Characterization of a MOSkin detector for in vivo skin dose measurements during
interventional radiology procedures
M. J. Safari, J. H. D. Wong, K. H. Ng, W. L. Jong, D. L. Cutajar, and A. B. Rosenfeld
Citation: Medical Physics 42, 2550 (2015); doi: 10.1118/1.4918576
View online: http://dx.doi.org/10.1118/1.4918576
View Table of Contents: http://scitation.aip.org/content/aapm/journal/medphys/42/5?ver=pdfcov
Published by the American Association of Physicists in Medicine
Articles you may be interested in
Characterizing energy dependence and count rate performance of a dual scintillator fiber-optic detector for
computed tomography
Med. Phys. 42, 1268 (2015); 10.1118/1.4906206
Comparison of measured and estimated maximum skin doses during CT fluoroscopy lung biopsies
Med. Phys. 41, 073901 (2014); 10.1118/1.4884231
Characterization of a cable-free system based on p-type MOSFET detectors for “in vivo” entrance skin dose
measurements in interventional radiology
Med. Phys. 39, 4866 (2012); 10.1118/1.4736806
Dosimetric evaluation of the OneDose™ MOSFET for measuring kilovoltage imaging dose from image-guided
radiotherapy procedures
Med. Phys. 37, 4880 (2010); 10.1118/1.3483099
A new approach in dose measurement and error analysis for narrow photon beams (beamlets) shaped by
different multileaf collimators using a small detector
Med. Phys. 31, 2020 (2004); 10.1118/1.1760191
Characterization of a MOSkin detector for in vivo skin dose measurements
during interventional radiology procedures
M. J. Safari, J. H. D. Wong, and K. H. Nga)
Department of Biomedical Imaging, Faculty of Medicine, University of Malaya,
Kuala Lumpur 50603, Malaysia and University of Malaya Research Imaging Centre, Faculty
of Medicine, University of Malaya, Kuala Lumpur 50603, Malaysia
W. L. Jong
Clinical Oncology Unit, Faculty of Medicine, University of Malaya, Kuala Lumpur 50603, Malaysia
D. L. Cutajar and A. B. Rosenfeld
Centre for Medical Radiation Physics, University of Wollongong, Wollongong, NSW 2522, Australia
(Received 3 September 2014; revised 20 March 2015; accepted for publication 4 April 2015;
published 23 April 2015)
Purpose: The MOSkin is a MOSFET detector designed especially for skin dose measurements. This
detector has been characterized for various factors affecting its response for megavoltage photon
beams and has been used for patient dose measurements during radiotherapy procedures. However,
the characteristics of this detector in kilovoltage photon beams and low dose ranges have not been
studied. The purpose of this study was to characterize the MOSkin detector to determine its suitability
for in vivo entrance skin dose measurements during interventional radiology procedures.
Methods: The calibration and reproducibility of the MOSkin detector and its dependency on different
radiation beam qualities were carried out using RQR standard radiation qualities in free-in-air geometry. Studies of the other characterization parameters, such as the dose linearity and dependency on
exposure angle, field size, frame rate, depth-dose, and source-to-surface distance (SSD), were carried
out using a solid water phantom under a clinical x-ray unit.
Results: The MOSkin detector showed good reproducibility (94%) and dose linearity (99%) for the
dose range of 2 to 213 cGy. The sensitivity did not significantly change with the variation of SSD
(±1%), field size (±1%), frame rate (±3%), or beam energy (±5%). The detector angular dependence
was within ±5% over 360◦ and the dose recorded by the MOSkin detector in different depths of a solid
water phantom was in good agreement with the Markus parallel plate ionization chamber to within
±3%.
