and gives off radiation that helps provide relief of pain. Metastron decays by beta emission with a physical half-life of 50.5 days. The maximum beta energy is 1.463 MeV (100%). The maximum range of β- from Strontium-89 in tissue is approximately 8 mm.

Lutetium Lu 177 dotatate (LUTATHERA, Advanced Accelerator Applications USA, Inc.) contains LU-177, which has a physical half-life of 6.647 days and emits beta particles with maximum energy of 0.498 MeV and average energy of 0.133 MeV and gamma rays [88]. Lutetium Lu 177 dotatate is a radiopharmaceutical used to treat adult patients with either metastatic or inoperable cancer known as gastroenteropancreatic neuroendocrine tumors (GEP-NETs) that are positive for the hormone receptor somatostatin, including GEP-NETs in the foregut, midgut, and hindgut [83].

Lutetium-177 (177Lu) Vipivotide Tetraxetan (PLUVICTO®, Novartis/Advanced Accelerator Applications) is a radioconjugate composed of prostate-specific membrane antigen (PSMA)-targeting ligand, combined with Lu-177, a beta-emitting radioisotope, to potentially combat tumor cells that express PSMA.

All radiopharmaceuticals should only be administered by authorized personnel in designated clinical settings after patient evaluation by a qualified physician.

2.4.2 Internal radiation dosimetry

The field of internal radiation dosimetry focuses on measuring the amount of radiation energy deposited in the body’s tissue by radionuclides. This involves examining the physical characteristics of radionuclides, as well as their pharmacokinetics and biokinetics [82]. One of the standard techniques to assess internal radiation doses from the administered radiopharmaceuticals in target organs is known as medical internal radiation dosimetry or MIRD. The MIRD system for calculating doses follows a structured method that utilizes information on the biological distribution of radiopharmaceuticals, the clearance rate of radiopharmaceuticals from the body, and the physical characteristics of radionuclides [82]. This approach is based on the absorbed dose method and provides more accurate results compared to other techniques. The MIRD method uses a simple model of the human body and includes source organs and target organs. The “target organ” is the recipient of the radiated energy from the source organs, and the “source organ” is any organ other than the target organ that contains radiopharmaceuticals. An organ can be both a target and a source [89]. In this case, the energy deposited in that organ is called self-dose.

The absorbed dose is calculated based on Eq. (14).

$$\mathit{\text{Absorbed Dose}}=\frac{\mathit{\text{Energy Absorbed from Ionizing Radiation}}}{\mathit{\text{Mass of Organ}}}$$E14

The unit of absorbed dose in the international system of Units (SI) is “gray” or “Gy,” which is equal to 1-joule energy absorbed per 1 Kg of absorber material. Its traditional unit is “rad” and is used more commonly in the United States. 1 rad is 100 ergs of energy absorbed per gram of medium; therefore, 1 rad = 0.01 Gy.

The amount of absorbed dose is influenced by various factors. These factors consist of the quantity of radioactivity present in the source organ, the duration for which radioactivity resides in the source organ, the type and quantity of radiation energy released by radioactivity in the source organ, and the proportion of the energy emitted by the source that is absorbed by the target organ.

To calculate the absorbed dose using the MIRD method, we need to quantify each abovementioned component step by step.

Step 1. Cumulated Activity, Ã

The amount of radiation received by a target organ depends on the level of radioactivity in the source organ and the time length that the radioactivity resides in the source organ. The product of these two factors is the cumulated activity (Ã) in the source organ, which can be written as Eq. 15 and has a unit of μCi-hr.

$$\tilde{\mathrm{A}}={\int }_0^{\infty }A\left(t\right)\mathit{dt}$$E15

Where A(t) is the radioactivity in a source organ over time of t (Figure 4). A(t) is different from person to person and varies in each organ. However, in human studies, it can be estimated through various methods, such as animal studies, which are then extrapolated with some uncertainty to humans. Additionally, imaging studies in normal human subjects and prior knowledge of the tracer kinetics are also utilized, sometimes in combination, to obtain these estimates [90]. A simplified mathematical model for A(t) can be described as:

Figure 4.

