Characterization of a gated fiber-optic-coupled detector for application
in clinical electron beam dosimetry
James A. Tanyia?
Department of Radiation Medicine, Oregon Health and Science University, Portland, Oregon 97239
and Department of Nuclear Engineering and Radiation Health Physics, Oregon State University,
Corvallis, Oregon 97331
Kevin D. Nitzling and Camille J. Lodwick
Department of Nuclear Engineering and Radiation Health Physics, Oregon State University,
Corvallis, Oregon 97331
Alan L. Huston and Brian L. Justus
Optical Sciences Division, Naval Research Laboratory, Washington, DC 20375
?Received 31 July 2010; revised 17 December 2010; accepted for publication 21 December 2010;
published 28 January 2011?
Purpose: Assessment of the fundamental dosimetric characteristics of a novel gated fiber-optic-
coupled dosimetry system for clinical electron beam irradiation.
Methods: The response of fiber-optic-coupled dosimetry system to clinical electron beam, with
nominal energy range of 6–20 MeV, was evaluated for reproducibility, linearity, and output depen-
dence on dose rate, dose per pulse, energy, and field size. The validity of the detector system’s
response was assessed in correspondence with a reference ionization chamber.
Results: The fiber-optic-coupled dosimetry system showed little dependence to dose rate variations
?coefficient of variation ?0.37%? and dose per pulse changes ?with 0.54% of reference chamber
measurements?. The reproducibility of the system was ?0.55% for dose fractions of ?100 cGy.
Energy dependence was within ?1.67% relative to the reference ionization chamber for the 6–20
MeV nominal electron beam energy range. The system exhibited excellent linear response ?R2
=1.000? compared to reference ionization chamber in the dose range of 1–1000 cGy. The output
factors were within ?0.54% of the corresponding reference ionization chamber measurements.
Conclusions: The dosimetric properties of the gated fiber-optic-coupled dosimetry system compare
favorably to the corresponding reference ionization chamber measurements and show considerable
potential for applications in clinical electron beam radiotherapy. © 2011 American Association of
Physicists in Medicine. ?DOI: 10.1118/1.3539737?
Key words: real-time dosimetry, pulsed electron beam, gated fiber optics, in vivo dosimetry
In vivo dosimetry plays a crucial role in radiotherapy quality
assurance by allowing the assessment of random and/or sys-
tematic treatment deviations, essentially revealing uncertain-
ties between prescribed and administered radiotherapy doses.
The most commonly used in vivo patient dosimetry systems
include thermoluminescence dosimeters ?TLDs?, silicon di-
odes, and metal oxide semiconductor field effect transistors
?MOSFETs?. Unlike TLDs that are not coupled to a signal
processing system, diodes have the advantage of providing
real-time feedback of accumulated dose. However, the re-
sponse of diodes is dependent on a number of parameters,
including the direction of the incident radiation, dose rate,
dose per pulse, temperature, and field size,1each of which
necessitates a corrective action. Furthermore, compensatory
energy-dependent buildup layer in diode manufacturing ren-
ders them bulky to the extent that they may perturb fluences.
Mitigating the drawbacks of the diode system has propelled
MOSFETs ?Ref. 2? into a favorable alternative for clinical
radiation therapy quality assurance. However, MOSFETs
have relatively short life span, exhibit directional and
energy-dependent response, and also require frequent
To overcome the shortfalls of TLD, diode, and MOSFET
dosimetry systems, a recent approach to in vivo dosimetry
involves the utilization of ultrasmall probes consisting of
near tissue-equivalent plastic scintillators connected to opti-
cal fibers.6–9In addition to their good spatial and temporal
resolution, scintillation detectors have favorable radiation re-
sponse characteristics that include reproducibility, linearity
of response with dose, dose-rate proportionality, and energy
independence when compared to other more commonly used
detector systems. Nonetheless, these detectors have not real-
ized their full potential for routine clinical applications in
radiation oncology due mainly to the limitations imposed by
signal coupling inefficiencies and noise capture. Because
dark current input to total noise is negligible, the main con-
tribution to the aforementioned limitations thus arises from
radiation-induced light produced in the optical fibers due to a
combination of Cerenkov emission and native fluorescence
or luminescence, depending on the type of material used for
Cerenkov radiation is generated in optical fibers when
961961 Med. Phys. 38 „2…, February 20110094-2405/2011/38„2…/961/7/$30.00 © 2011 Am. Assoc. Phys. Med.
charged particles enter the core with a velocity ?c/n, where c
is the speed of light in a vacuum and n is the refractive index
of the core material? greater than the local speed of light.