Conclusions: The MOSkin detector proved to be reliable when exposed to different field sizes, SSDs,
depths in solid water, dose rates, frame rates, and radiation incident angles within a clinical x-ray
beam. The MOSkin detector with water equivalent depth equal to 0.07 mm is a suitable detector for
in vivo skin dosimetry during interventional radiology procedures. C 2015 American Association of
Physicists in Medicine. [http://dx.doi.org/10.1118/1.4918576]
Key words: MOSFET detector, MOSkin, skin dose monitoring, skin dosimetry, interventional
radiology, in vivo dosimetry
1. INTRODUCTION
According to the United Nations Scientific Committee on the
Effects of Atomic Radiation (UNSCEAR) report, 3.6 × 109
diagnostic radiology x-ray examinations are performed worldwide annually, with this number increasing every year.1 Until
the late 1980s, diagnostic procedures were characterized as
low dose radiation procedures and were only linked with stochastic risks. Since the 1980s, fluoroscopically guided interventional procedures have become widespread and have been
used effectively to diagnose and treat numerous vascular and
cardiac diseases. Although interventional procedures provide
enormous advantages over invasive surgical procedures, long
periods of radiation exposure may increase the risk of deterministic effects in patients, thus causing radiation-induced skin
injuries.2–7
2550
Med. Phys. 42 (5), May 2015
The US Food and Drug Administration (FDA),8 the World
Health Organization (WHO),9 the International Commission
on Radiological Protection (ICRP),10 and the International
Atomic Energy Agency (IAEA)11 have all expressed concerns
regarding patient skin dose. They have also issued guidance
on the prevention of skin injuries in high dose interventional
procedures. In order to prevent severe radiation injuries, it is
important to evaluate the entrance skin dose (ESD) of patients
during long irradiation periods. To address these issues, several
radiation dose tracking systems have been developed and are
available for purchase. The Patient Exposure Monitoring Network (PEMNET®) System (Clinical Microsystems, Inc., Arlington, VA) was designed to calculate and display the realtime exposure rate and subsequently the patient’s ESD based
on the exposure parameters and the patient geometry information.12 The PEMNET®, however, does not differentiate the
0094-2405/2015/42(5)/2550/10/$30.00
© 2015 Am. Assoc. Phys. Med.
2550
2551
Safari et al.: Calibration of MOSkin for skin dose measurement
2551
F. 1. Schematic of (a) MOSkin dosimetry system, (b) MOSFET detector with an epoxy bubble encapsulation above the sensitive area, (c) MOSkin detector
face-up orientation, and (d) MOSkin detector face-down orientation.
radiation locations on the patient’s skin and cannot provide
the spatial dose distribution information required to calculate
the ESD. Various indirect beam-monitoring quantities have
been used to estimate the ESD in fluoroscopically guided
procedures, such as the kerma-area-product (KAP), fluoroscopy time, number of images, and dose level at the interventional reference point (IRP).13,14 Although these quantities are
widely used as an indicator for the entrance skin dose, several
studies have shown that they provide an unrealistic estimation
of the ESD in interventional radiology procedures.13–16
Recently, several skin dose estimation tools have been
developed using exposure parameters (exposure rate, kVp,
mAs, exposure angle, table height, etc.). This information is
extracted from DICOM tags and registered to standard and
anatomical patient model phantoms.17,18 These systems estimate the skin dose during IR procedures by taking into account
various parameters, such as the dose level at IRP, corrected
source skin distance, backscatter factor, and correction of mass
energy absorption coefficients of skin to air. The limitations
of these systems are that they have been implemented by a
limited number of manufacturers (e.g., Toshiba Dose Tracking
System) and are not commercially available in all fluoroscopy
machines. The anatomical model phantoms do not represent
the exact physical shape of the patients and mispositioning
of the patient can be an issue when considering the dose
evaluation uncertainty.
The evaluation of the dose absorbed by the skin of patients can be achieved by means of a direct dosimetry method.
Currently, different direct dosimetry methods have been used
to measure skin dose during interventional radiology proceMedical Physics, Vol. 42, No. 5, May 2015
dures: thermoluminescent dosimeters (TLDs),15,19 radiochromic films,15,20–22 glass dosimeters,23–25 MOSFET radiation
sensors,26–28 and scintillator dosimeters.29 The desirable properties of a detector for monitoring ESD during interventional
procedure include small physical size to preserve the image
quality, ability to track the dose every second, linear response
within the measured dose range, ease of usage, and an appropriate water equivalent depth (WED) for monitoring the skin
dose. According to the ICRP Report 59, for accurate skin dose
measurements, the WED value associated with the detector
should be equal to 0.07 mm.30 This depth is approximately
equal to the average depth of the basal layer of the epidermis
that comprises the most radiosensitive epithelial cells.
The MOSkin detector is a new type of MOSFET detector
developed and prototyped by the Centre for Medical Radiation Physics (CMRP), University of Wollongong, Australia
[Fig. 1(a)]. The MOSkin detector was designed especially for
skin dose measurements.31 This detector has previously been
tested and found to be suitable for skin dose measurements
in radiation therapy.31–33 This research focuses on the characterization of the MOSkin to determine its suitability for use
as an in vivo skin dosimeter during interventional radiology
procedures.
