Physics in Medicine & Biology     

TOPICAL REVIEW

Review of plastic and liquid scintillation dosimetryfor photon, electron, and proton therapyTo cite this article: Luc Beaulieu and Sam Beddar 2016 Phys. Med. Biol. 61 R305

 

View the article online for updates and enhancements.

Related contentCurrent status of scintillation dosimetry formegavoltage beamsL Beaulieu, M Goulet, L Archambault et al.

-

Development of a novel multi-point plasticscintillation detector with a single opticaltransmission line for radiation dosemeasurementFrançois Therriault-Proulx, LouisArchambault, Luc Beaulieu et al.

-

A method to correct for temperaturedependence and measure simultaneouslydose and temperature using a plasticscintillation detectorFrancois Therriault-Proulx, LandonWootton and Sam Beddar

-

Recent citationsA tomographic UV-sheet scanningtechnique for producing 3D fluorescenceimages of x-ray beams in a radio-fluorogenic gelTiantian Yao et al

-

Innovation and the future of advanceddosimetry: 2D to 5DMark Oldham

-

Real-time dosimetry with Yb-doped silicaoptical fibresIvan Veronese et al

-

This content was downloaded from IP address 54.70.40.11 on 25/10/2017 at 21:52

R305

Physics in Medicine & Biology

Review of plastic and liquid scintillation dosimetry for photon, electron, and proton therapy

Luc Beaulieu1,2 and Sam Beddar3,4

1 Département de physique, de génie physique et d’optique et Centre de recherche sur le cancer, Université Laval, Québec, QC, Canada2 Département de radio-oncologie et CRCHU de Québec, CHU de Québec, QC, Canada3 Department of Radiation Physics, The University of Texas MD Anderson Cancer Center, Houston, TX, USA4 The University of Texas Graduate School of Biomedical Sciences at Houston, Houston, TX, USA

E-mail: Beaulieu@phy.ulaval.ca and abeddar@mdanderson.org

Received 28 March 2016, revised 20 June 2016Accepted for publication 21 July 2016Published 3 October 2016

AbstractWhile scintillation dosimetry has been around for decades, the need for a dosimeter tailored to the reality of modern radiation therapy—in particular a real-time, water-equivalent, energy-independent dosimeter with high spatial resolution—has generated renewed interest in scintillators over the last 10 years. With the advent of at least one commercial plastic scintillation dosimeter and the ever-growing scientific literature on this subject, this topical review is intended to provide the medical physics community with a wide overview of scintillation physics, related optical concepts, and applications of plastic scintillation dosimetry.

Keywords: scintillators, plastic scintillation dosimetry, Cerenkov subtraction, small field dosimetry, in vivo dosimetry, 1D, 2D and 3D dosimetry, photon, electron and proton dosimetry

(Some figures may appear in colour only in the online journal)

1. Introduction

Scintillators are some of the oldest ‘real-time’ radiation detectors. The first scintillating detector, called a spinthariscope, was originally used by Sir William Crookes in the early 20th century. This detection system consisted of a zinc sulphide screen that produced a

L Beaulieu and S Beddar

Printed in the UK

R305

PMB

© 2016 Institute of Physics and Engineering in Medicine

2016

61

Phys. Med. Biol.

PMB

0031-9155

10.1088/0031-9155/61/20/R305

Topical Review

20

R305

R343

Physics in Medicine & Biology

IOP

Institute of Physics and Engineering in Medicine

0031-9155/16/20R305+39$33.00 © 2016 Institute of Physics and Engineering in Medicine Printed in the UK

Phys. Med. Biol. 61 (2016) R305–R343 doi:10.1088/0031-9155/61/20/R305

R306

Topical Review

small amount of visible light when struck by alpha particles; the light could be detected by the human eye through a microscope in a dark room (Birks 1964). Of course, modern scintillator-based systems now rely on advanced scintillating materials and photodetector devices (Knoll 1999). Inorganic scintillators, with their high photon interaction cross sec-tions and high light-production efficiency, have become the cornerstone of nuclear medi-cine imaging (Bushberg et al 2011), and copper-doped quartz sensors (Justus et al 2004) and Al2O3 crystals (Aznar et  al 2004, Andersen 2011) have been used with fiber optic-coupled laser stimulated readout systems for real-time dose measurements in radiotherapy. However, the high-Z materials of which these scintillators are made pose challenges for dosimetry applications.

Organic or plastic scintillators composed mainly of hydrocarbon molecules have interac-tion properties that are similar to those of water and human tissues, especially in the Compton-dominated energy range of the standard photon beams used in radiation therapy. It has been known for a long time that the light production in these scintillators is linearly proportional to the energy deposited in them by photons and electrons but independent of those particles’ initial energies (above ~125 keV) (Birks 1964). These interesting properties apply not only to solid plastic scintillators but also to liquid and vapor forms. Thus, in theory, plastic scintil-lation detectors (PSDs) or liquid scintillation detectors should require no correction factors when used in photon- and electron-beam dosimetry. However, for charged particles heavier than electrons, light output from plastic scintillators depends on linear energy transfer (LET), so a correction would be needed for proton beams in the Bragg peak region (Archambault et al 2008, Beddar et al 2009, Robertson et al 2013).

One of the first applications to employ the above-mentioned properties explicitly for radia-tion dosimetry was developed by Murai et al (1964), who used a PSD for integral dose meas-urements. Ge (1988) developed the earliest 2-dimensional (2D) PSD dosimetry system for use in radiotherapy, consisting of a scintillator sheet placed in a water tank and a video camera to measure the light emitted from the sheet. He also compared percent-depth dose curves (PPDs) and beam profiles of linac photon and electron beams measured with this PSD system to measurements obtained from ion chambers. While the two dosimeters tended to agree well overall, important differences were also noted but not fully explained at the time. The very first systematic study of PSDs as radiation dosimeters was conducted by Beddar et al (1992b, 1992c) and published in the form of two articles that are now considered seminal works in the field. Interest in scintillation dosimetry has increased tremendously since then, especially in the last decade, as researchers have begun to explore the potential of PSDs to solve outstand-ing challenges in small-field dosimetry, in vivo dosimetry, and real-time volumetric dosimetry (Beddar and Beaulieu 2016).

Although the light produced in a PSD is directly proportional to dose, light propagation, light collection, and various optical and physical phenomena (such as Cerenkov parasite light) pose numerous challenges for accurate and reproducible dosimetry. Light production in most plastic scintillators ranges from 7000 to 10 000 photons MeV−1. This represents a W equiva-lent of 100–140 eV/photon and an efficiency of about 2.5%. Therefore, a crucial element in PSD development is the optimization of the optical chain such that a maximum number of the emitted photons are collected while the linearity-to-dose of the original signal is preserved. This review will detail the intrinsic properties of plastic scintillators and survey the state of the art in the development of PSDs and their applications to photon, electron, and proton dosimetry.

Phys. Med. Biol. 61 (2016) R305

R307

Topical Review

2. Scintillation dosimetry

2.1. Basic physics properties

Most commercially available organic scintillators are, at the basic level, 2-component systems consisting of a bulk or solvent medium (e.g. polyvinyltoluene (PVT) for plastic scintillators or polystyrene for plastic scintillating fibers) and an organic fluor made of a long aromatic or cyclical hydrocarbon chain. Plastic scintillators produce light through interactions with charged particles and a mechanism of energy exchange between the bulk and organic fluor molecules called Förster resonance energy transfer or fluorescence resonance energy transfer. These interactions elevate aromatic molecules in the scintillator to excited vibrational and rotational energy states. The excited molecules then decay by a number of routes, including (1) fluorescence, a fast decay (on the order of nanoseconds) to the ground state via the emis-sion of a photon; (2) phosphorescence decay to a metastable state followed by slow decays to the ground state; or (3) delayed fluorescence, a return to the original excited state by ther-mal activation and then direct decay to the ground state over a longer time scale. Organic scintillators emit light at a lower frequency than they absorb it (a phenomenon known as the Stokes shift). Thus, scintillators are largely transparent to their own emitted light (Birks 1964). Finally, plastic scintillators emitting in the green or red wavebands are usually composed of an additional fluorescence compound, a secondary fluor, absorbing the blue–violet emit-ted lights and reemitting in the higher wavelength region. These are called ‘three- comp-onents’ system.

Plastic scintillators are similar to water in both their physical characteristics and their inter-actions with ionizing radiation. For example, the BC-400 scintillator (Saint-Gobain Crystals, Cleveland, OH, USA) used by Beddar et al (1992b, 1992c) has a density of 1.032 g cm−3 com-pared to 1.000 g cm−3 for water, a ratio of electrons to molecular weight of 0.5414 e- u−1 com-pared to 0.5551 e- g−1 for water, and an electron density of 3.272 e- g−1 compared to 3.343 e- g−1 for water. The BC-400 scintillator is composed of 8.47% hydrogen and 91.53% carbon by weight, while water is 11.19% hydrogen and 88.81% oxygen. Its mass energy absorption coefficient, mass collision stopping power, and mass angular scattering power are also nearly identical to those of water above 100 keV (figure 1).

Importantly for dosimetric purposes, the light emitted by plastic scintillators is directly proportional to the energy deposited by electrons with energies greater than approximately 125 keV. Thus, plastic scintillators are energy independent at the energies found in clinical electron beams and photon beams, which deposit their energy indirectly via electrons ionized in Compton interactions. For larger charged particles, such as protons, the threshold energy for a linear response is much higher (Beddar et al 1992b, 1992c).

In the above-mentioned two papers, Beddar et al also established a host of other physi-cal characteristics of the kinds of scintillators used for medical dosimetry. The scintillators’ response was found to be independent of total dose, dose rate, and angle of incidence. The scintillators exhibited relatively robust radiation hardness, losing only about 3% of their out-put after exposure to 10 kGy of irradiation. The light emission of the scintillators used for dosimetry was also found to be highly reproducible, with variations of much less than 1% between measurements.

Beddar et al (1992c) also investigated the temperature dependence of BC-400 scintilla-tor. They found a slight increase in output as temperature increased from 0 °C to 45 °C, small enough to be ignored without introducing more than 0.5% error into the measured dose. However, Wootton and Beddar (2013) and Buranurak et al (2013) have recently demonstrated that this result does not hold for all scintillators. Specifically, BCF-12 and BCF-60 scintillating

Phys. Med. Biol. 61 (2016) R305

R308

Topical Review

fibers exhibit a small decrease in output as temperatures increase. The effect can still be safely ignored for temperatures within a few degrees of the original temperature at calibration, but for some applications (such as in vivo dosimetry), this effect must be explicitly accounted

Figure 1. Radiation interaction characteristics of water, scintillator, and polystyrene plotted side by side. Plastic scintillator exhibits traits very similar to those of water (from Beddar et al (1992b)).

Phys. Med. Biol. 61 (2016) R305

R309

Topical Review

for. The temperature dependence of other scintillators and scintillating fibers requires further investigation.

2.2. Experimental validation

Beddar et  al (1992c) incorporated plastic scintillators into a dosimetry system to validate their usefulness for medical radiation dosimetry. They acquired percent depth-dose curves for 60Co and 10 and 18 MV photon beams and 6, 12, and 18 MeV electron beams and compared the results to measurements taken with an ion chamber. The PSD’s readings exhibited excel-lent agreement with those of the ion chamber (figure 2). They also captured a beam profile of a 60Co beam and compared it to results acquired with film and with an ion chamber. The PSD agreed nearly exactly with the film and differed only slightly from the ion chamber in the steepest regions of the beam penumbra; this difference can be accounted for by volume averaging of the ion chamber (figure 3). PSDs, therefore, have been found to exhibit excellent spatial resolution.

2.3. Component of a PSD and design optimization

A PSD comprises a scintillator or scintillating fiber, a photodetector for quantifying the light emitted by the scintillator, and a light guide for transporting light from the scintillator to the photodetector. The original PSD constructed by Beddar et al (1992b) consisted of a BC-400 scintillator optically coupled to a silica light guide with optical grease. A photomultiplier tube (PMT) was used to quantify the light output (figure 4). Since then, a great deal of research has been directed toward determining the optimal choice of materials to maximize the utility of PSDs.

While the BC-400 scintillators employed in Beddar et al early experiments are still used, BCF-12 and BCF-60 are also commonly used now. The latters are scintillating fibers, which differ from scintillators in that they have a thin layer of polymethyl methacrylate (PMMA) cladding. Archambault et al (2005) compared these and a few other scintillating fibers to tra-ditional scintillators and found that the cladding enhanced light collection by facilitating total internal reflection. BCF-60 is also used for its spectral properties: it has a tertiary wavelength-shifting component to convert blue light to green. This provides spectral separation between the emitted scintillation spectrum and the majority of the Cerenkov radiation (see section 3) produced in the light guide (Beddar et al 1992a). Ayotte et al (2006) demonstrated the value of polishing the abutting ends of the scintillator and optical fiber and of adding a layer of reflective material at the terminal end of the scintillator. They also researched optical coupling agents and found that nonpermanent couplings had the most efficient transmission; nonethe-less, they advised using permanent couplings to make a more robust detector. Specifically, they recommended using cyanoacrylate adhesive because of its high transmission efficiency and cost effectiveness (Ayotte et al 2006).

Silica light guides are now rarely used, having been replaced with plastic optical fiber. Plastic optical fiber is superior because it is more closely water equivalent, less expensive, and less brittle than silica fiber, resulting in a more robust detector. However, it has the dis-advantage of having much higher attenuation (Archambault et al 2005). Lambert et al (2008) have experimented with a hollow air-core fiber to address the problem of Cerenkov radiation produced in the fiber. The result is a detector that produces almost no Cerenkov contamination but is more fragile and has high light attenuation (Lambert et al 2008, 2010, Liu et al 2011, Archambault et al 2012b).

Phys. Med. Biol. 61 (2016) R305

R310

Topical Review

Figure 2. Depth dose curves for a range of photon beam qualities. Excellent agreement between ion chamber and scintillator is observed for each beam at all depths (from Beddar et al (1992c)).

Phys. Med. Biol. 61 (2016) R305

R311

Topical Review

Many groups favor the use of charge-coupled device (CCD) cameras over PMTs because CCD cameras can collect signals from multiple PSDs simultaneously (Archambault et al 2006, 2007). CCD cameras also do not require a high-voltage power supply, which is susceptible

Figure 3. Profile of a beam penumbra taken with ion chambers, calibrated film, and a plastic scintillation detector. The scintillator display better agreement with film than the ion chambers due to its small size (and thus less volume averaging) (from Beddar et al (1992c)).