Since Cerenkov radiation is created by charged particles, the
amount of Cerenkov radiation produced by an electron beam
will be significantly larger than that produced by a photon
beam. For example, for 6 MeV and 12 MeV electrons beams,
the maximum Cerenkov contribution compared to the light
output produced by a small-sized pure fused silica scintilla-
tor is of the order of ?12% and occurs when exposed at the
depths of maximum dose of the respective electron energies
in a 10?10 cm2field size.10The intensity of the radiation is
strongly dependent on the angle between the optical fiber
axis and the particle trajectory, reaching a maximum when
the Cerenkov cone is directed along the fiber axis. While
Cerenkov radiation is not directly related to the radiation
dose to the scintillator,11,12its magnitude can exceed the in-
tensity of the scintillation signal even at a wavelength where
the scintillation light is at its most intense. Thus, for accurate
dosimetric measurements, the Cerenkov radiation must be
removed from the scintillation signal.
Different methods have been proposed and tested to cor-
rect for the effects of Cerenkov radiation on scintillation do-
simetry. Beddar et al.10used a “background fiber” immedi-
ately adjacent to a signal fiber to perform background
subtraction of Cerenkov radiation from the scintillation sig-
nal. This technique, nevertheless, is limited by a potential
disparity in background signal generation in each fiber, par-
ticularly in high dose gradient regions. de Boer et al.13ex-
ploited spectral differences between scintillation light from a
long wavelength emitting scintillator and Cerenkov radiation
to reduce, but not eliminate, the interference caused by the
Cerenkov radiation. Fontbonne et al.14used a technique
known as chromatic filtration to again reduce, but not elimi-
nate, the Cerenkov background signal. Because linear-
accelerator-based radiotherapy beams are pulsed, scintilla-
tion signals may be resolved temporally from the prompt
Cerenkov and native luminescent radiation if the relaxation
time of the scintillator is long enough and the dose delivered
by the radiotherapy unit between pulses is negligible. To ob-
tain scintillation signals alone, the time interval when sam-
pling occurs is selected such that it is after the termination of
the signal induced by Cerenkov radiation and native lumi-
nescence and before the termination of scintillation signals; a
gating concept exploited by both Clift et al.15and Justus et
al.16However, unlike Clift et al.15who used a plastic scin-
tillator with relatively shorter emission times as the
radiation-sensitive material, Justus et al.16used copper ion
?Cu+?-doped quartz with relatively longer luminescence de-
cay times; hence, a more effective means of temporally re-
solving scintillation signals from Cerenkov radiation and na-
The gated Cu+-doped quartz optical fiber dosimeter16–19is
examined for clinical applications in the current study. Bene-
vides et al.20has assessed its performance for use in mam-
mography energy range, Hyer et al.21in diagnostic radiology,
and Tanyi et al.5in the therapeutic photon radiation realm. In
this study, a systematic investigation of its fundamental re-
sponse characteristics to clinical radiotherapy electron beams
II. MATERIALS AND METHODS
II.A. Gated fiber-optic-coupled dosimeter
The principle of operation of the gated fiber-optic-coupled
dosimeter ?FOCD? used in the current study has been de-
scribed elsewhere.5,16In brief, the radiation-sensitive portion
of the fiber dosimeters consists of a 1 mm length of
Cu+-doped fused quartz fiber fusion-spliced to one end of a 1
m long, multimode optical fiber. The diameter of the
Cu+-doped fused quartz fiber ?0.4 mm? is equal to that of the
silica glass core of the multimode fiber. The multimode fiber
is clad with silicone and a black Tefzel jacket to ensure light
tightness. The luminescence signal from the Cu+-doped
fused quartz is transmitted down the multimode optical fiber
to a 15-m-long fiber-optic patch cord for readout outside the
treatment room. The readout unit used for the current study
is equipped with three separate channels permitting simulta-
neous measurement of signals from three dosimeter fibers.