2. MATERIALS AND METHODS
The MOSkin detector was designed using a new packaging
structure in which the p-MOSFET sensor with a thick gate
oxide is hermetically sealed within a Kapton pigtail strip using
“drop-in” packaging technology [Fig. 1(a)]. A thin polyamide
2552
Safari et al.: Calibration of MOSkin for skin dose measurement
2552
T I. RQR standard radiation qualities.
Radiation beam
quality
RQR3
RQR4
RQR5
RQR6
RQR7
RQR8
RQR9
a
Tube potential
(kVp)
Effective energy
(keV)a
Added filtration
(mm Al)
Half value layer
(mm Al)
50
60
70
80
90
100
120
27.04
29.23
31.66
34.02
36.05
38.66
42.76
2.7
2.8
3.1
3.2
3.4
3.7
4.0
1.8
2.1
2.5
3.0
3.4
4.1
5.1
Effective energies were obtained from NIST website (Ref. 37).
film layer works as a moisture protector and build-up layer,
providing a WED of approximately 0.07 mm in tissue,31,34 as
compared to commercial MOSFET detectors, which utilize
wire-bonding and an epoxy bubble encapsulation above the
sensitive region [Fig. 1(b)]. The principle behind the operation
of MOSkin detectors is based on the shift of the threshold
voltage due to electron–hole generation within the gate oxide,
followed by the capture of holes on border traps. The threshold
voltage change, ∆Vth, is proportional to the absorbed dose. The
detector sensitivity is defined as
Sensitivity =
∆Vth (mV)
.
Dose (cGy)
(1)
The MOSkin detector has a finite lifetime due to the accumulation of radiation dose, which saturates the hole-traps
near the Si–SiO2 interface. The lifetime threshold voltage
was reported to be approximately 24 V, determined using
the CMRP designed electrometer rather than the MOSkin
detectors themselves.35 In this research, the MOSkin detector
response in terms of energy, field size, exposure angle, sourceto-surface distance (SSD), and dose rate and percent depthdose (PDD) dependences under various diagnostic beam qualities was investigated.
The MOSkin detector calibration and dependence on radiation energy were carried out using the standard radiation quality (RQR: radiation qualities in radiation beams emerging from
the x-ray tube assembly) according to IEC 61267 (Ref. 36) in
free air geometry. The free-in-air geometry was achieved by
placing the detectors on an elevated platform made of styrofoam. The radiation beams were generated by a Y-TU 160D02 x-ray machine (COMET AG, Flammat, Switzerland),
at the Secondary Standard Dosimetry Laboratory (SSDL) of
the Malaysian Nuclear Agency. Table I shows the parameters
of the various RQR standard qualities. The MOSkin detector
linearity and reproducibility, angular dependence, field size
dependence, PDD in solid water, dose rate dependence, frame
rate dependence, and source-to-surface distance dependence
were studied using a clinical C-arm fluoroscopy unit (Philips
Allura Xper FD20/20® x-ray unit, Philips Healthcare, Amsterdam, Netherlands) at the University of Malaya Medical Centre
(UMMC). Characterization procedures were carried out using
a 30×30×12 cm3 water equivalent plastic phantom (Gammex
457, Gammex, Middleton, WI) and the x-ray tube was positioned at a gantry angle of 180◦, perpendicular to the phantom
surface [Fig. 2(a)]. The detector was characterized for an 80
kVp photon beam (effective energy 42.7 keV) using nonmag-
F. 2. Characterization setups with C-arm x-ray tube fluoroscopy unit, (a) standard setup and (b) measurement setup for angular dependency test.
Medical Physics, Vol. 42, No. 5, May 2015
2553
Safari et al.: Calibration of MOSkin for skin dose measurement
2553
nified acquisition mode imaging (field of view: 48 cm), while
the exposure frame rate was fixed to 3 frames/s. The flat
panel detector was placed 120 cm from the x-ray tube focal
spot and a field size of 10 × 10 cm2 was used. This setup is
henceforth called the “standard setup” [Fig. 2(a)]. The MOSkin
detector was placed facing the x-ray tube unless otherwise
stated. The characterization procedures under RQR standard
radiation qualities were benchmarked against a 30 cm3 parallel plate ionization chamber (model 233612, PTW, Freiburg,
Germany), while measurements under the Philips Allura Xper
FD20/20® unit were benchmarked against a 0.055 cm3 Markus
parallel plate ion chamber (model 23343, PTW, Freiburg, Germany). All measurements were repeated three times and the
standard deviation (1 SD) of the results was reported.