A hypothetical time-activity curve that shows the level of radioactivity in a specific organ over time.

$$A\left(t\right)=A_0{e}^{-{\lambda }_{e}t}$$E16

where ${\lambda }_e={\lambda }_p+{\lambda }_{b.}$

λ_e, λ_p, and λ_b are the effective, physical decay, and biologic decay constants, respectively.

The decay constant can be calculated based on the half-life using Eq. 17.

$$\lambda =\frac{0.693}{T_{1/2}}$$E17

Therefore, cumulated activity is equal to

$$\tilde{\mathrm{A}}=\frac{A_0}{{\lambda }_e}=1.44T_e{A}_0$$E18

where T_e is the effective half-life and A₀ is the initial administered activity.

Step 2. Equilibrium Absorbed Dose Constant, ∆

After calculating the value of Ã, the next step is to determine the amount of radiation energy emitted by this cumulated activity. The amount of energy emitted per unit of cumulated activity in the source organ is given by the equilibrium absorbed dose constant ∆. This factor must be determined for each type of emission for the radionuclide and can be calculated using Eq. 19.

$${∆}_i=k{N}_i{E}_i$$E19

where E_i is the average energy of the emission i^th, expressed in the unit of MeV, and N_i is the relative frequency of that emission. The value of k in the SI unit is 1.6 × 10⁻¹³ and in the traditional unit is 2.13. The units for ∆ is Gy.kg/Bq.sec. and rad. g/μCi.hr. in SI and traditional units, respectively. Therefore, the radiation energy emitted by the source activity cumulated over time is given by àנ∆.

Step 3. Absorbed Fraction, ϕ

The energy absorbed by a target (r_k) from a source region (r_h) irradiation is:

$$\mathrm{Ã}\times {\text{∆}}_{\mathrm{i}}\times {\mathrm{\Phi }}_{\mathrm{i}}$$E20

Where ϕ represents the absorbed fraction, which indicates the proportion of energy released by the source and deposited in the target. The amount of radiation energy that reaches the target organ depends on its composition (e.g., lung, bone) and volume, as well as the distance attenuation between the source and target organs. This means that the absorbed fraction is influenced by the type and energy of the emission, as well as the anatomical relationship between the source and target pair. When performing a dosimetry calculation, it is necessary to determine a value of ϕ for each type of emission from the radionuclide and each source-target pair in the calculation.

The total energy absorbed can be mathematically expressed as:

$$\tilde{\mathrm{A}}\sum _i{\mathrm{\varphi }}_i\left(r_k\leftarrow r_h\right){∆}_i$$E21

The notation ϕ_i (r_k ← r_h) represents the absorbed fraction of energy delivered to the target organ (r_k) from a source organ or region (r_h) for the i^th emission of the radionuclide.

Now, the absorbed dose equation can be written as follows:

$$D\left(r_k\leftarrow r_h\right)=\frac{\tilde{\mathrm{A}}}{\mathit{\text{mass of the target}}\left(m_k\right)}\sum _i{\mathrm{\varphi }}_i\left(r_k\leftarrow r_h\right){∆}_i$$E22

The total dose of the target organ then is the summation of doses from all source organs inside the body.

Step 4. S-Factor

The S-factor is a measure of radiation dose per unit of activity, expressed in Gy/Bq.sec. The S-factor can be found in tables based on the radionuclide, source organ, and target organ. The MIRD pamphlet No.11 [91] tabulated many of the most used radionuclides for the standard phantom (The MIRD phantom is a simplified model of a 70 kg adult male) and expressed them in the conventional unit of rad/μCi.hr. Eq. 22 can be divided into two parts: (a) cumulated activity (Ã) and (b) factors that depend on radionuclide properties and the size and position of different organs in the phantom model. This latter quantity is called the S factor, and it is defined mathematically.

$$S\left(r_k\leftarrow r_h\right)=\frac{\sum _i{\mathrm{\varphi }}_i\left(r_k\leftarrow r_h\right){∆}_i}{m_k}$$E23

If we define a specific absorbed fraction as follows:

$$\Phi =\frac{{\mathrm{\varphi }}_i}{m_k}$$E24

then dose equation can be simplified as:

$$\overline{D}\left(r_k\leftarrow r_h\right)=\tilde{\mathrm{A}}\times S\left(r_k\leftarrow r_h\right)$$E25

where S is S-factor, or the mean absorbed dose per cumulated activity.