Figure 4. A diagram of the original PSD used by Beddar et al. Two PMTs collect signal from a scintillator and background fiber to isolate scintillation light from Cerenkov light (from Beddar et al (1992b)).

Phys. Med. Biol. 61 (2016) R305

R312

Topical Review

to drifting. They come with their own computer-controlled cooling and acquisition systems. CCD cameras come either in monochrome or with built-in RGB (Bayer pattern) color filters for (limited) spectral analysis. Furthermore, various technologies can enhance CCD sensitiv-ity, such as intensified CCD and electron-multiplying CCD (Lacroix et al 2009, 2010a).

Finally, photodiodes are also beginning to be used as photodetectors; the only currently available commercial PSD, the Exradin W1, uses a photodiode (Archambault et  al 2013). Photodiodes are advantageous because they can be read with an electrometer and are much less expensive than CCDs and PMTs. Results generated with photodiode-based PSDs are very promising thus far (Fontbonne et al 2002, Beierholm et al 2014, Carrasco et al 2015a). It should be pointed out, however, that PMTs are not obsolete: Liu et al have used a PMT array that allowed very fast acquisition of signals from multiple plastic scintillators simultaneously and with excellent sensitivity (Liu et al 2012). A systematic study of photodetector properties, including signal-to-noise ratio as a function of dose and dose rates over a wide photon energy range, has been conducted by Boivin et al (2015b). That study also provided recommenda-tions for photodetector choice based on the specific use intended for the PSD.

3. The Cerenkov effect and its solutions

Plastic scintillators have many desirable properties for dosimetry applications, however they are susceptible to detrimental stem effects, especially the emission of Cerenkov light from the optical guide fiber. Most optical guide materials, whether glass or plastic, have a relatively high index of refraction, n, well above the refractive index of vacuum, 1. Since the speed of light c in a material is influenced by the index of refraction, namely as c/n, electrons, either those emitted directly from high-energy beams or secondary electrons set in motion by pho-ton beams, travel faster than the speed of light in that material. The relationship between the threshold energy and the index n is simply given by:

⎛

⎝

⎜⎜⎜

⎞

⎠

⎟⎟⎟

=−

−E mc1

11 ,

n

th2

12

(1)

where the electron kinetic energy β = v/c = 1/n. For a medium with n = 1.5, that threshold is reached at around 175 keV. Note that the higher the refractive index, the lower the threshold energy. In a dielectric medium, traveling electrons polarize the medium, leading to dipole oscillations from perturbation of the orbital electrons. The relaxation process leads to emis-sion of visible light. Generally, the quantity of this emitted light is too small to be detected. However, since the light in the medium is moving more slowly than the exciting electrons, there is constructive interference in the direction of motion of the electrons. This is called the Cerenkov effect (Cerenkov 1937, Tamm and Frank 1937), after the 1958 Nobel laureate Pavel Alekseyevich Cerenkov. It appears as a violet–blue light and is commonplace in nuclear reac-tors, where spent fissile materials are placed in pools of water for cooling. Note that the refrac-tive index n of water is about 1.33, so the corresponding threshold energy is about 260 keV.

The Cerenkov emission spectrum peaks in the blue–violet region but is present at all vis-ible wavelengths; it follows a λ−3 intensity curve. Cerenkov light emission in the optical guide adds to the scintillation light emitted by the PSD sensor. As a form of stem effect, it must be removed for accurate dosimetry. The intensity of Cerenkov light emitted in the optical guide is up to 2 orders of magnitude per mm lower than the intensity of the scintillation light produced by the scintillator. However, the length of the light guides within the radiation field is usually much longer than the scintillation probes—several centimeters versus a few millimeters at

Phys. Med. Biol. 61 (2016) R305

R313

Topical Review

most for the scintillators. Therefore, the Cerenkov stem effect can account for a large fraction of the total light detected. In addition, the fraction of light accounted for by the Cerenkov effect changes depending on the irradiation geometry (Beddar et al 1992a, Archambault et al 2006) because Cerenkov light is emitted at an angle from the direction of motion of the elec-trons (cosθ = 1/βn). Therefore, the coupling of the Cerenkov light into an optical guide, in particular an optical fiber, is maximized at specific angles between the beam direction and the optical guide, from 40 to 50° as shown in figure 5 (de Boer et al 1993). Therefore, an efficient PSD system will maximize the collection of light from the scintillator and minimize or elimi-nate the contribution of Cerenkov light to the total light collected (Beddar et al 2003, 2004).

Another stem effect is the direct excitation of the polymer chain, or fluorescence, from the plastic optical fiber guides. The production of this form of light has been found to be orders of magnitude lower than that of Cerenkov light, but it is the dominant stem effect at lower ener-gies, where Cerenkov light emission is not possible. In most cases, however, it is negligible compared to scintillation light production (Therriault-Proulx et al 2013a).

3.1. Cerenkov light removal techniques

For high-energy beams, the production of Cerenkov light per unit of volume is proportional to dose. If the volume of the optical guide was to remain constant and perpendicular to the beam throughout the irradiation process, the Cerenkov light would simply remain a constant propor-tion of the total light collected. In that case, it would not need to be removed. However, many applications necessitate geometries in which different amounts of materials are irradiated—certainly different than the standard calibration conditions (e.g. a 10 × 10 cm2 field in a fixed beam configuration, i.e. 0°, at dmax). Because the amount of Cerenkov light that is superim-posed on the scintillation signal can vary greatly for the same delivered dose and can change the total amount of light detected, the Cerenkov signal must be removed. Several methods have been proposed to accomplish this.

3.1.1. Two-fiber subtraction. Beddar et al proposed the first method to efficiently remove the Cerenkov stem signal, a two-fiber approach, shown in figure 4 (Beddar et al 1992b). They recognized, as depicted in figure 6(a), that the total visible light produced is a superposition of two components (neglecting the even lower intensity fluorescence light): scintillation light and Cerenkov light. In this method, two nearly identical assemblies were used side-by-side; one was composed only of the optical guide without a scintillator. Any spurious light pro-duced in the optical guide could therefore be measured independently and subtracted from the measurements taken by the guide-plus-scintillator assembly, leaving only the scintillation sig-nal. This setup assumed that both assemblies would see the same fluence; this condition is eas-ily met for in-field measurements in standard (large) fields, where no strong dose gradient is present. Many authors have replicated this concept over the years (Flühs et al 1996, Lee et al 2007, Ishikawa et al 2009, Lambert et al 2010). Létourneau et al (1999) applied this method to radiosurgery measurements, placing the two assemblies on top of each other instead of side-by-side to minimize the effect of the strong gradient in the small fields used in radiosurgery. The two-fiber subtraction method has become the reference method to compare and validate other new Cerenkov light removal techniques under standard irradiation conditions.

3.1.2. Basic spectral filtering. Basic spectral or optical filtering was first proposed by de Boer et al (1993) to take advantage of the fact that some plastic scintillators emit light in a wavelength region far away from the blue–violet region that dominates Cerenkov light. With an appropriate optical filter (or a photodetector that is sensitive only to the scintillator’s

Phys. Med. Biol. 61 (2016) R305

R314

Topical Review

wavelength), the scintillation light would predominate over the Cerenkov signal (figure 6(b)). This technique has the distinctive advantage over the two-fiber method of requiring only a single light guide-scintillator assembly, making it useful for measurement conditions involv-ing large dose gradients. de Boer and colleagues showed that this method could remove over 50% of the Cerenkov signal. However, because Cerenkov light is emitted at all wavelengths, it was impossible to completely remove it. Clift et al (2000) further optimized this approach and obtained even better filtration of Cerenkov light. Archambault et al (2006) compared the two-fiber technique (figure 7), basic spectral filtering, and the chromatic removal approach (described below) under various irradiation conditions. They demonstrated that the basic fil-tering method was the least effective of the three, with deviations of up to 8% from an ion chamber measurement for large fields, but less than 3% for fields smaller than 10 × 10 cm2.

3.1.3. Chromatic removal technique. Fontbonne et  al (2002) proposed a novel approach to Cerenkov stem effect removal that expanded the basic optical filtering technique to two wavelengths. Since the total light measured in a PSD assembly is, in its simplest form, the superposition of two signals, Fontbonne and colleagues hypothesized that using a single PSD assembly but splitting the collected light into two separate wavelength bands would allow them to isolate the scintillator’s signal from the Cerenkov light (figure 6(c)). This problem can be formulated as a set of two linear equations with two unknowns, which in matrix form gives (Fontbonne et al 2002):

⎛⎝⎜

⎞⎠⎟

⎛⎝⎜

⎞⎠⎟

⎛⎝⎜

⎞⎠⎟∗ =

G BG B

K

KDD

1 1

2 2

g

b

1

2 (2)

The first matrix corresponds to a set of two measurement conditions (1) and (2) in two wave-length regions, for example, green and blue (G and B). kg and kb are calibration factors, and D1 and D2 are the expected doses for the given measurement conditions. Generally, the conditions

Figure 5. Measured charged produced by the Cerenkov light as a function of fiber-beam angle for 20 MeV electrons and 125 kVp x-rays (de Boer et al 1993).

Phys. Med. Biol. 61 (2016) R305

R315

Topical Review

are selected such that D1 = D2, but under very different expected Cerenkov light production scenarios: one maximizes Cerenkov light and one minimizes it. Once the calibration factors are known, the relation between scintillation light and dose, absent Cerenkov contamination, can be obtained under any irradiation conditions and for any photon or electron beam energy, at least for 60Co and above:

= ∗ + ∗D k G k Bg b (3)

Spectral bands other than blue and green can be chosen for this approach. Blue and green are usually selected because of the predominance of these wavelengths in the total light sig-nal from either blue or green scintillators and the availability of optical filters or photode-tectors with built-in color filters, such as CCD cameras and photodiodes (Fontbonne et al 2002, Archambault et al 2007, Lacroix et al 2008, Guillot et al 2011b, Therriault-Proulx et al

Figure 6. Schematic representation of the scintillation and Cerenkov signals in PSD (a) and associated optical removal techniques: the basic optical filtering (b), the chromatic removal (c), the multi-spectral removal (d) and the hyperspectral method (e). See text for more details.

Phys. Med. Biol. 61 (2016) R305

R316

Topical Review

2011b). Overall, this is an efficient subtraction method that is applicable to various multi-fiber array configurations. It also enables dose measurements in high-gradient regions such as the penumbra (Archambault et al 2006, Lacroix et al 2008), small fields (Gagnon et al 2012, Yoo et al 2012), and in vivo fields (Archambault et al 2010, Cartwright et al 2010).

By definition, the calibration condition corresponds to the boundary conditions in terms of Cerenkov light production. As such, as long as measurements are performed within these boundaries, the method has been shown to be quite efficient for both photon- and electron-beam characterization (Frelin et al 2005, Lacroix et al 2010b, 2011) (figure 7). These usually correspond to field sizes from 5 × 5 cm2 to 30 × 30 cm2. For small-field dosimetry, a different set of calibration conditions is usually needed (Morin et al 2013).

Figure 7. Deviation (%) between dose measurements made by a scintillator and that of an ion chamber for various field sizes and three Cerenkov removal techniques for BCF-12 scintillator (a) and BCF-60 scintillator (b). No Cerenkov subtraction is also displayed for comparison (no denoising). Note that 10 × 10 cm2 was used as the reference calibration field for no denoising, simple filtering and background subtraction. Figure reproduced with permission from Archambault et al (2006). Copyright 2006 AAPM.

Phys. Med. Biol. 61 (2016) R305

R317

Topical Review

3.1.4. Multi-/hyperspectral filtering. While the chromatic removal technique described above is efficient, it can only disentangle two component signals. If three or more signals need to be deconvolved, a different approach is needed. Archambault et al (2012c) recognized that the chromatic removal technique described in section 3.1.3 could be generalized to an arbi-trary number of signal sources N, measured in L wavelength bands. As long as L ⩾ N and at least L calibration measurements are performed, the resulting matrix system can be solved. Therriault-Proulx et al (2012) further made the distinction between a multispectral approach, where L ⩾ N using a limited number of large wavelength bands (figure 6(d)), and a true hyper-spectral approach using very narrow bands (figure 6(e)) where L � N. Although the details of these approaches are beyond the scope of this review, it is important to underline that the multi/hyperspectral method can extract scintillation signals in the presence of multiple stem effect contaminants (Cerenkov and fluorescence combined), allow for multiple scintillators to be attached to a single light guide or multipoint PSDs (mPSDs), and disentangle the temper-ature dependence of scintillation light, allowing both temperature and dose measurements in vivo (Therriault-Proulx et al 2015). Finally, Darafsheh et al (2015) recently used a simplified implementation of the hyperspectral method, which they called spectroscopic subtraction, for Cerenkov removal in a high-resolution PSD. Overall, the hyperspectral approach is a much more powerful and widely applicable approach than other stem effect removal techniques because it enables the deconvolution of an arbitrary number of signals and is limited primarily by the overall light collection efficiency.

3.1.5. Temporal filtering. Temporal filtering or timing was proposed by Clift et al (2002). In this technique, the light signal from a slow scintillation in between beam pulses is read. Since Cerenkov light production is a fast process and is correlated to beam pulse, one can use a slow-decaying scintillator whose signal will remain after the Cerenkov signal has decayed out. In this case, a BC-444G scintillator with 264 ns decay (compared to a few tens of nanoseconds for most plastic scintillators and scintillating fibers) was used. This approach has been com-pared to the two-fiber technique and shown to be highly efficient, but at the price of losing a sizable fraction (44%) of the scintillation light (Clift et al 2002). The need for timing electron-ics with links to the linac beam pulse signal adds complexity to this technique, but most mod-ern linacs now make transistor-transistor logic output available for treatment gating purposes.

3.1.6. Air-core fibers. Finally, another method bypasses the Cerenkov problem entirely by using air as the medium inside a hollow fiber—the so-called air-core fibers. With air’s refrac-tive index of 1, it is impossible to produce Cerenkov light in the range of clinical photon and electron beams from a standard linac. However, light propagation requires that the refractive index of the cladding be less than 1 for total internal reflection. Therefore, air-core fibers are generally coated with extremely thin layers of reflective dielectric materials or a high-Z metal. This approach was first proposed by Lambert et al (2008, 2010) and has been shown to be highly efficient. Generally, these fibers are very rigid and also exhibit light attenuation proper-ties that are up to 10 times higher than those of plastic core fibers, so only a few centimeters of air-core fibers can be used before they must be connected to regular optical guides. Because of the rigidity, their use for in vivo application, e.g. insertion inside catheters or applicators, tends to be avoided (Liu et al 2011).