Two types of uncertainties were considered in the analysis
of the MOSkin characteristics. The MOSkin reader shows the
threshold voltage change with an uncertainty of 1 mV, and the
immediate readout of the detector after irradiation can generate
voltage creep-up, up to 4 mV.31 The creep-up voltage depends
on the time between successive readouts, which peaks 10 s
following the end of irradiation. This component of uncertainty can be reduced with a one-minute post-irradiation wait
time.38 The significance of these uncertainties is dependent on
the total dose delivery to the MOSkin detector, and it is negligible during high dose delivery when the change in threshold
voltage is large. In this research, the dose delivered to the
detector was in the lower range; therefore, these uncertainties
were taken into account in the results. Furthermore, a oneminute postirradiation wait time was used before readout. A
study on temperature dependence of the MOSFET detector reported a threshold voltage variation of 50 mV for a temperature
change from 20 to 40 ◦C (Ref. 39) for the used readout current.
To avoid a temperature effect on MOSkin detector reading, the
MOSkin detector was placed on the styrofoam or solid water
phantom, inside the measurement room, for approximately 5
min before starting the measurement to allow for temperature equalization. The ambient temperature was continuously
monitored throughout the experimental procedures.
Automatic brightness control (ABC) was used to control
the x-ray tube output. ABC controls the light output of the
image receptor by adjusting the kVp and/or mA of the x-ray
tube using a preprogrammed kVp-mA curve.40 The small size
and design of the MOSkin detector do not change the ABC
exposure parameters when it is in the beam. In this study,
the MOSkin detector was set up, per the standard setup, and
the sensitivity of the MOSkin detector was studied for the
dose range of 2 to 213 cGy. The MOSkin’s response was
benchmarked against the Markus chamber. Any deviation in
exposure time and/or mAs was corrected based on the exposure parameters recorded from the console (mAs, number of
images, KAP). The reproducibility of the MOSkin detector
was assessed based on the mean of the standard deviations for
three sets of measurements.
2.A. Calibration
2.E. Depth-dose measurements
The MOSkin detector was calibrated under RQR7 beam
quality (effective energy 36.1 keV) and the exposure fixed at
140 mAs. The detector was placed on a styrofoam board at
a SSD of 100 cm and field size of 13 cm diameter. The PTW
ion chamber measurements for comparison were corrected for
temperature, pressure, and energy, while the reproducibility
of the PTW ion chamber was measured to be better than
99%.
MOSkin and Markus detectors were initially set up, per the
standard setup, followed by placement at different depths in
a solid water phantom. For all depth-dose measurements, the
detectors were placed along the beam central axis separately to
minimize the uncertainty caused by the detector positioning.
Depth-dose measurements were tested for 0–35 mm depths
from the surface of the solid water phantom.
2.B. Energy dependence measurements
Semiconductor detectors are known to be energy-dependent, particularly in the kV beam energy range. The high
atomic number (Z) of the detector sensitive volume (silicon
oxide, Z = 14) is expected to over-respond at low kV energies due to the photoelectric absorption effect. This study
Medical Physics, Vol. 42, No. 5, May 2015
was carried out to investigate the magnitude of the energy
dependence of the MOSkin detector for the beam energy range
commonly used in diagnostic radiological procedures. The
MOSkin energy dependence was studied for the beam energy
range of RQR3 to RQR9 (effective energy 27.04–42.76 keV).
The exposure was fixed at 140 mA s. The MOSkin detector was
placed at a SSD of 100 cm and a field size of 13 cm diameter
was used.
2.C. Dose linearity and reproducibility measurements
2.D. Field size dependence measurements
Field size dependence is an important factor for point dose
recording during interventional radiology procedures since
multiple field sizes are often used within a treatment procedure. The ABC system adjusts the tube output to maintain the
image brightness at a constant level for different exposure field
sizes. As the exposure field size decreases, the tube current
increases, and subsequently, the exposure dose increases. The
MOSkin was set up, per the standard setup, at a SSD of 70 cm
with selected field sizes from 5 × 5 cm to 20 × 20 cm.
2.F. Source-to-surface distance dependence
The MOSkin detector was set up, per the standard setup,
and placed at a SSD of 90 cm with an exposure field size
of 10 × 10 cm, with the SSD adjusted by moving the couch
toward the x-ray tube. The detector response at different distances (70–90 cm) was studied. The MOSkin detector dose rate
dependence was also evaluated using this setup.