The readout unit uses three photon counting photomultiplier
tube modules ?model HC124-03, Hamamatsu Corporation,
Bridgewater, NJ? to detect the luminescence signal in each of
the three optical fiber dosimeters. The Cu+-doped lumines-
cence signal in each fiber dosimeter is separated from Cer-
enkov radiation and native luminescence generated in the
silica multimode fiber by using a software gating method.
Originally, gate synchronization signal was provided by an
electronic sync pulse taken from the power supply of a medi-
cal linear accelerator ?linac?.16,19However, due to variations
in the polarity, magnitude, and overall quality of the elec-
tronic sync pulses as a function of the individual linac unit,
the electronics in the dosimeter had to be modified for each
machine. In order to provide a gate synchronization signal
that could be used universally, a scattered-photon trigger
detector,5consisting of a block of plastic scintillator ?Saint-
Gobain BC-408? having dimensions of ?10 cm?2.5 cm
?2 cm, edge-coupled directly to the face of a photomulti-
plier tube module ?model HC124-03, Hamamatsu Corpora-
tion, Bridgewater, NJ?, was used. The scattered-photon trig-
ger detector was placed on a wall in the treatment room,
?4 m from the radiation source. All other features of the
gated optical fiber dosimeter system, including the software
and data acquisition procedures, have been described in
some detail previously.16Briefly, the output pulses from the
single photon counting module were collected by a 32-bit
counter configured to collect the data in a semiperiod buff-
ered counting mode. The counter sum was transferred into a
buffer on both the rising and the falling edges of each gate
pulse from the scattered-photon trigger. The signal collected
during the gate pulse includes the Cerenkov background as
well as native fiber luminescence, while the signal collected
between pulses corresponds to the Cu+-doped luminescence
signal. An 80 MHz clock was used in conjunction with an
additional 32 bit counter to accurately measure the duration
of each gate pulse and the interval between pulses. After
962Tanyi et al.: Gated fiber-optic dosimeter962
Medical Physics, Vol. 38, No. 2, February 2011
subtracting the photomultiplier tube dark noise, the accumu-
lated dose and the dose rate were calculated and presented in
II.B. Reference dosimetry
Three FOCDs were used in the current study. A 0.6 cm3
PTW N30004 ionization chamber with graphite wall with
thickness of 0.079 g/cm2, an inner aluminum electrode of
1.0 mm diameter, and an air cavity radius of 0.305 cm, in-
terfaced with a PTW UNIDOS E electrometer ?Physikalisch-
Technische Werkstaetten, Freiburg, Germany?, was used to
provide reference dosimetry to quantitate the validity of the
FOCD response. The calibration of the reference ionization
chamber and electrometer combination is traceable to an Ac-
credited Dosimetry Calibration Laboratory ?ADCL?.
II.C. Linear accelerator
A Varian Trilogy treatment platform ?Varian Medical Sys-
tems, Palo Alto, CA?, capable of producing electron beams
with nominal energies of 6, 9, 12, 16, and 20 MeV, was used
in the current study. This treatment unit is also equipped with
electron collimation ?cone? applicators with dimensions of
6?6, 10?10, 15?15, 20?20, and 25?25 cm2.
II.D. Phantom and measurement standardization
Measurements were performed in a 30?30?16.5 cm3
Plastic Water®phantom ?CNMC Co., Inc., Nashville TN?
with slabs of varying thickness. Unless otherwise stated,
measurements were performed at standard calibration setup;
that is, source-to-surface distance ?SSD? of 100 cm, cone
size of 10?10 cm2at SSD, nominal gantry angle of 0° ?In-
ternational Electrotechnical Commission or IEC standards?,
and depth of maximum dose ?1.2 cm for 6 MeV, 2.0 cm for
9 MeV, 3.0 cm for 12 MeV, 3.0 cm for 16 MeV, and 3.0 cm
for 20 MeV? as the point of measurement. With each dosim-
eter centered on the central axis of the treatment unit beam
port, the nominal delivered dose was 1 cGy for every 1
monitor unit ?MU? of radiation. Finally, for all the experi-
ments except for reproducibility measurements and unless
otherwise stated, the response of each FOCD was computed
as an average of five sequential readings, delivered and re-
corded at a machine dose rate of 600 MU/min.