3.2. Using Cerenkov light for dosimetry

The previous section  has reviewed techniques to remove Cerenkov light, treating it as a contamination signal or stem effect, but because Cerenkov light emission is, on its own,

Phys. Med. Biol. 61 (2016) R305

R318

Topical Review

proportional to dose, it may have applications in dosimetry. As such, Cerenkov light could be used for beam dose measurements (Glaser et  al 2014, Roussakis et  al 2015) and for in vivo surface measurements or ‘Cherenkoscopy’ (Zhang et al 2013, Holt et al 2014). This application has been explored as part of a clinical trial of in vivo dosimetry during external beam radiotherapy for patients with breast cancer (Jarvis et al 2014). One limitation of using Cerenkov light for dose measurement is that its intensity is orders of magnitude lower at proto n therapy beam energies than at photon and electron radiotherapy beam energies due to the lower secondary electron energies. One study, however, found that Cerenkov light production in a plastic optical fiber could potentially be used for depth-dose measurements in proton therapy (Jang et al 2012).

4. External beam applications

Before we turn to dose measurement applications for PSDs, it is important to recall that PSDs are water equivalent both in terms of photon energy absorption interaction cross sec-tions (μen/ρ) and electron stopping powers over a wide range of energies that encompasses all radiation therapy sources, from 192Ir all the way to the high-energy electron beams produced by modern linacs (Beaulieu et al 2013). As such, PSDs offer minimal perturbation of photon and electron fluence over the reference medium, i.e. water. A single calibration factor, either at 60Co or 6 MV, can therefore be used for all of the remaining beams in the energy range, from 192Ir and up, as stated previously (Beddar et al 1992c, Frelin et al 2005, Lacroix et al 2010b). This is a major advantage for beam characterization because the out-of-field energy fluence, composed mainly of scatter radiation, usually moves to a lower energy. For dosimeters whose response is energy dependent, a change in the calibration factor (or the introduction of a correction factor) relative to in-field measurements is required. PSDs are mostly immune to changes in energy as long as the majority of electron energy fluence remains above about 125 keV (Attix 1986). We refer the reader to section 5.3 for a more detailed discussion on the energy dependence of plastic scintillators.

Similarly, PSDs have been shown to be independent of dose rate at both low and high dose rates. More recently, both our group’s measurements of dose rates up to 2400 MU min−1 and published studies using 6 and 10 MV flattening filter-free beams confirm that this dose rate independence persisted even at the highest dose rate achievable on a radiosurgery sys-tem, while a specific Pion correction was needed for ion chambers (Beierholm et  al 2014, Underwood et al 2015, Beddar and Beaulieu 2016).

Finally, the most common PSD systems are built using 1 mm diameter scintillating fibers or plastic scintillators, though smaller diameters can also be readily found. Thus, the effec-tive measurement volumes of PSDs found in the literature range from 0.0023 cm3 down to 0.000 196 cm3, all much smaller than the commonly accepted definition of 0.01 cm3 for microchambers (Gagnon et al 2012, Ralston et al 2012, Morin et al 2013, Beierholm et al 2014).

While many detectors have one or two of the above-mentioned qualities, PSDs are among the few real-time dosimeters that possess all of them in one package. A new PSD for external beam dosimetry, the Exradin W1 (Standard Imaging, Middleton, WI, USA), is now com-mercially available. Several studies have looked into its stability, reproducibility, dose rate dependence, and other parameters, including uncertainty budgets (Beierholm et  al 2014, 2015, Carrasco et al 2015a, 2015b). For the most part, it has been found that this PSD exhibits the same characteristics as those described above.

Phys. Med. Biol. 61 (2016) R305

R319

Topical Review

4.1. Large-field photon and electron dosimetry

Standard treatment field parameters, such as depth-dose curves and beam profiles, were the first quantities measured by PSDs; the results were reported in the 1992 papers of Beddar et al (1992b, 1992c). PSDs were extremely accurate compared to a reference ion chamber for the depth-dose curves and provided better definition of the beam penumbra because they are smaller than ion chambers, as can be seen in figures 2 and 3. These results have been repro-duced many times by different groups (Frelin et al 2005, Archambault et al 2006, Lee et al 2007, Lacroix et al 2008).

To accurately calculate these parameters for electron beams, data from a reference ion chamber must be transformed from ionization in air to dose in water. This transformation involves a number of correction factors to account for the perturbation of the fluence by the presence of the air cavity, the chamber wall, and so on (Almond et al 1999). These factors are generally obtained from detailed Monte Carlo (MC) simulations or by using a perturba-tion-free detector. However, electron dose-to-water depth curves, including correct evalua-tion of the photon bremsstrahlung tail, are measured directly by PSDs (Beddar et al 1992c). Furthermore, because PSDs are insensitive to changes in beam quality, all electron beams can be measured with a standard calibration process that uses a 6 MV photon beam (figure 8). These results are consistent across many different PSD configurations. Lacroix et al (2010b) demonstrated using MC simulations that from a pure material composition perspective, the depth-dose curves calculated in water for a range of electron beams from 6 to 20 MeV were indistinguishable from those calculated in the PSD. Based on these results, the authors went on to use PSD-measured depth-dose curves in water to experimentally extract effective points of measurement for parallel plate chambers and depth-dependent perturbation factors for sili-con diodes (Lacroix et al 2011).

4.2. Small field dosimetry

Modern radiation therapy now delivers extremely targeted treatments that spare organs at risk and at the same time increase the dose to the target volume. Techniques such as intensity- modulated radiation therapy (IMRT), volumetric arc therapy (VMAT), stereotactic radio-surgery, and ablative radiation therapy provide exquisitely precise treatment delivery, often referred to as ‘dose painting’. These techniques superpose a large number of small fields and, therefore, raise new challenges for dosimetry. Most calibration and dosimetry standards were developed and defined for standard conditions, such as a 10 × 10 cm2 field. However, as field sizes drop significantly below 5 × 5 cm2, dose measurements become more problematic, and the true dose can be underestimated (Das et al 2008, Alfonso et al 2010).

The challenges faced in the measurements of small fields are (Laub and Wong 2003, Das et al 2008, Ding et al 2008, Alfonso et al 2010):

• source occlusion, which occurs when the field opening is such that the radiation source is partially shadowed;

• loss of charged particle equilibrium due to a charged particle range that is longer than the field size, preventing replacement along the measurement axis;

• a partial volume effect when the detector is too large; • the non-water equivalence of the dosimeter itself.

All of these phenomena contribute to incorrect dose measurements, with the largest dis-crepancies observed in the smallest fields (Sánchez-Doblado et al 2007, Klein et al 2010, Francescon et al 2012).

Phys. Med. Biol. 61 (2016) R305

R320

Topical Review

However, the characteristics of PSDs either make them immune to these deficiencies or mitigate them. MC studies by Francescon et al (2014b), Kamio and Bouchard (2014), and Papaconstadopoulos et  al (2014) showed that the closest agreement with a perfect detec-tor (represented by the MC code) was achieved by a PSD. Figure 9 presents the results of Francescon and colleagues’ study (Francescon et  al 2014b), which showed that the PSD (Exradin W1, in this case) had deviations from the MC-predicted doses of 0.3% for the small-est (5 mm) field. For that same field size, other dosimeters had deviations ranging from 7.4% to 35.0%. These simulation studies further showed that the only discernible disadvantage, from a theoretical perspective, of the Exradin W1 PSD was its overall size, which could account for 1.0–1.5% deviation due to volume averaging. Nonetheless, this was the smallest correction required by any of the small-field dosimeters studied in those works (Kamio and Bouchard 2014, Papaconstadopoulos et al 2014).

The first article reporting the experimental use of a PSD in small-field dosimetry was published by Létourneau et al (1999). They used a system very similar to the one shown in figure 4. In 2000, Westermark et al (2000) performed the first experimental comparison of multiple dosimeters, including a PSD, for ‘narrow’ photon beams. Beddar et al (2001) com-pared a miniature PSD system (1.6 × 10−3 cm3) to a diode and a to 0.1 cm3 PTW ion chamber (Freiburg, Germany). Some key papers have studied the output measurements of PSDs and compared these measurements to those obtained with other small-field dosimeters, such as diodes, EBT films, ion chambers, the MicroLion dosimeter and others (Ralston et al 2012, Morin et al 2013, Azangwe et al 2014, Francescon et al 2014a, Underwood et al 2015).

A new generation of PSDs that use advanced Cerenkov subtraction techniques (see sec-tion  3) has been designed specifically for small-field dosimetry (Archambault et  al 2007, Klein et al 2010, Lambert et al 2010, Gagnon et al 2012, Ralston et al 2012, Morin et al 2013). Interestingly, the experiments of Ralston et al (2012) and Morin et al (2013) have shown not only that PSDs need almost no correction, validating the theoretical studies discussed earlier,

Figure 8. Depth-dose curve of 6, 12 and 18 MeV electron beams from a Varian Linac measured by a NACP ion chambers and a PSD. Curves were normalized to dmax. Figure reproduced with permission from Lacroix et al (2010b). Copyright 2010 AAPM.

Phys. Med. Biol. 61 (2016) R305

R321

Topical Review

but also that PSDs can be used to experimentally extract correction factors for other dosime-ters. Morin and colleagues built a 0.000 196 cm3 PSD using 0.5 mm diameter plastic scintillat-ing optical fiber. Using reference MC data, they estimated that the volume averaging effect for the smallest (5 mm) CyberKnife field was less than 0.5% (Morin et al 2013). The relative total scatter factor of this PSD was compared to those of a 1 mm diameter PSD, three diodes, and a microLion ion chamber for all circular field sizes available on a CyberKnife unit. A com-parison of the MC and PSD data for the particular CyberKnife unit examined is presented in figure 10. For fields of 20 mm or larger, the measurements from all dosimeters were within the MC uncertainties. However, for the two smallest fields, significant substantial deviations were observed for some dosimeters. The 1 mm diameter PSD showed a 1.0–1.5% volume averag-ing effect, in line with the theoretical MC studies, while the 0.5 mm diameter PSD remained within the MC uncertainties for all fields without any corrections. The 0.5 mm diameter PSD was then used to experimentally extract detector- and treatment-unit-specific correction fac-tors, as defined by Alfonso et al (2008), for three diodes and a microLion dosimeter that were used to measure the output of 0.75 cm and 0.5 cm CyberKnife collimators (figure 10). The measured factors compared favorably to those generated by MC modeling and did not require accurate knowledge of the CyberKnife’s effective beam energy or any other complex MC modeling of the unit.

More recently, Francescon et al (2014a, 2014b) and Underwood et al (2015) have valid-ated the commercial Exradin W1 PSD for small-field dosimetry. Their results showed that the Exradin W1 accurately measured the depth-dose curves, tissue-maximum ratios, and off-axis ratios in the small-field conditions of a CyberKnife system. Furthermore, the on-axis output measurement of the PSD was almost perfect, and the correction factors were below 1% for all of the small fields; those for the microDiamond (PTW 60019) and Ediode (PTW 60017) detectors were as high as 6% and 8%, respectively (Underwood et al 2015). Similar

Figure 9. Reproduction of table 2 of Francescon, Kilby and Satariano Monte Carlo study of correction factors for various small-field dosimeters, including W1 scintillators (Francescon et al 2014b).

Phys. Med. Biol. 61 (2016) R305

R322

Topical Review

performance was shown for beam profile measurements. As in the studies discussed above, correction factors were extracted using the Exradin W1 PSD, and the experimental correction factors were found to be in good agreement with MC data. Underwood and colleagues, how-ever, underlined that the calibration process of the Exradin W1 PSD for small-field measure-ments may require further exploration.

In conclusion, the literature on PSDs indicates that PSDs, if made small enough to over-come volume averaging effect for the radiation beam under consideration, are considered as correctionless detectors in the constrained conditions of small-field dosimetry (Kamio and Bouchard 2014).

5. In vivo dosimetry with PSDs

5.1. External beam radiotherapy

Archambault et al (2010) demonstrated the feasibility of using PSDs for real-time in vivo dosimetry in both acrylic and anthropomorphic phantoms. They measured dose in 200-, 20-, and 2 cGy irradiations in 150 ms intervals, achieving accuracy of better than 1%. Subsequently, Klein et al (2012) used PSDs attached to endorectal balloons to measure dose delivered by IMRT and VMAT plans in both a deformable anthropomorphic phantom and a rigid IMRT quality assurance (QA) phantom. The measured dose was found to be within 1% of the dose calculated by the treatment planning system. In a recent review manuscript (Medical Physics Vision 20/20 series), Mijnheer et al (2013) discussed in details on the advantages of using PSDs for in vivo dosimetry in external beam radiation therapy.

Figure 10. Relative total scatter factors normalized to Monte Carlo as function of CyberKnife collimator size for various diodes (PTW 60008, PTW 60012, SFD Diode), a MicroLion chamber and two PSD (0.5 and 1.0 mm diameters). The dashed lines represent the Monte Carlo uncertainties of 1%. Figure reproduced with permission from Morin et al (2013). Copyright 2010 AAPM.

Phys. Med. Biol. 61 (2016) R305

R323

Topical Review

Wootton et al (2014) then used PSDs for in vivo monitoring of rectal wall dose in patients undergoing IMRT for prostate cancer. They enrolled five patients in an in vivo protocol and performed 142 total measurements during treatments. A daily computed tomography (CT) scan was acquired to determine the expected dose to the detectors, which was then compared to the measured dose. The average difference between measured and expected doses over all 142 measurements was −0.3%, with a standard deviation of 8.7%. The average difference for specific patients ranged from −3.3% to 3.2%, with standard deviations of between 5.5% and 7.0% for four of the five patients and 13.9% for the fifth (Wootton et al 2014).

This success has been followed by the development of the first commercial PSD specifi-cally designed for in vivo dosimetry. Radiadyne LLC (Houston, TX) has incorporated a pair of PSDs into their commercially available endorectal balloons for prostate immobilization (figure 11). Investigation of the utility of this new product is ongoing.