2554
Safari et al.: Calibration of MOSkin for skin dose measurement
2554
F. 3. Energy dependence of the MOSkin detector for diagnostic energy range, under RQR standard beam qualities. The values were normalized to 1 at
36.1 keV and the average standard deviation of three measurements is presented.
2.G. Frame rate dependence
The Philips Allura Xper FD20/20® system classifies the
frame rate for angiography procedures into two main categories, i.e., cardiac and vascular applications. The vascular
application allows operators to select acquisition exposures
with frame rates of 2, 3, 4, and 6 frames/s while the frame rates
of 15 and 30 frames/s are for cardiac applications. In this study,
the MOSkin detector was set up, per the standard setup, and
both application types were used to study the dependence of
the MOSkin detector to the exposure frame rate. The exposure
parameters were recorded from the console (kV, mA, ms, and
filter). The effective energies (keV) were calculated based on
measured HVLs using an Unfors detector (Unfors Raysafe
AB, Billdal, Sweden).
2.H. Angular dependence measurements
The C-arm x-ray tube can rotate from 120◦ left anterior
oblique (LAO) to 180◦ right anterior oblique (RAO) and hence,
cannot cover the entire angular range. Due to this limitation,
the solid water phantom was placed at the edge of the patient
support couch [Fig. 2(b)] at a SSD of 81 cm (isocenter of
rotation). The MOSkin detector was placed in the center of the
surface of the solid water phantom on the central axis.
The C-arm x-ray tube was rotated from 30◦ RAO to 150◦
RAO in 20◦ intervals, and the angular response of the MOSkin
detector was assessed for face-up and face-down orientations
[Figs. 1(c) and 1(d)]. Other exposure parameters followed the
standard setup.
3. RESULTS AND DISCUSSION
3.A. Detector calibration
The sensitivity of the MOSkin detector for a RQR7 beam
quality (effective energy 36.1 keV) photon beam was measured
to be 11.56 ± 0.36 mV cGy−1. Previous research reported
that the sensitivity of the MOSkin detector for 150 kVp xray beam (effective energy 64.87 keV) was approximately
6.70 mV cGy−1, which was measured in comparison with
EBT2 film.41 The sensitivity of the MOSkin detector in a megavoltage beam was 2.63 ± 0.01 mV cGy−1.42 The higher sensitivity measured at lower beam energy is due to the increasing
dominance of photoelectric absorption and consequently, an
F. 4. Linearity of the MOSkin detector benchmarked against the Markus ion chamber. Average standard deviation of this measurement was ±2 mV.
Medical Physics, Vol. 42, No. 5, May 2015
2555
Safari et al.: Calibration of MOSkin for skin dose measurement
2555
F. 5. The field size effects on response for the MOSkin detector. All the readings were normalized to a field size of 10 × 10 cm.
increasing response of the MOSkin detector at low energy. The
MOSkin detector sensitivity decreases as the cumulative dose
increases. Our recent study showed that the MOSkin detector’s
sensitivity decreased by 0.09 mV/cGy with every increase
of 10 Gy accumulated dose in megavoltage radiation, which
is equivalent to sensitivity reduction of 1.5%/V increase in
the threshold voltage.43 Hence, it is recommended to recalibrate the detector periodically over its useful lifetime based
on the accuracy needed. The lifetime of a MOSkin detector is
determined by the initial voltage (∼10 V), the sensitivity of
detector, and the maximum threshold voltage (∼24 V). When
used under diagnostic beam energies, the detector lifetime is
approximately equivalent to 12 Gy of radiation exposure.
3.B. Energy dependence
Figure 3 shows the MOSkin detector’s response, normalized to RQR7 beam quality. As expected, the MOSkin detector showed enhanced response at lower beam energies. The
detector over-responded by a factor of 1.15 at 50 kVp (effective beam energy 27.04 keV). At 120 kVp (effective beam
energy 42.76 keV), the detector under-responded by a factor
of 0.86. However, for the tube potentials that are commonly
used in interventional radiology procedures (70–100 kVp),
the detector response varies within ±5%. The ratio of massenergy absorption coefficient of silicon (Si) to air was defined
as
Ratio of mass-energy absorption coefficient =
(2)
where (µen/ρ)Si and (µen/ρ)air are mass-energy absorption
coefficients of Si and air, respectively. The measured ratio was
normalized to 1 at 36.1 keV. As Fig. 3 illustrates, the change of
the MOSkin detector’s sensitivity for different effective beam
energies shows the same tendency with the ratio of the massenergy absorption coefficient of Si to air. This comparison
does not take into account the change of recombination of
electron-hole pairs in the gate of the MOSFET with decreasing
photon energy.44 The ratio of the mass-energy absorption
coefficients of Si to air has the same trend as the ratio of Si to
water.