Step 5. Effective Dose Calculation

The effective dose to the whole body is given by Eq. (26).

$$E=\sum _T{W}_T{D}_T{W}_R=\sum _T{W}_T{H}_T$$E26

The unit used for measuring the effective dose, E, is Sievert. D_T and W_T are the average absorbed dose and tissue equivalent factor for organ T. W_R is equal to unity for all radiations used in diagnostic nuclear medicine such as gamma rays, X-rays, electrons, and positrons. The equivalent dose, H_T, specifies a quantity that considers the relative biologic damage caused by radiation with a particular tissue or organ and is given by Eq. (27).

$$H_T=D_T{W}_R$$E27

The unit of H_T is also Sievert, so care must be taken to use this unit as it is associated with both equivalent dose and effective dose. The effective dose is primarily designed for assessing radiation risks and measuring radiation doses received by workers in the radiation industry. It can also be applied to clinical nuclear medicine. The effective dose demonstrates the total body dose that would cause the same overall risk as the nonuniform dose distribution delivered. This is accomplished by assigning different weighting factors to the doses received by individual organs [89].

3. Proton therapy dosimetry

The primary distinction between photon and proton beams lies in the inherent physical characteristics of the proton beam. Protons are directly ionizing charged particles with a positive charge and a finite range. Consequently, proton and photon beams have a completely different dose distribution. Unlike photons, protons release an immense amount of energy at the end of their path and produce a maximum peak near the end of their range. The maximum peak is called the “Bragg peak,” which is a distinct property of protons, and the proton dose falls to zero instantly beyond the Bragg peak. This can be used to increase the dose to targets while minimizing radiation exposure to normal tissue [92, 93]. Correspondingly, it is a crucial task to determine the location of the dose deposition to ensure the target, not OARs, received the prescribed dose. This information can be extracted using IVD. In both brachytherapy and proton therapy, various detectors play a crucial role in ensuring accurate and safe treatment delivery. These detectors include MOSFET (Metal-Oxide-Semiconductor Field-Effect Transistor), PSD (Position-Sensitive Detector), OSLD (Optically Stimulated Luminescence Dosimeter), TLD (Thermoluminescent Dosimeter), and microDiamond. MOSFET detectors are valued for their real-time response and sensitivity, making them suitable for measuring dose distribution during treatment. PSDs are used to determine the position of charged particles in the proton beam, contributing to beam monitoring and verification. OSLDs and TLDs, as passive detectors, are employed to measure the absorbed dose of ionizing radiation in both modalities. MicroDiamond detectors, with their high spatial resolution, offer valuable insights into dose delivery and beam characteristics in proton therapy. While the applications of these detectors are similar in both brachytherapy and proton therapy, their specific utilization may vary depending on the treatment facility’s protocols and equipment. Up-to-date research and consultation with experts in the field are essential to stay informed about the latest advancements and best practices [94, 95, 96, 97, 98]. However, these approaches are only applicable when the detectors could be placed on the skin or accessible body cavities (e.g., in prostate cancer where dosimeters can be located inside the rectum).

The dose or range of proton can also be monitored using imaging methods such as positron emission tomography (PET), prompt gamma (PG) imaging, ionoacoustic imaging, and follow-up MRI. Due to the lack of exit dose in proton therapy, EPID cannot be used.

The in-vivo range verification with PET can be performed online (in-beam) monitoring or offline. In the online technique, a small field-of-view dual head PET is attached to the gantry, allowing taking images during or immediately after treatment, while in the offline method, the patient is transferred to the PET/CT scanner room after treatment. The reason for using PET is that protons collide with atoms in tissues (non-elastic collision), causing nuclear reactions that create positrons (⁺β). Combining the information from the PET image with the patient’s anatomy can provide data to determine the location of the positron emitters in the patient’s body to identify the path of the proton beam and monitor the location of the tumor and dose deposition path [99, 100, 101].