II.E. Dose-rate and dose per pulse dependence
The FOCDs were irradiated with the 6 MeV electron
beam using the standard calibration configuration to investi-
gate the dependence of their response on machine dose rate.
Dose-rate response dependence was assessed for the follow-
ing machine settings: 100, 200, 300, 400, 500, 600, and 1000
MU/min. For each dose rate, a reading corresponding to 100
MU at the standard calibration position was recorded and the
results were normalized to the response at 600 MU/min for
postexposure computation of a coefficient of variation
?COV?. The dose per pulse ?that is, the product of the pulse
width and the instantaneous dose rate? response dependence
of the FOCDs was also assessed for the 6 MeV beam energy
by changing the SSD from 100 to 125 cm, in 5 cm incre-
ments, and the recorded readings normalized to the reading
at 100 cm SSD. Corresponding measurements were also per-
formed with the reference ionization chamber.
For each available electron energy, the FOCDs were re-
peatedly irradiated at standard calibration setting and 600
MU/min dose rate to 100 MU 25 times. The mean of each 25
consecutive readings was found, and the percentage differ-
ence between each consecutive reading and its corresponding
mean was calculated. A COV was also calculated and used to
define the dosimetry system’s reproducibility.
II.G. Energy dependence
Energy dependence was evaluated by computing the ratio
of the integrated scintillation output of each FOCD at a
specified depth and at depth of maximum dose for a 10
?10 cm2cone size and 100 cm SSD. For the purpose of this
study, the specified depths were 2.0 cm for 6 MeV, 3.0 cm
for 9 MeV, 4.0 cm for 12 MeV, 5.5 cm for 16 MeV, and 6.5
cm for 20 MeV. Corresponding ratios, termed ionization ra-
tios, were also computed for the reference ionization cham-
To identify the useful dynamic range of the dosimetry
system and to ensure that the scintillator output is free from
any fatigue effect for the relatively short exposure times and
accumulated doses typical of the useful dynamic range, lin-
earity measurements were performed on the FOCDs across a
range of potential clinical doses for each of the available
electron energies. The delivered doses ranged from machine
settings of 1 to 1000 MU and measurements were normal-
ized to the reading at 100 MU and compared to similar ref-
erence ionization chamber measurements.
II.I. Field size dependence
Fiber-optic-coupled dosimeter field size response depen-
dence was performed at standard calibration setup, except for
the electron cone applicator dimension that was variable, and
FIG. 1. Reproducibility of 100 cGy dose fractions from the 6 MeV beam.
963 Tanyi et al.: Gated fiber-optic dosimeter 963
Medical Physics, Vol. 38, No. 2, February 2011
reported as output factors. Thus, for each of the available
electron energies, the FOCD readings for different cones
were recorded and normalized to the reading for the 10
?10 cm2cone. Corresponding measurements were also per-
formed for the reference ionization chamber.
Figure 1 shows the percent deviation of 25 sequential
FOCD signals from their mean value as a response to 100
cGy dose fractions from the 6 MeV electron beam energy.
The reproducibility of the FOCD system was within ?0.55%
The response of the FOCD system was in excellent agree-
ment with reference ionization chamber measurements for
low dose ?that is, 1–10 cGy? and high dose ?that is, 11–1000
cGy? ranges for nominal electron beam energies in the range
of 6–20 MeV ?see Figs. 2?a? and 2?b? for the 6 MeV beam
energy?. The dose response of the FOCD system for all elec-
tron energies was linear; the linear regression correlation co-
efficients ?R2? were found to be equal to 1.000.
III.C. Dose rate and dose per pulse dependence
Figure 3 shows the average dose rate dependence of the
FOCD system at dose rates ranging from 100 to 1000 MU/
min. The FOCD response was within ?0.50% of the refer-
ence ionization chamber measurements and remained uni-
form well within the reproducibility of the FOCD system.