Finally, Therriault-Proulx et al (2012) have developed a novel mPSD that has promising applications for in vivo dosimetry. Traditional PSDs can only measure a single dose point per detector assembly. However, the mPSD can measure dose at multiple positions simultane-ously by using several different types of scintillating fibers in conjunction with a spectro-meter that distinguishes the spectra of the different fibers. For in vivo dosimetry purposes, it is desirable to maximize the amount of data acquired (that is, to measure the dose at as many points as possible) to ensure broad agreement between expected and delivered doses. There is a limit, however, to the number of traditional PSDs a patient can tolerate and that can be used without compromising clinical workflow. The mPSD solves these problems by maximizing the amount of data that can be acquired with the fewest detectors. A prototype was found to be highly accurate, achieving 2.3%, 1.6%, and 0.3% accuracy at three positions measured simultaneously (Therriault-Proulx et al 2013b).

5.2. Brachytherapy

One of the first scintillator dosimeter used in brachytherapy was the OPTIDOS system by PTW (Freiburg, Germany) and employ for calibration and dosimetry of beta emitters used for intravascular brachytherapy (Rosenthal et al 2003, Hakanen et al 2004). In fact, an early design of PSD for this type of application (Quast et  al 1998, Bambynek et  al 2000) was directly based on the two-fiber subtraction method presented in section 3. However, like the intravascular brachytherapy procedure, this particular scintillation dosimeter is not avail-able anymore. In the following section, we will therefore focused on current applications in brachytherapy, in particular for high-dose-rate (HDR) brachytherapy.

Figure 11. A commercial in vivo PSD for monitoring rectal wall dose during radiation therapy for prostate cancer. Two PSDs are embedded into the balloon and can be read out by connecting the optical connectors to a permanently installed system.

Phys. Med. Biol. 61 (2016) R305

R324

Topical Review

In HDR brachytherapy, a remote afterloader moves a high-activity radioactive source (typi-cally 20 000–40 000 U) inside catheters to precalculated dwell positions for associated dwell times. The high dose gradients around the source, combined with the hypofractionation of the dose and the multiple possible sources of error, make in vivo dosimetry a necessity during this type of treatment. Below is a nonexhaustive list of possible treatment errors that could be prevented by accurate in vivo dosimetry:

• Afterloader malfunction • Errors in treatment plan selection • Applicator reconstruction and fusion errors • Applicator length and source-indexer length errors • Incorrect source step size • Incorrect dwell times • Unintentional switching of guide tubes • Errors in Dose recordings • Errors in Replacement of the source between fractions (different activities, dwell times)

A good dosimeter for HDR brachytherapy has to possess specific characteristics: small size, good sensitivity and wide dynamic range, energy and angular independence, and the ability to be read in real time. In a Vision 20/20 article, Tanderup et  al (2013) conducted an in-depth review of the different types of detectors (thermoluminescent detectors (TLDs), metal–oxide–semiconductor field-effect transistors (MOSFETs), diodes, alanine detec-tors, radioluminescence detectors, and PSDs), evaluating each of them over a wide range of essential characteristics for in vivo HDR brachytherapy applications. PSDs showed very good performance across the board and, therefore, seem well suited for in vivo dose measure-ment during HDR brachytherapy. This had been experimentally demonstrated by Lambert et al (2007), who found that PSDs compared favorably to MOSFETs, diamond detectors, and TLDs for HDR brachytherapy dosimetry.

The first study on the use of PSDs in 192Ir HDR brachytherapy was performed in 1996 by Arnfield et al (1996). The dosimetry system was made of a 3 mm long × 1 mm diameter PVT plastic scintillator (BC-408S, Saint-Gobain Crystals, Cleveland, OH, USA) coupled to a 10 m silica fiber. Light was detected using a PMT. A Cerenkov stem effect removal approach consisting of simple spectral filtration was used, but only succeeded at subtracting part of the stem effect-induced light. Unfortunately, this detector was never used for in vivo applications.

A PSD-based dosimetry system was first used in a clinic for HDR brachytherapy for a 2011 study by Suchowerska et al (2011). This system, BrachyFOD, had first been described and tested in a phantom in 2006 (Lambert et al 2006) and was similar in design to the systems used by Beddar et al (1992b) and Létourneau et al (1999) for external beam dosimetry. A 4 mm long × 0.5 mm diameter PVT scintillator (BC-400, Saint-Gobain Crystals) was used, and the light was guided by a PMMA optical fiber to an OPTIDOS reader (PTW, Freiburg, Gernany), composed of a PMT and an electrometer. In this clinical study, the dose to the urethra was monitored in 24 patients undergoing brachytherapy for prostate cancer. A maximum dose dif-ference of 9% in the total integrated dose was reported between the measured values and the expected values from the treatment planning system. No stem effect removal approach was implemented in these studies, and no temperature dependence correction was reported by the authors.

Some groups have also studied plastic scintillators for verification of the spatial and tem-poral accuracy of brachytherapy treatment delivery. For example, Kojima et al (2009) used a CCD camera to detect the position and duration of the light emitted by a plastic scintil-lator block when an HDR brachytherapy afterloader positioned the radioactive source in a

Phys. Med. Biol. 61 (2016) R305

R325

Topical Review

phantom. This system was shown to be accurate, but its use is limited to pretreatment QA applications because of its design.

As mentioned above, Archambault et al (2010) used a CCD camera to measure the light output from an array of PSDs during in vivo external beam dosimetry. Cartwright et al (2010) used a similar system to simultaneously measure the light from 16 scintillation detectors placed around a rectal applicator, allowing them to determine the dose at the rectal wall. Each PSD was composed of a 4 mm long × 1 mm diameter plastic scintillator (BC-400) coupled to a 2 m long × 1 mm diameter polymer fiber. The dose measured with this system was, in gen-eral, within a 3% accuracy range of the treatment planning system, and the system allowed the source position to be determined within an average of 2 mm. The uncertainty in this dosimetry system was shown to be dependent on dose rate (0.2% at 1 Gy s−1 and 50% at 0.7 mGy s−1), and no Cerenkov removal method was implemented.

A silicon photodiode with an appropriate photo-amplification system can also be used as the photodetector in the PSDs for brachytherapy dosimetry. Lee et al (2010), in a phan-tom study, coupled a BCF-12 plastic scintillating fiber (10 mm long × 1 mm diameter) with a PMMA plastic optical fiber to guide the light to a photodiode system. Repeatability within 1.6% and a standard deviation of 0.94% were obtained for a source-to-detector distance of 5 mm. The PSD measurements differed from measurements taken with an EBT film by a max-imum of 1.94% and an average of 0.93%. The Cerenkov effect was shown to be important, but was measured separately with a 2-step process, which is not practical for an in vivo dosimeter.

In external beam radiation therapy, an effective way to remove the Cerenkov effect is the chromatic removal approach. In a study using an RGB photodiode connected to a dual-channel electrometer, Therriault-Proulx et al (2011a) showed that the chromatic removal approach also performed well for 192Ir HDR brachytherapy. The first system capable of complete stem effect removal for HDR brachytherapy was used for a clinically relevant in-phantom simulation of a typical prostate cancer treatment (Therriault-Proulx et al 2011b). The prostate and surround-ing organs at risk (bladder, urethra, rectum) were contoured on the CT images of an arbitrarily selected patient. A customized water phantom, including a catheter-positioning template, was used. Eleven catheters were inserted for dose delivery from the 192Ir source and two additional catheters were used to mimic dosimeters in the urethra and at the rectal wall. The PSD used in this study was composed of a 3 mm long × 1 mm diameter green scintillating fiber (BCF-60) coupled to a 7 m long × 1 mm diameter PMMA optical fiber. The integrated doses over the entire treatment plan and for each catheter were measured and compared to the expected values from the treatment planning system and to the dose rates for the individual dwell positions.

Another interesting avenue for in vivo dosimetry during HDR brachytherapy is the use of single-fiber mPSDs. mPSDs are able to measure the dose simultaneously at many positions with a single detector insertion and are particularly useful when space is limited (e.g. in cath-eters or orifices). Therriault-Proulx et al (2013b) demonstrated the use of a three-point mPSD for 192Ir HDR brachytherapy with the customized water phantom described earlier in this section. They employed a hyperspectral filtering approach using a spectrometer to assess the contrib ution of each light-emitting element (BCF-10, BCF-12, BCF-60, and stem effects) to the total light output (Archambault et al 2012c, Therriault-Proulx et al 2012, 2013b). Average relative differences between the measured and expected values of 3.4% (±2.1%), 3.0% (±0.7%), and 4.5% (±1.0%) were obtained for source positions going radially from 1 cm to 5 cm and over a longitudinal range of 10.5 cm. The mPSD was also shown to be very effective in measuring the position of the radioactive source because it employs the knowledge of the relative positions of the multiple scintillating elements. Given the promising results obtained with the mPSD in that study, such a dosimeter could be for clinical studies.

Phys. Med. Biol. 61 (2016) R305

R326

Topical Review

5.3. Superficial therapy, radiology, and interventional radiology

Obtaining accurate dosimetry in the kilovoltage energy range used in radiology is compli-cated by the photoelectric effect, which increases as the energy decreases (E−3) and strongly depends on the composition of the dosimeter (~Zn, with n = 3–4). While primary calibration standards do exist for kilovoltage units, most interventional radiology units use air kerma at a fixed point in space to report dose. This quantity is, however, highly inaccurate for char-acterizing entrance skin dose to dose to organs, with differences of 35% or more from the actual delivered dose. Most dosimeters are also energy dependent in this energy range due to their composition. This includes PSDs, as demonstrated in figure 2 of Lessard et al (2012). Furthermore, tissues themselves begin to show interaction cross section differences for photon energies below 100 keV relative to water, air, or other tissue until an average energy of about 30 keV is reached—approximately the lower end of a CT scanner’s (80–90 kVp) or fluoros-copy unit’s energy (e.g. figure 1 of Beaulieu et al (2012)).

The second challenge for PSDs in radiologic applications involves quenching, which is related to the possible non-linear relationship between light production and deposited energy as the secondary electron energy fluence shifts toward lower energy and, thus, higher LET. An increase in alternative energy dissipation, or quenching of light production, has been docu-mented for plastic scintillators at low energies (Williamson et al 1999, Frelin et al 2008a). Frelin et al (2008a) and Williamson et al (1999) demonstrated that the energy at which sig-nificant quenching begins to occur depends strongly on the scintillator’s chemical composi-tion, starting at 40 keV and reaching to 160 keV for electrons, and 55 keV and 410 keV for photons. A systematic study by Lessard et al (2012) using plastic scintillating fibers instead of plastic scintillators (which differ by the bulk medium) showed that the effect of the photon interaction cross-section (μen/ρ) was a more important correction factor, leaving a ±5% effect that could be accounted for by quenching (relative to the MC dose prediction) in the range of 80–150 kVp (effective energy of 26–86 keV for the specific unit used). They concluded that the level of accuracy was sufficient for use in real-time in vivo dosimetry and that quenching could be ignored. It is important to note that the PSD in above-cited study by Lessard et al was calibrated using a reference ion chamber in the radiology energy range. An overall large sys-tematic shift (quenching effect) might still exist, but such an effect could only be extracted if the PSD were initially calibrated in an energy range where quenching is negligible (e.g. a 60Co or 6 MV beam) and then used at radiology energy (Boivin et al 2016). In order to compensate for this effect, Williamson et al has suggested incorporating high-Z elements, such as chlo-rine, within the plastic scintillator matrix (Williamson et al 1999). More recently, Nowotny and Taubeck proposed to incorporate inorganic scintillators (in powder form) to polystyrene plastic scintillator in order to ensure a flat energy response for x-ray application but could not fully correct for light quenching effect (Nowotny and Taubeck 2009). Finally, the lower secondary electron energy found in radiologic applications does not produce Cerenkov radia-tion. Fluorescence stem effects remain a possibility but are generally considered negligible (Marckmann et al 2006, Lessard et al 2012, Therriault-Proulx et al 2013a).

Jones and Hintenlang published one of the first applications of a PSD to radiology (Jones and Hintenlang 2008). This was followed by a more in-depth study by Hyer et al (2009). The latter showed energy dependence attributable to the photoelectric effect of almost 40% between 40 kVp and 120 kVp. That study further showed that in a measurement of a depth-dose curve, the PSD exhibited an energy dependence of about 10% attributable to beam hard-ening. Lessard et al (2012) later developed an effective correction, thereby eliminating most of this dependence (figure 12). However, for each individual energy, Hyer et al (2009) demon-strated that the light output was linearly related to dose.

Phys. Med. Biol. 61 (2016) R305

R327

Topical Review

Hyer et al (2010) subsequently used the PSD system they had previously developed to measure the entrance dose and the dose to various organ locations within an anthropomorphic phantom. Similarly, Yoo et al (2014) used a 1 mm diameter × 10 mm long BCF-12 scintil-lating fiber to measure entrance surface dose on an anthropomorphic thorax phantom. One interesting result of the latter study was that the PSD was transparent to radiological beams and thus would not interfere with diagnostic or interventional procedures. The same group also developed a second system that eliminated potential fluorescence contamination by using the subtraction method of Beddar et al (1992b) and Yoo et al (2013).

More recently, Boivin et  al (2015b) presented a systematic comparison of the perfor-mance of PSD systems that used different photodetectors exposed to dose rates ranging from 0.07– 300 mGy s−1 and beam energies from 120 kVp to 23 MV. The data suggested that PSDs would have applications in interventional radiology, diagnostic radiology, and radiotherapy, and the authors provided guidelines for selecting the appropriate PSD system components. The same group then developed and optimized a PSD system for real-time dosimetry during diagnostic and interventional procedures (Boivin et al 2015a). However, at this point, no PSD system has been used for in vivo patient measurements in the field of diagnostic radiology.

6. 1D and 2D plastic scintillation dosimetry

The applications presented above are generally limited to a single, small (1–10 mm long) scintillation sensor connected to a light guide. Because the basic interaction physics and other properties of plastic scintillators make them indistinguishable from water for most photon and electron beams found in clinical practice, PSDs are among the few real-time dosimeters that can be packed closely together. Archambault et al (2007) nicely demonstrated this by using a large block of plastic scintillators immersed in water. Therefore, an obvious extension would be to use multiple single-scintillator probes to form either a one-dimensional (1D) or two-dimensional (2D) measurement array.