Prompt gamma imaging is the second approach for providing real-time in-vivo range verification of the proton using a Compton camera [102, 103, 104, 105, 106]. The camera detects MeV prompt γ-rays emitted from excited nuclei along the proton path to create an image using two stages of detectors. The first stage includes photon-scattering detectors, while the second stage has photon-absorbing detectors [105]. Data obtained from the energy deposition and spatial coordinates of gamma rays, resulting from at least two interactions in the stages of Charged-Particle Cancer Therapy (CC), could potentially be utilized to precisely determine the range of the ion beam with an accuracy of up to 1 millimeter. In this context, Charged-Particle Cancer Therapy refers to a treatment modality that employs charged particles, such as protons or heavy ions, to target cancer cells with high precision and minimal damage to surrounding healthy tissues. When these charged particles interact with the patient’s tissue, they produce secondary particles, including gamma rays, through nuclear interactions. By carefully analyzing the energy and spatial information of these gamma rays from multiple interactions, medical physicists and radiation therapists can gain valuable insights into the distribution of the ion beam within the patient’s body. This data can be processed using advanced algorithms and computational techniques to accurately determine the range of the charged particles, helping to ensure that the tumor receives the prescribed radiation dose while sparing nearby healthy tissues. This level of accuracy in determining the ion beam range is of utmost importance in radiation therapy to maximize the treatment’s effectiveness and minimize potential side effects. It allows for precise treatment planning and delivery, leading to better outcomes for cancer patients undergoing Charged-Particle Cancer Therapy [104, 107].

A third technique for range verification in proton therapy is ionoacoustic, a nonnuclear technique [107, 108, 109]. This approach is a direct technique to localize the Bragg peak based on the thermoacoustic effect. The local energy deposition of a pulsed ion beam within a short time (μsec) in a limited tissue volume (mm³) can generate an ultrasound signal (ionoacoustic signal) with a frequency in the of range 0.1 to 10 MHz. Detecting this signal allows the user to precisely localize the Bragg peak’s position [107, 108].

The fourth approach is to use MRI under certain conditions to determine in-vivo range verification of proton [110, 111, 112, 113]. One well-known example is the use of MRI to detect the fatty conversion in the vertebral bone marrow after proton therapy to visualize the proton beam and its range [111]. A similar observation has also been reported for changes in the liver [110]. This technique is a late approach, usually 3 to 6 months after proton therapy [110].

In summary, these imaging techniques for in-vivo range verification in proton therapy look promising, but they are still in the research and developmental phase.

4. Neutron therapy dosimetry

Neutrons can penetrate tissues deeply and can interact with atomic nuclei, releasing energy that damages cancer cells and have a higher relative biological effectiveness (RBE) compared to photons. This means that neutron radiation can cause more significant damage to cancer cells per unit of absorbed dose, potentially improving treatment outcomes.

The three different techniques for neutron therapy in radiation oncology are: (a) ²⁵²Cf neutron brachytherapy, (b) Fast neutron therapy (FNT), and (c) Boron neutron capture therapy (BNCT).

4.1 ²⁵²Cf neutron brachytherapy

The isotope Californium-252 is a source of gamma/neutron radiation with a half-life of 2.645 years and was introduced for brachytherapy in the late 1960s and early 1970s [114]. This form of brachytherapy is particularly useful for treating deep-seated tumors that are challenging to reach with other radiation sources. ²⁵²Cf was usually used in combination with external photon beam therapy since it was discovered that irradiating tumors with ²⁵²Cf before photon irradiation is more effective [115]. Currently, ²⁵²Cf is not being clinically implanted due to its large size and low activity levels [114].

4.2 Fast neutron therapy

Fast neutron therapy was common during the 1980s and 1990s and became an available option to treat certain tumors such as advanced prostate cancer [116], breast cancer [117], and bone and soft tissue sarcomas [118, 119, 120]. Fast neutrons are the production of one of the nuclear reactions such as fusion interaction (d + T or d + D), stripping reactions (d + Be), and inelastic interactions (p + Be). Among all these nuclear reactions, p + Be is the most preferable option for modern, high-energy facilities [121].