Figure 4 shows the relative response of the FOCDs and the
reference ionization chamber plotted against SSDs ranging
from 100 to 125 cm for the 6 MeV electron beam. The
FOCD response was within ?0.54% of the reference ioniza-
tion chamber measurements.
III.D. Field size dependence
Table II shows a comparison of the output factors mea-
sured with each FOCD and the reference ionization chamber.
The FOCD output factors were in good agreement with those
of the reference ionization chamber. Quantitatively, the
maximum difference between the FOCD response and that of
the reference ion chamber was within ?0.54% ?6 MeV?,
?0.35% ?9 MeV?, ?0.50% ?12 MeV?, ?0.42% ?16 MeV?,
and ?0.48% ?20 MeV?.
III.E. Energy dependence
Figure 5 shows the energy response of the FOCD system
for clinical electron beams in the range of 6–20 MeV. The
ratio of each FOCD’s integrated scintillation output at a
TABLE I. Percentage standard deviations for signal relative to the average
Fiber 1 Fiber 2Fiber 3
FIG. 2. Response curves of three FOCDs and a reference ionization chamber for ?a? low doses ?range: 1–10 cGy? and ?b? high doses ?range: 11–1000 cGy?
for the 6 MeV beam, measured at standard calibration settings; the linearity coefficient R2is 1.000 for the FOCDs.
FIG. 3. Dose-rate response of three FOCDs and a reference ionization cham-
ber under pulsed radiation from the 6 MeV electron beam. Dose-rate re-
sponse was normalized to unity by the response at 600 MU/min.
964Tanyi et al.: Gated fiber-optic dosimeter964
Medical Physics, Vol. 38, No. 2, February 2011
specified depth to that at depth of maximum dose for a 10
?10 cm2cone size and 100 cm SSD correlated well with
the corresponding reference ionization chamber ionization
ratio and was within ?1.67% for all the energies investi-
gated, indicating little or no energy dependence.
In the current study, the characteristics of the Cu+-doped
scintillation detector for clinical electron beams with nomi-
nal energy in the range of 6–20 MeV are assessed. The out-
put of a linear accelerator is a train of pulses, each typically
?5 ?s wide. The pulse repetition rate varies between ?50
and ?600 Hz; as such, the time interval between electron
pulses varies from ?20 ms ?corresponding to the machine
dose rate of 100 MU/min? to ?2 ms ?corresponding to the
machine dose rate of 1000 MU/min?. During each electron
pulse, the signal from each dosimeter fiber includes scintil-
lation from the Cu+-doped fiber sensor plus Cerenkov radia-
tion and native luminescence from all portions of the multi-
mode fiber exposed to the electron beam. The decay of the
Cerenkov emission is on the order of a picosecond, while the
native fluorescence decays on a nanosecond time scale;
hence, both the Cerenkov and the native luminescence emis-
sions are immediately terminated after the electron pulse ter-
minates. In contrast, the luminescence signal from the
Cu+-doped glass persists for several hundred microseconds
FIG. 4. Dose per pulse response of three FOCDs and a reference ionization
chamber under pulsed radiation from the 6 MeV electron beam. Dose per
pulse response was normalized to unity by the response at 100 cm source-
TABLE II. Output factors measured with a reference ionization chamber and three fiber-optic-coupled dosimeters
in a solid water phantom. FOCD is fiber-optic coupled dosimeter.
6 1015 2025
FIG. 5. Energy response of four detectors: Three FOCDs and a reference
ionization chamber. Energy response was computed as the ratio of each
detector’s signal output at two depths; a specified depth and at depth of
maximum dose, for a 10?10 cm2cone size and 100 cm SSD.
965 Tanyi et al.: Gated fiber-optic dosimeter 965
Medical Physics, Vol. 38, No. 2, February 2011
after the electron pulse terminates. Because of the differ-
ences in these lifetimes, the signal due to Cerenkov radiation
and the native luminescence is efficiently separated from the
scintillation signal between electron pulses by gated
detection.16Since the scintillation signal is measured after
each individual electron pulse, the measurements and the
dose response are expected to be independent of the dose rate
?Fig. 3?. The observed results confirm this expectation, also
indicating that no ion-recombination effect exists in the
FOCD system. Hence, no correction is needed for the FOCD
reading while changing the dose rate between 100 and 1000
MU/min. In addition to changes in machine dose rate, SSD
variation can, in effect, also introduce a change in the dose
per pulse at a point of interest. The results in the current
study also indicate that the FOCD response is independent of
dose per pulse changes as a function of SSD variation ?Fig.