The first discussion of such a PSD array used for dosimetry can be found in a 1996 paper by Flühs et al (1996), which dealt with ophthalmic plaque dosimetry. An array of 16 PSDs was used, and the two-fiber subtraction method was implemented to account for stem effects. The PSDs were mounted on a cylindrical piston that could be moved up and down and rotated about its long axis. By combining these two degrees of liberty, this setup enabled a full angu-lar characterization of the ophthalmic plaque on a plane (2D) and in three dimensions (3D). The authors further automated the translation and rotation process via a computer-controlled interface and demonstrated 3D 125I eye plaque dosimetry comprised of 1000 measurement points in about 10 min (Bambynek et al 1999). These concepts have been replicated, with some modifications over the last 10 years (Sliski et al 2006, Eichmann et al 2009). Finally a larger array consisting of 40 PSDs (with 40 subtraction fibers) was built for use in external beam dosimetry measurements (Becks et al 2000). Signals from all 80 collecting fibers were read by a CCD photodetector after passing through an image intensifier. Depth-dose curves were measured with the array. However, depending on the measurement conditions, the results were shown to differ significantly from those obtained with an ion chamber. The differences were attributed to the handling of the Cerenkov light, dose gradient, among other factors.

To account for Cerenkov and other stem effects in 3D applications, Archambault et al (2007) published the first experiments with a PSD array that applied the chromatic removal technique using a color CCD photodetector. They were able to show a good agreement between PSD and ion chamber measurements for the depth-dose curve, especially in high-dose-gradient regions; the key gain came from the use of a single fiber for both the dose signal and Cerenkov correction signal. Bartesaghi et al (2007) also used the two-fiber subtraction method with an

Phys. Med. Biol. 61 (2016) R305

R328

Topical Review

array of eight PSDs alongside eight clear fibers, connected to a multi-anode PMT, to measure the depth-dose curves of a 6 MeV electron beam and a 6 MV photon beam within 2%. This concept was extended by Lacroix et al (2008), who inserted a linear configuration of 29 PSDs in a solid water phantom (30 × 30 cm2 plaque). A robust pin-and-dowel positioning system was employed to bring the collecting optical fibers to a color CCD photodetector for chromatic removal. They successfully acquired a 6 MV depth-dose curve (with a maximum deviation of 1.6%) and simultaneous profile measurements for 4 × 4 cm2, 10 × 10 cm2, and 20 × 20 cm2 fields. More recently, Liu et al (2012) proposed a smaller-scale, 16-PSD linear array. The main design difference lies in the use of a multi-anode PMT array for light detection.

Lee et al (2008) designed a 5 × 5 PSD array in which the PSDs were spaced 5 or 10 mm center-to-center. They used 10 mm long scintillators. Depth-dose curves were extracted for 6 MV and 15 MV photon beams. That study did not consider Cerenkov subtraction. Guillot et al (2011a), however, took the need for Cerenkov subtraction into account when they built a very large, 26 × 26 cm2 2D array composed of 781 PSDs equally spaced 10 mm apart, except on the central X and Y lines, where 5 mm spacing was used to extract more accurate beam pro-files. To apply the chromatic removal method for such a large number of PSD signals, Guillot and colleagues used two CCD photodetectors in combination with a dichroic beam splitter and two optical filters. The detector was set on the treatment table and could be irradiated from any angle. Treatment plans were applied on a CT model of the detector. The study demonstrated the water-equivalence of this large fiber array in terms of CT-number (relative to water) and through the acquisition of standard and IMRT beams using the true delivery angles. Guillot et al (2013) also conducted a performance analysis using signal detection theory and extrac-tion of detector-specific receiver operating characteristic curves for various scenarios.

Figure 12. (Top) Depth-dose curve of a 100 kVp orthovoltage beam (44 keV effective energy) comparing PSD, corrected PSD, Graftchromic film, Monte Carlo and literature. (Bottom) relative difference to Monte Carlo of the various data sets. Figure reproduced with permission from Lessard et al (2012). Copyright 2012 AAPM.

Phys. Med. Biol. 61 (2016) R305

R329

Topical Review

An interesting inverted-T PSD array design was proposed by Yoo et al (2012). It enabled simultaneous measurements of a beam profile along one axis (in-plane or cross-plane) and a depth-dose curve. Fifty square fibers were read by a color complementary metal–oxide– semiconductor CCD photodetector. The array was applied to 1 × 1 cm2 and 3 × 3 cm2 small-field measurements. Gagnon et al (2012) proposed a cross-hair array composed of 49 PSDs (also read by a color CCD photodetector) that could be dedicated to radiosurgery measure-ments. Output factors and profiles for cones from 4 to 40 mm were measured and compared to measurements taken with diodes (shielded and unshielded), a micro-ion chamber (Exradin A16), and EBT films. Results confirmed that this PSD array performed on par with EBT films for profile measurements.

The results presented by Guillot et al (2011a) demonstrated the limitations of using a single PSD probe to build larger arrays. Goulet et al (2010, 2011) explored a completely different approach to this problem: using a collection of long scintillating fibers. They suggested using a long fiber as a 1D (line dose) integrator and developed a mathematical formalism for that purpose. One potential application demonstrated by that group was a real-time fluence moni-tor (transmission monitor) that could be inserted directly in a linac treatment head accessory tray, similar in concept to the device for advanced verification of IMRT deliveries (DAVID (Poppe et al 2006)) and other such detectors (figure 13). By having one long fiber associated with each leaf pair of a multileaf collimator-equipped linac and read at both ends simultane-ously, Goulet et al (2011) were able to identify not only small changes in overall fluence, errors in field opening, and systematic shifts in collimator positions, but also errors as small as 1 mm in a single leaf positioning. They determined the beam transmission factor to be 0.983, which is smaller than that of any other devices available at that time (see table 4 of Goulet et al (2011)). Izaguirre et al (2012) and Knewtson et al (2015) from Washington University have since reproduced this experiment using a different photodetector apparatus but using the same principle and reached similar results.

The concept multiple long fibers was pushed further by applying the planar configuration proposed above on the treatment table and performing multiple measurements by rotating the devices on the central axis around the perpendicular plane relative to the beam. This acquisi-tion scheme was coupled to a tomographic reconstruction algorithm (much like CT image reconstruction) to obtain a high-resolution dose pattern of 1 × 1 mm2 at all points in the meas-urement plane. This technique, called tomodosimetry, was demonstrated experimentally by Goulet et al (2012). A similar system using a configuration of a few long scintillating fibers distributed on the surface of two concentric cylindrical planes was also proposed by the same group. Numerical simulations showed that this system had potential for use in high-resolution (1 × 1 × 1 mm3 voxels) 3D volumetric dosimetry (Goulet et al 2013a, 2013b). However, the necessity for rotation in both the 2D and 3D designs makes clinical translation of this approach difficult. Other 3D geometries (discussed in the next section) are more attractive at this time.

A last alternative for 2D dosimetry is the use of scintillator sheets. This concept was pro-posed as early as 1959 by Olde and Brannen (1959), who used a PMT and a focusing lens. However, the first complete demonstration of such system was conducted by Ge, in his 1988 master’s thesis at Washington University in St. Louis (Ge 1988). Ge’s system was com-posed of a scintillator sheet immersed in a water tank and read by a optical camera. A similar approach was used a few years later by Perera et al (1992), also from Washington University, to measure the 2D dose distribution of a brachytherapy source.

This basic configuration has been slightly modified to position the scintillating sheet perpend icularly to the beam for both proton (Boon et al 1998) and photon beam measure-ments (Petric et al 2006). In that case, a mirror is used to allow a CCD camera to read the light at 90° from the beam incidence. Petric et al (2006) measured beam profiles with a maximum

Phys. Med. Biol. 61 (2016) R305

R330

Topical Review

deviation of 8% at the beam edge and identified a few key issues, such as optical scattering and glaring effects, that required corrections. Frelin et al (2008b) introduced an innovative deconvolution technique that, among other things, took care of the Cerenkov light contamina-tion that is produced in either water or clear plastic phantoms. This system, called DosiMap, was tested against standard fields. Depth-dose curves and beam profiles agreed with an ion chamber to a maximum deviation of 4%. The system was further tested for electron beams and IMRT fields (Collomb-Patton et al 2008, 2009).

7. 3D volumetric scintillation dosimetry

Volumetric scintillation is a promising new area of study that aims to make fast, high-resolu-tion, and accurate measurements of absorbed dose distributions by imaging the light emitted by a scintillating medium in 3D. One of the attractive features of this methodology is that, unlike other detectors, the medium/phantom used for the measurement is the detector itself, which can be water-equivalent if chosen judiciously and is read in real-time.

7.1. External beams

The development of 2D and interpolation 3D dosimetry (from 2D measurements) has been a very active area for photon radiotherapy treatment QA systems, as evidenced by the avail-ability of many such systems, such as the ArcCHECK, ScandiDos Delta 4, MatriXX, and

Figure 13. A scintillating fiber-based fluence (transmission) monitor. The system shown is read by a CCD photodetector (A) and mount in the accessory tray (B). The configuration is set such that each leaf pair (here for a Varian MLC collimator) is covered by one fiber (C). Figure reproduced with permission from Goulet et al (2011). Copyright 2011 AAPM.

Phys. Med. Biol. 61 (2016) R305

R331

Topical Review

MapCheck. However, since volumetric scintillation is a promising new method for fast, accu-rate, and high-resolution measurement in proton therapy, a few studies have examined its potential for QA of external photon beams. The most important studies in this area were undertaken by Pönisch et al (2009), who used a volume of LS, and Kroll et al (2013) and Goulet et al (2014), who used a volume of solid plastic scintillator.

Pönisch et al (2009), using an extension of the system that was described by Beddar et al (2009), investigated the LS system for application in real-time high-energy photon beam dosimetry. Figure 14 illustrates how that study verified the dose distribution of a simple four-field-box technique. Excellent agreement was found between the corrected scintillation light and the dose distribution calculated with the treatment planning system throughout both the high- and low-dose regions. The modified detector system consisted of a light-shielded, LS-filled acrylic tank (outer dimensions: 17.8 × 14.0 × 12.7 cm3, wall thickness: 3 mm), a camera objective, and a high-sensitivity electron-multiplying CCD camera (Luca EM, Andor Technology, South Windsor, CT). The light data measured with the CCD camera were filtered to correct for scattering of the optical light inside the LS. The depth-dose and lateral profiles as well as the 2D dose distributions were found to agree with the results from the treatment plan-ning system. The authors concluded that the LS system had several valuable characteristics, including near water equivalence, dose rate independence, linearity with dose, and robust-ness for single or multiple acquisitions. They also concluded that the use of this system for dosimetry QA checks could save a great deal of time over both film and scanning ion chamber measurements.

While the previous system did use a volumetric detector, dose reconstruction was limited to 2D. Kroll et al (2013) tried to introduce a 3D tomographic reconstruction method based on a right prism with a regular hexagonal base solid scintillator block read by four CCD cameras: three coplanar CCDs spaced 120° apart perpendicularly to the beam direction and one rear CCD. With this experimental setup, they acquired depth-dose curves for electron and proton beams. Bremsstrahlung photons made 3D tomographic dose reconstruction using the rear CCD, which was placed in the axis of the beam, difficult. Furthermore, proton beam measure-ments were plagued by a strong quenching effect, similar to that described above. However, that system was able to detect dose patterns down to 50 mGy, and its scintillator size allowed measurements of electrons of 19 MeV or less and protons of 110 MeV or less. The geometry used in this study mandated that all beams be delivered at a fixed angle, which would be a limitation for modern photon external beam radiotherapy.

This limitation was later overcome by Goulet et al (2014), who introduced a measurement geometry allowing real-time monitoring of treatment delivery for IMRT, VMAT, and stereo-tactic body radiotherapy. The system they proposed is shown in figure 15. They developed and validated a novel type of high-resolution 3D dosimeter based on the real-time light acquisition of a plastic scintillator volume using a plenoptic camera and tomographic reconstruction algo-rithms. With this system, they were able to perform millimeter-resolution, water-equivalent dosimetry of an IMRT and a VMAT plan over a whole 3D volume. Their results suggested that using plenoptic camera technology will enhance volumetric scintillation dosimetry and the full 3D reconstruction of dose distribution in real time. The set-up shown in figure 15 was able to achieve 2 × 2 × 2 mm3 spatial resolution using a basic plenoptic camera and the linac portal imaging device to constrain the tomographic reconstruction in the transverse plane.

7.1.1. Proton beams. In 2006, Fukushima et  al (2006) reported on the development of an easy-to-handle range measurement tool using a rectangular (5 × 5 × 40 cm3) block of organic plastic scintillator (BIRCON, BC-400). They recorded the scintillation light using a high-definition video camera (Sony, HDR-HC1) as the CCD camera (1920 × 1080 pixels, at 30

Phys. Med. Biol. 61 (2016) R305

R332

Topical Review

Figure 14. Experimental validation of the LS system for the QA of photon RT: (a) CT simulation, (b) treatment planning, (c) treatment delivery, (d) and image acquisition/processing of the light distribution produced within the LS detector compared to the expected dose distribution (Beddar 2015).

Figure 15. (Left): A plastic block scintillator system using a plenoptic camera in conjunction with an EPID for 3D photon QA. (Right) Axial (a), sagittal (b), and coronal (c) planes from a 3D reconstruction of an IMRT brain treatment measured with the plenoptic camera and the portal imaging device. The 2D gamma maps for the criteria (3% dose, 3 mm distance) are also shown. Figure reproduced with permission from Goulet et al (2014). Copyright 2014 AAPM.

Phys. Med. Biol. 61 (2016) R305

R333

Topical Review

frames per second (fps), recording images in 8-bit grayscale). Using this detector in 2D mode, Fukushima and colleagues were able to measure range by measuring the depth of the bright-ness distribution and to observe the time structure of the range of the proton beam. They con-cluded that their tool was superior to the conventional method, which used an ion chamber and a water phantom, in terms of setup time, irradiation time, and range variation measurement. They also concluded that this tool would be useful for proton range measurement and control in the clinical setting.