4.3 Boron neutron capture therapy

Boron neutron capture therapy (BNCT) is a radiotherapy technique to eradicate tumors such as melanoma, brain tumors, and head and neck cancers. This treatment is based on the nuclear reaction known as boron neutron capture (¹⁰B (n, α)⁷Li) in which boron-10 absorbs low thermal neutrons (<0.5 eV), resulting in the production of helium nuclei (⁴He₂) and recoiling lithium-7 (⁷Li₃) atoms [122, 123]. These particles deposit energy in a short range (<10 μm) due to the high linear energy transfer, which results in substantial damage to malignant cells accumulated with boron-10 [124].

4.4 Neutron therapy dosimetry

Neutron fields are usually a mixture of photon and neutron, which makes the dosimetry more complicated. Neutrons like photons ionize indirectly; however, they have different RBE; hence, it is necessary to measure them separately [125, 126]. In neutron dosimetry, the energy of the neutron and the composition of the tissue (hydrogen, carbon, nitrogen, oxygen, calcium, and phosphorus) for dosimetry are key factors. Because the probability of interaction between the neutron and the element within the body depends on the neutron energy and the target nuclide. Neutrons are classified based on their energies nevertheless; there is no agreement to classify them precisely. The following is an approximation:

• Thermal (0.025 eV)

• Slow (<10 eV)

• Intermediate (10 eV – 100 keV)

• Fast (>100 keV)

The recommended dosimetry method to measure the absorbed dose in fast-neutron fields is described by the international commission on radiation units and Measurements (ICRU) protocol No. 45 using an ion chamber [127]. The tissue equivalent material should be used for the wall, gas, electrode, and water for the phantom. The absorbed dose is computed based on the application of the Bragg-Gray cavity theory (Eq. 28).

$$D=\frac{Q_n}{m}\times \frac{W_n}{e}{\left(r_{m,g}\right)}_n{d}_T\frac{1}{1+\delta }$$E28

Where Q is the charge of one sign produced in the cavity, and W_n is the average amount of energy required to form an ion pair. This value is calculated as a mean for all the charged particles that are produced during neutron interactions in the gas material being used. (r_m,g)_n is the gas-to-wall absorbed-dose conversion factor; d_T is the displacement correction factor, and δ is a correction factor to account for the difference in the response of the detector, such as the ionization chamber, to different types of radiation, such as neutrons and photons [127]. Another detector is a low-pressure Tissue Equivalent Proportional Counter (TEPC), which can be used for BNCT measurements and is able to measure the absorbed dose and information about the microscopic nature and beam quality [128]. However, collecting data with TEPC is slow and time-consuming. To obtain beam data more efficiently, a set of three ion chambers including an A-150 TEP chamber, an Mg chamber, and another Mg chamber filled with B-10 on its inner surface were usually utilized [128]. The Si diodes can also be used for neutron dosimetry since their sensitivity to photons is negligible. They also have higher spatial resolution compared to ion chambers, which can improve penumbra measurements [125, 129]. Another method used for in-vivo dosimetry of neutrons includes the use of activation detectors. These detectors contain materials that activate when exposed to neutrons, resulting in the emission of gamma rays. The intensity of the gamma rays can be measured, indicating the neutron dose. TLDs and solid-state detectors can also be used for in-vivo neutron dosimetry.

Unfortunately, due to advanced development in photon and proton therapy, neutron therapy is not a common cancer therapy technique at the time of this writing. Therefore, obtaining updated information about the dosimetry of this treatment technique is quite challenging.

5. Conclusion

In this chapter of the book, we have presented a concise outline of sophisticated methodologies and tools employed in the field of dosimetry, primarily emphasizing the assessment of photon and electron dosages. Distinctive detectors exhibit variations in their performance across diverse dosimetric situations, encompassing factors like small versus large fields, low versus high spatial detector precision, or the use of photon versus proton beams. The utilization of an inappropriate detector could result in inaccurate quality assurance outcomes. Consequently, it is imperative for medical physicists to possess a comprehensive understanding of detector constraints when quantifying doses, thereby ensuring precision in clinical applications.