The reproducibility of the scintillation detector under
electron irradiation was examined at 600 MU/min with each
measurement representing the integrated signal obtained
from an irradiation of 100 MU at standard calibration setting.
The percentage standard deviations of the recorded signals
are contained well within a 0.55% envelope relative to the
average reading ?Fig. 1 and Table I?.
The gated measurement of scintillation light generated by
the Cu+-doped quartz dosimeter will be a direct measure of
dose only if the quartz has ?1? high ionizing radiation sensi-
tivity and scintillation efficiency ?that is, the fraction of the
electron kinetic energy converted into detectable fluorescent
light is large and is independent of the energy of the charged
particle? and ?2? high transparency to the scintillation light.
As such, the light emitted should have a linear dependence
on the energy deposited by the charged particles interacting
within the detector. In the current study, it was demonstrated
that the conversion is linear with a proportional light yield to
deposited energy over a dose range of 1–1000 cGy for elec-
tron energies in the range of 6–20 MeV ?see Figs. 2?a? and
2?b? for the 6 MeV beam energy?.
Despite an effective atomic number of ?10.8 and a den-
sity of 2.2 g/cm3, Cu+-doped fused quartz, composed pri-
marily of silica ?99.7%?, showed no observable energy re-
sponse dependence in the range of 6–20 MeV ?Fig. 5?,
attributable in part to the constancy of the water-to-silica
collision mass stopping power ratios in the electron energy
range assessed.22Hence, no energy correction should be nec-
essary while performing in vivo measurements.
Finally, cone sizes ranging from 6?6 to 25?25 cm2
were used to investigate the influence of a potential “stem
effect” caused by fluorescence/Cerenkov radiation generated
during electron beam irradiation. For each field size, the FO-
CDs were placed on the central axis of the beam. Thus, the
length of the optical fiber that was exposed to radiation var-
ied from 3 cm ?for the 6?6 cm2cone? to 12.5 cm ?for the
25?25 cm2cone?. Because Cerenkov radiation is generated
only in that portion of the optical fiber that is exposed to
radiation, the Cerenkov interference is expected to increase
as the field size increases. The gated signal outputs of the
FOCDs as a function of field ?cone? size, compared to the
corresponding measurements from the reference ionization
chamber, are shown in Table II. There is excellent agreement
between the gated signal output of the FOCDs and the ref-
erence ionization chamber output for all cone sizes investi-
gated, indicating efficient separation of Cerenkov radiation
and native luminescence from scintillation signal by gated
Long-term stability and changes in response due to radia-
tion damage are beyond the scope of the current study and
are not examined. Regarding directional response depen-
dence, the work of Benevides et al.20has shown that the
FOCD axial-angular response is nearly uniform without any
marked variations in sensitivity. This is in conformity with
the ?2% uniform directional response reported by Miller et
al..23for therapeutic photon and electron energies. Like
Miller et al.,23Benevides et al.20also showed that the FOCD
tilt or normal-to-axial-angular response exhibited a marked
decrease in sensitivity along the long axis of the detector,
which is attributed to ?1? photon attenuation in the length of
the optical fiber connecting the dosimeter to the photodetec-
tor and ?2? decrease in cross-sectional area of the sensitive
element presented to the incident beam stemming from the
fact that the sensitive element of the FOCD is homogeneous
and cylindrically symmetric. As a result of these findings,
directional response dependence is beyond the scope of the
The Cu+-doped quartz dosimeter has demonstrated repro-
ducible dose measurements of electron beams with a linear
response to absorbed dose and response independent of dose
rate, dose per pulse, and energy. This performance is
achieved by gating the detection of the scintillation signal
with each electron pulse, thereby effectively eliminating the
background signal due to Cerenkov radiation.
a?Author to whom correspondence should be addressed. Electronic mail:
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