In 2009, Beddar et al (2009) were the first team to study in detail the feasibility of using a 3D liquid scintillator (LS) detector system for the verification and full characterization of proton beams in real time for proton therapy applications. Their first LS detector system con-sisted of a light–tight gray polyvinyl chloride (PVC) phantom containing a smaller transparent acrylic inner tank filled with a LS at one end. At the other end of the tank was a CCD cam-era, as illustrated in figures 16(a) and (c). The scintillation light produced within the 3D liquid volume during irradiation with pristine proton Bragg peaks was measured at acquisition rates of 20 and 10 fps, thereby providing a series of consecutive image frame sequences. The depth scintillation light profiles from single images with acquisition times of 50 and 100 ms were identical. These measurements were then compared to ion chamber measurements and MC simulations. The light distribution measured from the images acquired at rates of 20 and 10 fps had standard deviations of 1.1% and 0.7%, respectively, in the plateau region of the Bragg curve. Beddar et al (2009) have also demonstrated the feasibility of using a 3D volumetric scin-tillator to measure the range, intensity, position, and dose distribution of a proton beam in 2D (as depicted in figure 17) and verified that the system was capable of acquiring multiple images within a single proton pulse with good signal-to-noise ratio and submillimeter image resolution. Although the size of this first LS detector was limited (7 × 7 × 14 cm3), they were able to dem-onstrate that (1) a larger detector volume could be used, (2) the method could be extended to 3D by increasing the number of cameras viewing the LS volume from multiple angles (figure 16(b)), and (3) the method could be used for the QA of intensity modulated proton therapy beams.

Members of the same team continued to develop this line of research, with Archambault et al (2012a) developing a larger LS detector system (20 × 20 × 20 cm3) to further character-ize proton beams and Robertson et al (2013) further exploring quenching effects in LSs. Both studies demonstrated that this method was not only feasible, but also had great potential for performing accurate, high-spatial-resolution, and nearly real-time volumetric dosimetry in proton therapy.

Archambault et al (2012a), using a larger LS detector viewed by a CCD camera with a field view of ~ 26 × 20 cm2 and a pixel size of 0.4 mm, measured the spatial location, inten-sity, and range (energy) of proton beamlets in near real-time. Pencil beams with nominal ranges in water of 9.5–17.6 cm were scanned to irradiate an area of 10 × 10 cm2. Images were acquired at 50 ms per frame. The proton ranges measured from light distributions produced in the LS were accurate to within an average of 0.3 mm. The largest reported deviation observed between nominal and measured ranges was 0.6 mm; the laterally measured pencil beam posi-tion was accurate to within 0.4 mm, on average. The intensity of single proton spots was measured to a precision of 3% for the smallest spot intensity (0.005 MU) and 0.5% for the largest spot (0.04 MU).

Robertson et al (2013) undertook the challenging task of studying the problem of quench-ing exhibited by scintillating materials, with the goal of accurately converting the light signals to absorbed dose. They aimed to develop an ionization quenching correction method to restore the linear dose response of scintillators irradiated by proton beams. The quenching correction method they used was based on an empirical model developed by Birks that predicts quench-ing on the basis of the LET of the proton beam (Birks 1951, 1964). This method required prior

Phys. Med. Biol. 61 (2016) R305

R334

Topical Review

knowledge of the LET distribution of the beam and calculation of the Birks model parameters for the LS that was used (BC-531). They reported that the calculated and measured Bragg peak heights agreed within 3% for all energies except 85.6 MeV, where the agreement was within 10%. The quality of the quenching correction was poorer for sharp low-energy Bragg peaks because of blurring and detector-size effects. Robertson and colleagues concluded that future detectors should improve blurring correction methods and optimize pixel size to improve accuracy for low-energy Bragg peaks.

Figure 16. (a) Schematic of the LS detector system with one CCD camera (b) the configuration with two CCD cameras to resolve the 3rd dimension and (c) the actual LS detector system illustrated in (a) being set-up on a robotic couch at the scanning beam gantry at the Proton Therapy Center of the University of Texas MD Anderson Cancer Center in Houston TX (reproduced from: Beddar (2015)).

Figure 17. A consecutive series of images of the LS detector system exposed to a 120 MeV proton beam. Each image represents the total light output seen within a 100 ms time interval. Figure reproduced with permission from Beddar et  al (2009). Copyright 2009 AAPM.

Phys. Med. Biol. 61 (2016) R305

R335

Topical Review

Robertson et al (2014) performed another study to theoretically and experimentally charac-terize and correct for the optical artefacts that affect the accuracy of 3D LS detector systems, including photon scattering, refraction, camera perspective, vignetting, lens distortion, the lens point spread function, stray radiation, and noise in the camera. The largest optical arte-facts in these detector systems are blurring from the lens point spread function and refraction at the tank window-air interface. Photon scatter in the scintillator was not found to be a signifi-cant source of artefacts. The correction methods used by Robertson and colleagues effectively mitigated the impact of the artefacts, increasing the average gamma analysis pass rate from 66% to 98% for gamma criteria of 2% dose difference and 2 mm distance to agreement. They also concluded that the optical artefact correction methods described in their study could be directly applied to future systems that incorporate multiple cameras to facilitate a full 3D reconstruction of the light signal.

Another successful simulation study was conducted by Hui et  al (2014), who assessed the feasibility of reconstructing 3D scintillation light distributions from proton pencil beams using limited viewing angles. Their proposed system consisted of a tank filled with LS imaged by three CCD cameras at three orthogonal viewing angles. Because of the limited number of viewing angles, Hui and colleagues developed a profile-based technique to obtain an ini-tial estimate that could improve the quality of the 3D reconstruction. They found that their

Figure 18. 3D scintillation dosimetry using a liquid scintillator: (top) photograph of the set-up and (bottom) schematic representation of the dosimetry system (Kirov et al 2005).

Phys. Med. Biol. 61 (2016) R305

R336

Topical Review

profile-based technique was able to reconstruct a single-energy proton beam (i.e. a 161.6 MeV proton beam) in 3D with a gamma pass rate (3%/3 mm local) of 100.0%. For a single-energy layer of an intensity-modulated proton therapy prostate treatment plan (i.e. a collection of 69 161.6 MeV proton beams adopted from an actual proton treatment plan for prostate can-cer), the proposed method was able to reconstruct the 3D light distribution with a gamma pass rate (3%/3 mm local) of 99.7%. In addition, they found that the proposed method was effec-tive in detecting errors in the treatment plan. They concluded, as have the other authors whose work is discussed here, that LS can be a very useful tool for 3D proton beam QA and has the potential to enable efficient, high-resolution 3D QA of clinical proton beams.

In a subsequent paper, Hui et al (2015) refined the Beddar group’s method by developing a geometric calibration system to accurately calculate physical distances within the LS detec-tor, taking into account optical artefacts. They also assessed the accuracy, consistency, and robustness of proton-beam range measurements taken with LS detectors and this new geomet-ric calibration system. Measurements were made on 3 different days to evaluate the setup’s day-to-day robustness, and three sets of measurements were made on each day to evaluate its consistency from delivery to delivery. All proton beam ranges measured using the LS sys-tem were within half a millimeter of the nominal range. The delivery-to-delivery standard deviation of the range measurement was ± 0.04 mm, and the day-to-day standard deviation was ± 0.10 mm. In addition to being accurate and robust, the LS detector was able to measure the range of 94 proton beams (72.5–221.8 MeV) in just two deliveries. (Because a maximum of 60 different energies could be used during each delivery from the synchrotron, two separate deliveries were used.) These qualities make such a detector a perfect tool for range measure-ment of spot-scanning proton beams.

7.2. Brachytherapy

In the previous section, we described efforts to obtain measurement points around ocular brachytherapy sources (Flühs et al 1996, Bambynek et al 1999). Those papers introduced the notion of 3D measurement using translation and rotation. While not itself introducing a form of volumetric scintillation dosimetry, this work set the stage for extended spatial measure-ments. Kirov et al (2005) combined multiple planar 2D dose pattern acquisition using a CCD camera with a lens focusing on a rotation stage and a tomographic algorithm to reconstruct 3D dose patterns around a 106Ru eye plaque. The experimental setup is presented in figure 18. While the acquisition process was fast, the overall precision of the system was rather poor. That paper provided an in-depth discussion of the discrepancies observed, including scintilla-tion quenching, the limitations of the algorithm, and other phenomena.

8. Conclusion and future perspectives

There has been a great deal of interest in using organic scintillation materials for dosimetry in photon, electron, and proton therapy, mainly because they have favorable dosimetric char-acteristics and advantages over other commonly used materials, dosimeters, and detectors. During the last decade or so, our field has witnessed an explosion of innovation in these new 1D, 2D, and 3D dosimetry methods. A growing body of research shows that solid plastic and LSs can be applied to or used in:

• small-field, radiosurgery, and CyberKnife dosimetry • real-time in vivo dosimetry for external beam radiation therapy and brachytherapy • multiple dose points PSD (mPSDs) using a hyperspectral approach

Phys. Med. Biol. 61 (2016) R305

R337

Topical Review

• radiologic imaging and interventional radiology • volumetric dosimetry for proton, electron, photon, and brachytherapy applications

Further investigations, innovations, and developments will improve the status of scintilla-tion dosimetry as a method for laboratory research or for additional clinical applications, espe-cially in light of the limitations of more commonly used methods. Significant research and development efforts are especially warranted on the hyperspectral approach using multiple PSDs within a single optical fiber line and the volumetric approach using 3D liquid or solid scintillation detectors. Indeed, the latter approach is opening the window for a new paradigm in which the dose distribution would be imaged in near real-time instead of being measured point-by-point or with planar detection. Quenching is another challenging area that requires further investigation to extend scintillation dosimetry to light and heavy ion therapies other than proton therapy.

Acknowledgments

The authors would like to acknowledge the National Cancer Institute (NCI) of the National Institutes of Health (NIH) of the United States and the Natural Sciences and Engineering Research Council (NSERC) of the Canadian government, who helped support not only the research but also the students and postdoctoral fellows who have contributed to scintillation dosimetry. The views and content expressed are solely the responsibility of the authors and does not necessarily represent the official views of the NIH or NSERC.

We would like to thank Amy Ninetto of the Department of Scientific Publications at MD Anderson Cancer Center for her help editing this paper. We also would like to acknowledge all of our students and postdoctoral fellows and thank them for their hard work, dedication, and contributions to the field of scintillation dosimetry. Their work is being referenced in this very first review paper on the subject. We are also grateful to our colleagues, who were very sup-portive of this research and who challenged us to advance this field to the point where it stands now. It is here to stay. Finally, we must acknowledge and thank all the authors worldwide who have collaborated with us as well as those who have contributed in their own initiatives to advance the field of scintillation dosimetry.

References

Alfonso R et al 2008 A new formalism for reference dosimetry of small and nonstandard fields Med. Phys. 35 5179–86

Alfonso R et al 2010 Status of IAEA/AAPM Recommendations on small and composite field dosimet Med. Phys. 37 3096

Almond P R, Biggs P J, Coursey B M, Hanson W F, Huq M S, Nath R and Rogers D W 1999 AAPM’s TG-51 protocol for clinical reference dosimetry of high-energy photon and electron beams Med. Phys. 26 1847–70

Andersen C 2011 Fiber-coupled Luminescence dosimetry in therapeutic and diagnostic radiology AIP Conf. Proc. 1345 100

Archambault L, Arsenault J, Gingras L, Beddar A S, Roy R and Beaulieu L 2005 Plastic scintillation dosimetry: optimal selection of scintillating fibers and scintillators Med. Phys. 32 2271–8

Archambault  L, Beaulieu  L and Beddar  S  A 2012b Comment on ‘Plastic scintillation dosimetry: comparison of three solutions for the Cerenkov challenge’ Phys. Med. Biol. 57 3661–5

Archambault L, Beaulieu L, Klein D, Bourguoin A, Wootton L, Therriault-Proulx F and Beddar S 2013 Performance of the first commercial miniature plastic scintillation detector Radiother. Oncol. 106 S26

Phys. Med. Biol. 61 (2016) R305

R338

Topical Review

Archambault  L, Beddar  A  S, Gingras  L, Lacroix  F, Roy  R and Beaulieu  L 2007 Water-equivalent dosimeter array for small-field external beam radiotherapy Med. Phys. 34 1583–92

Archambault  L, Beddar  A  S, Gingras  L, Roy  R and Beaulieu  L 2006 Measurement accuracy and Cerenkov removal for high performance, high spatial resolution scintillation dosimetry Med. Phys. 33 128

Archambault L, Briere T M, Pönisch F, Beaulieu L, Kuban D A, Lee A and Beddar S 2010 Toward a real-time in vivo dosimetry system using plastic scintillation detectors Int. J. Radiat. Oncol. Biol. Phys. 78 280–7

Archambault L L, Poenisch F F, Sahoo N N, Robertson D D, Lee A A, Gillin M T M, Mohan R R and Beddar S S 2012a Verification of proton range, position, and intensity in IMPT with a 3D liquid scintillator detector system Med. Phys. 39 1239–46

Archambault  L, Polf  J  C, Beaulieu  L and Beddar  S 2008 Characterizing the response of miniature scintillation detectors when irradiated with proton beams Phys. Med. Biol. 53 1865–76

Archambault L, Therriault-Proulx F, Beddar S and Beaulieu L 2012c A mathematical formalism for hyperspectral, multipoint plastic scintillation detectors Phys. Med. Biol. 57 7133–45

Arnfield M R, Gaballa H E, Zwicker R D, Islam Q and Schmidt-Ullrich R 1996 Radiation-induced light in optical fibers and plastic scintillators: application to brachytherapy dosimetry IEEE Trans. Nucl. Sci. 43 2077–84

Attix F H 1986 Introduction to Radiological Physics and Radiation Dosimetry (New York: Wiley)Ayotte G, Archambault L, Gingras L, Lacroix F, Beddar A S and Beaulieu L 2006 Surface preparation

and coupling in plastic scintillator dosimetry Med. Phys. 33 3519–25Azangwe G, Grochowska P, Georg D and Izewska J 2014 Detector to detector corrections: a comprehensive

experimental study of detector specific correction factors for beam output measurements for small radiotherapy Med. Phys. 41 072103

Aznar M C, Andersen C E, Bøtter-Jensen L, Bäck S A J, Mattsson S, Kjaer-Kristoffersen F and Medin J 2004 Real-time optical-fibre luminescence dosimetry for radiotherapy: physical characteristics and applications in photon beams. Phys. Med. Biol. 49 1655–69

Bambynek M, Flühs D, Heintz M, Kolanoski H, Wegener D and Quast U 1999 Fluorescence 125I eye applicator Med. Phys. 26 2476–81

Bambynek M, Flühs D, Quast U, Wegener D and Soares C G 2000 A high-precision, high-resolution and fast dosimetry system for beta sources applied in cardiovascular brachytherapy Med. Phys. 27 662–7

Bartesaghi  G, Conti  V, Bolognini  D, Grigioni  S, Mascagna  V, Prest  M, Scazzi  S, Mozzanica  A, Cappelletti P and Frigerio M 2007 A scintillating fiber dosimeter for radiotherapy Nucl. Instrum. Methods Phys. Res. A 581 80–3

Beaulieu  L, Carlsson Tedgren  A, Carrier  J-F, Davis  S  D, Mourtada  F, Rivard  M  J, Thomson  R  M, Verhaegen F, Wareing T A and Williamson J F 2012 Report of the Task Group 186 on model-based dose calculation methods in brachytherapy beyond the TG-43 formalism: current status and recommendations for clinical implementation Med. Phys. 39 6208–36

Beaulieu L, Goulet M and Archambault L 2013 Current status of scintillation dosimetry for megavoltage beams J. Phys.: Conf. Ser. 012013

Becks K H, Drees J, Goldmann K, Gregor I M and Heintz M 2000 A multichannel dosimeter based on scintillating fibers for medical applications Nucl. Instrum. Methods Phys. Res. A 454 147–51

Beddar S 2015 Real-time volumetric scintillation dosimetry J. Phys: Conf. Ser. 573 12005–12Beddar S, Archambault L, Sahoo N, Poenisch F, Chen G T, Gillin M T and Mohan R 2009 Exploration

of the potential of liquid scintillators for real-time 3D dosimetry of intensity modulated proton beams Med. Phys. 36 1736–43

Beddar S and Beaulieu L 2016 Scintillation Dosimetry ed S Beddar and L Beaulieu (Boca Raton, FL: CRC Press)

Beddar  A  S, Kinsella  K  J, Ikhlef  A and Sibata  C  H 2001 A miniature scintillator-fiberoptic-PMT detectorsystem for the dosimetry of small fields in stereotactic radiosurgery IEEE Trans. Nucl. Sci. 48 924–8

Beddar A S, Law S, Suchowerska N and Mackie T R 2003 Plastic scintillation dosimetry: optimization of light collection efficiency Phys. Med. Biol. 48 1141–52

Beddar A S, Mackie T R and Attix F H 1992a Cerenkov light generated in optical fibres and other light pipes irradiated by electron beams Phys. Med. Biol. 37 925–35

Beddar A S, Mackie T R and Attix F H 1992b Water-equivalent plastic scintillation detectors for high-energy beam dosimetry: I. Physical characteristics and theoretical consideration Phys. Med. Biol. 37 1883–900

Phys. Med. Biol. 61 (2016) R305

R339

Topical Review

Beddar A S, Mackie T R and Attix F H 1992c Water-equivalent plastic scintillation detectors for high-energy beam dosimetry: II. Properties and measurements Phys. Med. Biol. 37 1901–13

Beddar A S, Suchowerska N and Law S H 2004 Plastic scintillation dosimetry for radiation therapy: minimizing capture of Cerenkov radiation noise Phys. Med. Biol. 49 783–90

Beierholm A R, Behrens C F and Andersen C E 2014 Dosimetric characterization of the Exradin W1 plastic scintillator detector through comparison with an in-house developed scintillator system Radiat. Meas. 69 50–6

Beierholm A R, Behrens C F and Andersen C E 2015 Comment on ‘characterization of the Exradin W1 scintillator for use in radiotherapy’ [Med. Phys. 42, 297–304 (2015)] Med. Phys. 42 4414–6

Birks  J  B 1951 Scintillations from organic crystals: specific fluorescence and relative response to different radiations Proc. Phys. Soc. A 64 874–7

Birks J B 1964 The Theory and Practice of Scintillation Counting (Oxford: Pergamon)Boivin J, Beddar S, Bonde C, Schmidt D, Culberson W, Guillemette M and Beaulieu L 2016 A systematic

characterization of the low-energy photon response of plastic scintillation detectors Phys. Med. Biol. 61 5569–86

Boivin  J, Beddar  S, Guillemette  M and Beaulieu  L 2015a A novel tool for in vivo dosimetry in diagnostic and interventional radiology using plastic scintillation detectors IFMBE Proc. vol 51, ed D A Jaffray (Berlin: Springer) pp 680–4

Boivin  J, Beddar  S, Guillemette  M and Beaulieu  L 2015b Systematic evaluation of photodetector performance for plastic scintillation dosimetry Med. Phys. 42 6211–20

Boon S N, Van Luijk P, Schippers J M, Meertens H, Denis J M, Vynckier S, Medin J and Grusell E 1998 Fast 2D phantom dosimetry for scanning proton beams Med. Phys. 25 464

Buranurak  S, Andersen  C  E, Beierholm  A  R and Lindvold  L  R 2013 Temperature variations as a source of uncertainty in medical fiber-coupled organic plastic scintillator dosimetry Radiat. Meas. 56 307–11

Bushberg J T, Siebert J A, Leidholdt E M Jr and Boone J M 2011 The Essential Physics of Medical Imaging ed L W Wilkins (Philadelphia, PA: Lippincott Williams & Wilkins)

Carrasco P, Jornet N, Jordi O, Lizondo M, Latorre-Musoll A, Eudaldo T, Ruiz A and Ribas M 2015a Characterization of the Exradin W1 scintillator for use in radiotherapy Med. Phys. 42 297–304

Carrasco P, Jornet N, Jordi O, Lizondo M, Latorre-Musoll A, Eudaldo T, Ruiz A and Ribas M 2015b Response to ‘comment on “characterization of the Exradin W1 scintillator for use in radiotherapy”’ [Med. Phys. 42, 297–304 (2015)] Med. Phys. 42 4417–8

Cartwright L E, Suchowerska N, Yin Y, Lambert J, Haque M and McKenzie D R 2010 Dose mapping of the rectal wall during brachytherapy with an array of scintillation dosimeters Med. Phys. 37 2247–55

Cerenkov  P  A 1937 Visible radiation produced by electrons moving in a medium with velocities exceeding that of light Phys. Rev. 52 378–9

Clift M A, Johnston P N and Webb D V 2002 A temporal method of avoiding the Cerenkov radiation generated in organic scintillator dosimeters by pulsed mega-voltage electron and photon beams Phys. Med. Biol. 47 1421–33

Clift  M  A, Sutton  R  A and Webb  D  V 2000 Dealing with Cerenkov radiation generated in organic scintillator dosimeters by bremsstrahlung beams Phys. Med. Biol. 45 1165–82

Collomb-Patton V, Boher P, Leroux T, Fontbonne J M, Batalla A and Vela A 2008 DOSIMAP: a high-resolution 2-D tissue equivalent dosemeter for linac QA and IMRT verification Radiat. Prot. Dosim. 131 100–9

Collomb-Patton V, Boher P, Leroux T, Fontbonne J M, Vela A and Batalla A 2009 The DOSIMAP, a high spatial resolution tissue equivalent 2D dosimeter for LINAC QA and IMRT verification Med. Phys. 36 317

Darafsheh  A, Zhang  R, Kanick  S  C, Pogue  B  W and Finlay  J  C 2015 Spectroscopic separation of Čerenkov radiation in high-resolution radiation fiber dosimeters J. Biomed. Opt. 20 095001–8

Das I J, Ding G X and Ahnesjo A 2008 Small fields: nonequilibrium radiation dosimetry Med. Phys. 35 206–15

de Boer  S  F, Beddar  A  S and Rawlinson  J  A 1993 Optical filtering and spectral measurements of radiation-induced light in plastic scintillationdosimetry Phys. Med. Biol. 38 945–58

Ding G X, Duggan D M and Coffey C W 2008 A theoretical approach for non-equilibrium radiation dosimetry Phys. Med. Biol. 53 3493–9

Eichmann M, Flühs D and Spaan B 2009 Development of a high precision dosimetry system for the measurement of surface dose rate distribution for eye applicators Med. Phys. 36 4634–43

Phys. Med. Biol. 61 (2016) R305

R340

Topical Review

Flühs D, Heintz M, Indenkämpen F and Wieczorek C 1996 Direct reading measurement of absorbed dose with plastic scintillators—the general concept and applications to ophthalmic plaque dosimetry Med. Phys. 23 427–34

Fontbonne J M et al 2002 Scintillating fiber dosimeter for radiation therapy accelerator IEEE Trans. Nucl. Sci. 49 2223–7

Francescon P, Beddar S, Satariano N and Das I J 2014a Variation of kQclin, Qmsr (fclin, fmsr) for the small-field dosimetric parameters percentage depth dose, tissue-maximum ratio, and off-axis ratio Med. Phys. 41 101708

Francescon P, Kilby W and Satariano N 2014b Monte Carlo simulated correction factors for output factor measurement with the CyberKnife system-results for new detectors and correction factor dependence on measurement distance and detector orientation Phys. Med. Biol. 59 N11–7

Francescon P, Kilby W, Satariano N and Cora S 2012 Monte Carlo simulated correction factors for machine specific reference field dose calibration and output factor measurement using fixed and iris collimators on the CyberKnife system Phys. Med. Biol. 57 3741–58

Frelin  A  M, Fontbonne  J  M, Ban  G, Colin  J and Labalme  M 2008a Comparative study of plastic scintillators for dosimetric applications IEEE Trans. Nucl. Sci. 55 2749–56

Frelin A M, Fontbonne J M, Ban G, Colin J, Labalme M, Batalla A, Isambert A, Vela A and Leroux T 2005 Spectral discrimination of \vCerenkov radiation in scintillating dosimeters Med. Phys. 32 3000

Frelin A M, Fontbonne J M, Ban G, Colin J, Labalme M, Batalla A, Vela A, Boher P, Braud M and Leroux  T 2008b The DosiMap, a new 2D scintillating dosimeter for IMRT quality assurance: characterization of two Cerenkov discrimination methods Med. Phys. 35 1651

Fukushima Y, Hamada M, Nishio T and Maruyama K 2006 Development of an easy-to-handle range measurement tool using a plastic scintillator for proton beam therapy Phys. Med. Biol. 51 5927–36

Gagnon  J-C, Thériault  D, Guillot  M, Archambault  L, Beddar  S, Gingras  L and Beaulieu  L 2012 Dosimetric performance and array assessment of plastic scintillation detectors for stereotactic radiosurgery quality assurance Med. Phys. 39 429–36

Ge W S 1988 A Two-dimensional Scintillation Dosimetry System for Radiotherapy (St-Louis: Washington University)

Glaser A K, Zhang R, Gladstone D J and Pogue B W 2014 Optical dosimetry of radiotherapy beams using Cherenkov radiation: the relationship between light emission and dose Phys. Med. Biol. 59 3789–811

Goulet M, Archambault L, Beaulieu L and Gingras L 2012 High resolution 2D dose measurement device based on a few long scintillating fibers and tomographic reconstruction Med. Phys. 39 4840

Goulet M, Archambault L, Beaulieu L and Gingras L 2013a 3D tomodosimetry using long scintillating fibers: a feasibility study Med. Phys. 40 101703

Goulet M, Gingras L and Beaulieu L 2010 Experimental validation of a scintillating fiber detector for real-time quality control of MLC-driven radiotherapy treatment Med. Phys. 37 3261–1

Goulet  M, Gingras  L and Beaulieu  L 2011 Real-time verification of multileaf collimator-driven radiotherapy using a novel optical attenuation-based fluence monitor Med. Phys. 38 1459–67

Goulet  M, Gingras  L and Beaulieu  L 2013b 3D tomodosimetry using scintillating fibers: proof-of-concept J. Phys.: Conf. Ser. 444 012023

Goulet  M, Rilling  M, Gingras  L, Beddar  S, Beaulieu  L and Archambault  L 2014 Novel, full 3D scintillation dosimetry using a static plenoptic camera Med. Phys. 082101 1–13

Guillot  M, Beaulieu  L, Archambault  L, Beddar  S and Gingras  L 2011a A new water-equivalent 2D plastic scintillation detectors array for the dosimetry of megavoltage energy photon beams in radiation therapy Med. Phys. 38 6763–74

Guillot M, Gingras L, Archambault L, Beddar S and Beaulieu L 2011b Spectral method for the correction of the Cerenkov light effect in plastic scintillation detectors: a comparison study of calibration procedures and validation in Cerenkov light-dominated situations Med. Phys. 38 2140–50

Guillot M, Gingras L, Archambault L, Beddar S and Beaulieu L 2013 Performance assessment of a 2D array of plastic scintillation detectors for IMRT quality assurance Phys. Med. Biol. 58 4439–54

Hakanen A T, Sipilä P M and Kosunen A 2004 Calibration of a scintillation dosemeter for beta rays using an extrapolation ionization chamber Med. Phys. 31 1123–7

Holt R W, Zhang R, Esipova T V, Vinogradov S A, Glaser A K, Gladstone D J and Pogue B W 2014 Cherenkov excited phosphorescence-based pO2 estimation during multi-beam radiation therapy: phantom and simulation studies Phys. Med. Biol. 59 5317–28

Phys. Med. Biol. 61 (2016) R305

R341

Topical Review

Hui C K, Robertson D, Alsanea F and Beddar S 2015 Fast range measurement of spot scanning proton beams using a volumetric liquid scintillator detector Biomed. Phys. Eng. Express 1 025204

Hui C, Robertson D and Beddar S 2014 3D reconstruction of scintillation light emission from proton pencil beams using limited viewing angles-a simulation study Phys. Med. Biol. 59 4477–92

Hyer  D  E, Fisher  R  F and Hintenlang  D  E 2009 Characterization of a water-equivalent fiber-optic coupled dosimeter for use in diagnostic radiology Med. Phys. 36 1711–6

Hyer D E, Serago C F, Kim S, Li J G and Hintenlang D E 2010 An organ and effective dose study of XVI and OBI cone-beam CT systems J. Appl. Clin. Med. Phys. 11 181–97

Ishikawa M et al 2009 A feasibility study of novel plastic scintillation dosimetry with pulse-counting mode. Phys. Med. Biol. 54 2079–92

Izaguirre  E  W, Price  S  G, Smajlovic  S, Yaddanapudi  S, Wooten  H and Mutic  S 2012 SU-E-T-159: sensitivity of in line real time scintillating fiber detectors for external beam treatment verification and patient safety Med. Phys. 39 3739

Jang K W, Yoo W J, Shin S H, Shin D and Lee B 2012 Fiber-optic Cerenkov radiation sensor for proton therapy dosimetry Opt. Express 20 13907–14

Jarvis L A, Zhang R, Gladstone D J, Jiang S, Hitchcock W, Friedman O D, Glaser A K, Jermyn M and Pogue B W 2014 Cherenkov video imaging allows for the first visualization of radiation therapy in real time Int. J. Radiat. Oncol. Biol. Phys. 89 615–22

Jones A K and Hintenlang D 2008 Potential clinical utility of a fibre optic-coupled dosemeter for dose measurements in diagnostic radiology Radiat. Prot. Dosim. 132 80–7

Justus B L, Falkenstein P, Huston A L, Plazas M C, Ning H and Miller R W 2004 Gated fiber-optic-coupled detector for in vivo real-time radiation dosimetry Appl. Opt. 43 1663–8

Kamio Y and Bouchard H 2014 Correction-less dosimetry of nonstandard photon fields: a new criterion to determine the usability of radiation detectors Phys. Med. Biol. 59 4973–5002

Kirov A S, Piao J Z, Mathur N K, Miller T R, Devic S, Trichter S, Zaider M, Soares C G and LoSasso T 2005 The three-dimensional scintillation dosimetry method: test for a 106Ru eye plaque applicator Phys. Med. Biol. 50 3063–81

Klein D, Briere T M, Kudchadker R, Archambault L, Beaulieu L, Lee A and Beddar S 2012 In-phantom dose verification of prostate IMRT and VMAT deliveries using plastic scintillation detectors Radiat. Meas. 47 921–9

Klein D M, Tailor R C, Archambault L, Wang L, Therriault-Proulx F and Beddar A S 2010 Measuring output factors of small fields formed by collimator jaws and multileaf collimator using plastic scintillation detectors Med. Phys. 37 5541

Knewtson T, Pokhrel S, Hernandez-Morales D, Price S, Loyalka S, Rangaraj D and Izaguirre E 2015 WE-AB-BRB-09: real time in vivo scintillating fiber array detector for medical LINACS Med. Phys. 42 3651–1

Knoll G F 1999 Radiation Detection and Measurement 3rd edn (Wiley, New York)Kojima H, Hanada T, Katsuta S, Yorozu A and Maruyama K 2009 New method for obtaining position

and time structure of source in HDR remote afterloading brachytherapy unit utilizing light emission from scintillator J. Appl. Clin. Med. Phys. 10 2983

Kroll  F, Pawelke  J and Karsch  L 2013 Preliminary investigations on the determination of three-dimensional dose distributions using scintillator blocks and optical tomography Med. Phys. 40 082104

Lacroix F, Archambault L, Gingras L, Guillot M, Beddar A S and Beaulieu L 2008 Clinical prototype of a plastic water-equivalent scintillating fiber dosimeter array for QA applications Med. Phys. 35 3682

Lacroix F, Beaulieu L, Archambault L and Beddar A S 2010a Simulation of the precision limits of plastic scintillation detectors using optimal component selection Med. Phys. 37 412

Lacroix F, Beddar A S, Guillot M, Beaulieu L and Gingras L 2009 A design methodology using signal-to-noise ratio for plastic scintillation detectors design and performance optimization Med. Phys. 36 5214–20

Lacroix F, Guillot M, McEwen M, Cojocaru C, Gingras L, Beddar A S and Beaulieu L 2010b Extraction of depth-dependent perturbation factors for parallel-plate chambers in electron beams using a plastic scintillation detector Med. Phys. 37 4331–42

Lacroix  F, Guillot  M, McEwen  M, Gingras  L and Beaulieu  L 2011 Extraction of depth-dependent perturbation factors for silicon diodes using a plastic scintillation detector Med. Phys. 38 5441–7

Lambert J, McKenzie D R, Law S, Elsey J and Suchowerska N 2006 A plastic scintillation dosimeter for high dose rate brachytherapy Phys. Med. Biol. 51 5505–16

Phys. Med. Biol. 61 (2016) R305

R342

Topical Review

Lambert  J, Nakano T, Law S, Elsey  J, McKenzie D R and Suchowerska N 2007 In vivo dosimeters for HDR brachytherapy: a comparison of a diamond detector, MOSFET, TLD, and scintillation detector Med. Phys. 34 1759

Lambert J, Yin Y, McKenzie D R, Law S H, Ralston A and Suchowerska N 2010 A prototype scintillation dosimeter customized for small and dynamic megavoltage radiation fields Phys. Med. Biol. 55 1115–26

Lambert J, Yin Y, McKenzie D R, Law S and Suchowerska N 2008 Cerenkov-free scintillation dosimetry in external beam radiotherapy with an air core light guide Phys. Med. Biol. 53 3071–80

Laub W U and Wong T 2003 The volume effect of detectors in the dosimetry of small fields used in IMRT Med. Phys. 30 341–7

Lee B, Jang K W, Cho D H, Yoo W J, Tack G-R, Chung S-C, Kim S and Cho H 2007 Measurements and elimination of Cherenkov light in fiber-optic scintillating detector for electron beam therapy dosimetry Nucl. Instrum. Methods Phys. Res. A 579 344–8

Lee B, Yoo W J, Jang K W and Cho D H 2010 Measurements of relative depth doses using fiber-optic radiation sensor and EBT film for brachytherapy dosimetry IEEE Trans. Nucl. Sci. 57 1496–501

Lee  B et  al 2008 Measurement of two-dimensional photon beam distributions using a fiber-optic radiation sensor for small field radiation therapy IEEE Trans. Nucl. Sci. 55 2632–6

Lessard F, Archambault L, Plamondon M, Després P, Therriault-Proulx F, Beddar S and Beaulieu L 2012 Validating plastic scintillation detectors for photon dosimetry in the radiologic energy range Med. Phys. 39 5308–16

Létourneau D, Pouliot J and Roy R 1999 Miniature scintillating detector for small field radiation therapy Med. Phys. 26 2555–61

Liu P Z Y, Suchowerska N, Abolfathi P and McKenzie D R 2012 Real-time scintillation array dosimetry for radiotherapy: the advantages of photomultiplier detectors Med. Phys. 39 1688–95

Liu  P  Z  Y, Suchowerska  N, Lambert  J, Abolfathi  P and McKenzie  D  R 2011 Plastic scintillation dosimetry: comparison of three solutions for the Cerenkov challenge Phys. Med. Biol. 56 5805–21

Marckmann C J, Aznar M C, Andersen C E and Bøtter-Jensen L 2006 Influence of the stem effect on radioluminescence signals from optical fibre Al2O3:C dosemeters Radiat. Prot. Dosim. 119 363–7

Mijnheer B, Beddar S, Izewska J and Reft C 2013 In vivo dosimetry in external beam radiotherapy Med. Phys. 40 070903

Morin  J, Béliveau-Nadeau  D, Chung  E, Seuntjens  J, Thériault  D, Archambault  L, Beddar  S and Beaulieu L 2013 A comparative study of small field total scatter factors and dose profiles using plastic scintillation detectors and other stereotactic dosimeters: the case of the CyberKnife Med. Phys. 40 011719

Murai T, Nakamura T and Yamamoto A 1964 A new method of integral dose measurement with a plastic scintillator phantom J. Radiat. Res. 29 23–34

Nowotny R and Taubeck A 2009 A method for the production of composite scintillators for dosimetry in diagnostic radiology Phys. Med. Biol. 54 1457–68

Olde G L and Brannen E 1959 Three-dimensional scintillation dosimeter Rev. Sci. Instrum. 30 1014Papaconstadopoulos P, Tessier F and Seuntjens J 2014 On the correction, perturbation and modification

of small field detectors in relative dosimetry Phys. Med. Biol. 59 5937–52Perera H, Williamson J F, Monthofer S P, Binns W R, Klarmann J, Fuller G L and Wong J W 1992

Rapid two-dimensional dose measurement in brachytherapy using plastic scintillator sheet: linearity, signal-to-noise ratio, and energy response characteristics Int. J. Radiat. Oncol. Biol. Phys. 23 1059–69

Petric M P, Robar J L and Clark B G 2006 Development and characterization of a tissue equivalent plastic scintillator based dosimetry system Med. Phys. 33 96

Poppe B, Thieke C, Beyer D, Kollhoff R, Djouguela A, Rühmann A, Willborn K C and Harder D 2006 DAVID—a translucent multi-wire transmission ionization chamber for in vivo verification of IMRT and conformal irradiation techniques Phys. Med. Biol. 51 1237–48

Pönisch F, Archambault L, Briere T M, Sahoo N, Mohan R, Beddar S and Gillin M T 2009 Liquid scintillator for 2D dosimetry for high-energy photon beams Med. Phys. 36 1478

Quast  U, Flühs  D and Bambynek  M 1998 Endovascular brachytherapy—treatment planning and radiation protection Herz 23 337–46

Ralston  A, Liu  P, Warrener  K, McKenzie  D and Suchowerska  N 2012 Small field diode correction factors derived using an air core fibre optic scintillation dosimeter and EBT2 film Phys. Med. Biol. 57 2587–602

Robertson D, Hui C, Archambault L, Mohan R and Beddar S 2014 Optical artefact characterization and correction in volumetric scintillation dosimetry Phys. Med. Biol. 59 23–42

Phys. Med. Biol. 61 (2016) R305

R343

Topical Review

Robertson D, Mirkovic D, Sahoo N and Beddar S 2013 Quenching correction for volumetric scintillation dosimetry of proton beams Phys. Med. Biol. 58 261–73

Rosenthal P, Weber W, Förster A, Orth O, Köhler B and Seiler F 2003 Calibration and validation of a quality assurance system for 90Sr/90Y radiation source trains Phys. Med. Biol. 48 573–85

Roussakis Y, Zhang R, Heyes G, Webster G, Mason S, Green S, Pogue B and Dehghani H 2015 Real-time Cherenkov emission portal imaging during CyberKnife(®) radiotherapy Phys. Med. Biol. 60 N419–25

Sánchez-Doblado F, Hartmann G H, Pena J, Capote R, Paiusco M, Rhein B, Leal A and Lagares J I 2007 Uncertainty estimation in intensity-modulated radiotherapy absolute dosimetry verification Int. J. Radiat. Oncol. Biol. Phys. 68 301–10

Sliski  A, Soares  C and Mitch  M  G 2006 A fibre optic scintillator dosemeter for absorbed dose measurements of low-energy x-ray-emitting brachytherapy sources Radiat. Prot. Dosim. 120 24–7

Suchowerska N, Jackson M, Lambert J, Yin Y B, Hruby G and McKenzie D R 2011 Clinical trials of a urethral dose measurement system in brachytherapy using scintillation detectors Int. J. Radiat. Oncol. Biol. Phys. 79 609–15

Tamm I E and Frank I M 1937 Coherent radiation of fast electrons in a medium Dokl. Akad. Nauk SSSR 14 107–12

Tanderup  K, Beddar  S, Andersen  C  E, Kertzscher  G and Cygler  J  E 2013 In vivo dosimetry in brachytherapy Med. Phys. 40 070902

Therriault-Proulx F, Archambault L, Beaulieu L and Beddar S 2012 Development of a novel multi-point plastic scintillation detector with a single optical transmission line for radiation dose measurement Phys. Med. Biol. 57 7147–59

Therriault-Proulx  F, Beaulieu  L, Archambault  L and Beddar  S 2013a On the nature of the light produced within PMMA optical light guides in scintillation fiber-optic dosimetry Phys. Med. Biol. 58 2073–84

Therriault-Proulx  F, Beddar  S and Beaulieu  L 2013b On the use of a single-fiber multipoint plastic scintillation detector for 192Ir high-dose-rate brachytherapy Med. Phys. 40 062101

Therriault-Proulx  F, Beddar  S, Briere  T  M, Archambault  L and Beaulieu  L 2011a Technical note: removing the stem effect when performing Ir-192 HDR brachytherapy in vivo dosimetry using plastic scintillation detectors: a relevant and necessary step Med. Phys. 38 2176–9

Therriault-Proulx F, Briere T M, Mourtada F, Aubin S, Beddar S and Beaulieu L 2011b A phantom study of an in vivo dosimetry system using plastic scintillation detectors for real-time verification of 192Ir HDR brachytherapy Med. Phys. 38 2542

Therriault-Proulx F, Wootton L and Beddar S 2015 A method to correct for temperature dependence and measure simultaneously dose and temperature using a plastic scintillation detector Phys. Med. Biol. 60 7927–39

Underwood T S A, Rowland B C, Ferrand R and Vieillevigne L 2015 Application of the Exradin W1 scintillator to determine Ediode 60017 and microDiamond 60019 correction factors for relative dosimetry within small MV and FFF fields Phys. Med. Biol. 6669–83

Westermark M, Arndt J, Nilsson B and Brahme A 2000 Comparative dosimetry in narrow high-energy photon beams Phys. Med. Biol. 45 685–702

Williamson J F, Dempsey J F, Kirov A S, Monroe J I, Binns W R and Hedtjarn H 1999 Plastic scintillator response to low-energy photons Phys. Med. Biol. 44 857–71

Wootton L and Beddar S 2013 Temperature dependence of BCF plastic scintillation detectors Phys. Med. Biol. 58 2955–67

Wootton L, Kudchadker R, Lee A and Beddar S 2014 Real-time in vivo rectal wall dosimetry using plastic scintillation detectors for patients with prostate cancer Phys. Med. Biol. 59 647–60

Yoo W J, Moon J, Jang K W, Han K-T, Shin S H, Jeon D, Park J-Y, Park B G and Lee B 2012 Integral T-shaped phantom-dosimeter system to measure transverse and longitudinal dose distributions simultaneously for stereotactic radiosurgery dosimetry Sensors 12 6404–14

Yoo W J, Shin S H, Jeon D and Han K T 2013 Development of a fiber-optic dosimeter based on modified direct measurement for real-time dosimetry during radiation diagnosis Meas. Sci. Technol.

Yoo W, Shin S, Jeon D, Hong S, Sim H, Kim S, Jang K, Cho S, Youn W and Lee B 2014 Measurement of entrance surface dose on an anthropomorphic thorax phantom using a miniature fiber-optic dosimeter Sensors 14 6305–16

Zhang R, Gladstone D J, Jarvis L A, Strawbridge R R, Jack Hoopes P, Friedman O D, Glaser A K and Pogue B W 2013 Real-time in vivo Cherenkoscopy imaging during external beam radiation therapy J. Biomed. Opt. 18 110504

Phys. Med. Biol. 61 (2016) R305