F. 6. Depth dose in solid water for FSD of 10 × 10 cm. All the readings are normalized to the surface dose (100%).
Medical Physics, Vol. 42, No. 5, May 2015
(µen/ρ)Si
,
(µen/ρ)air
2556
Safari et al.: Calibration of MOSkin for skin dose measurement
2556
F. 7. Source-to-surface distance and dose rate dependence of the MOSkin detector for an 80 kVp x-ray beam with an exposure field size of 10 × 10 cm.
3.C. Dose linearity and reproducibility
The MOSkin detector showed a linear response with the
amount of delivered dose (2–213 cGy) measured by Markus
ion chamber to be better than 0.99 (Fig. 4). The reproducibility
of the MOSkin measurements was found to be better than
94%.
3.D. Field size dependence
The ABC controlled the brightness of the image for different exposure field sizes by adapting the milliampere-second. A
smaller exposure field size had a higher beam exposure (mAs).
For this study, the tube potential remained constant throughout
the measurements at 80 kVp (effective beam energy 42.76
keV), and the beam exposure changed from 61 to 8 mAs
for 5 × 5 cm to 20 × 20 cm exposure field sizes, respectively.
Figure 5 illustrates that the MOSkin detector is independent of
field size variation (within ±1%).
3.E. Depth-dose data
The percentage depth-dose response of the MOSkin detector was investigated previously,41,45 which compared the
MOSkin detector response with Markus chamber and Monte
Carlo simulations in water. In their reports, the MOSkin detector showed an over-response compared with the Markus
chamber measurements,41 while Monte Carlo simulations
showed MOSkin to have good agreement with simulated water
response (±3%).45 In this study, Fig. 6 shows the percentage
depth-dose curves for the MOSkin and Markus ion chamber.
Both detector responses were normalized to their response at
the surface of the solid water phantom. The MOSkin detector
appears to be in good agreement with the Markus chamber to
within ±3%. The largest deviation between the MOSkin and
Markus detectors’ response was 4% at 2 mm. These data are
in good agreement with the Monte Carlo simulations for the
100 kVp beam (effective photon energy of 48.6 keV).45 The
MOSkin detector for this study was calibrated and measured
for the effective photon energy of 42.76 keV.
3.F. Source-to-surface distance dependence
Figure 7 shows the dose measured by the MOSkin and
the Markus chamber corrected for inverse square law. The
MOSkin showed good agreement with the Markus chamber
measurements (±1%). The error bars increased with greater
distance due to the lower dose delivered to the detector. Consequently, the effects of reader and voltage creep uncertainties
become noticeable. The SSD parameter is an important feature
T II. Effective energy and dose rate values for different frame rates of cardiac and vascular applications, controlled by AEC system.
Application
types
Vascular
Cardiac
a
Frame rate
(frames/s)
Tube potential
(kVp)
Added filtration
(mm)
HVL
(mm Al)
Effective energy (keV)a
Dose rate (mGy/s)
2
3
4
6
15
30
83.50
83.48
83.45
78.64
70.84
70.93
0.1 Cu + 1.0 Al
0.1 Cu + 1.0 Al
0.1 Cu + 1.0 Al
0
0
0
5.07
5.09
5.07
3.22
2.57
2.58
42.54
42.61
42.54
34.84
31.55
31.60
1.68
1.99
2.53
3.68
0.58
1.19
Effective energies were obtained from NIST website (Ref. 37).
Medical Physics, Vol. 42, No. 5, May 2015
2557
Safari et al.: Calibration of MOSkin for skin dose measurement
2557
F. 8. MOSkin detector frame rate and dose rate dependences of vascular and cardiac applications.
for interventional radiology dosimetry, as the x-ray tube is
continuously rotated around the patient during interventional
radiology procedures. Depending on the location of the patient’s body, the x-ray tube is placed at different distances from
the body surface. The FDA has expressed concerns about SSD
in fluoroscopic procedures and has suggested using a SSD
of 38 cm as a threshold for stationary fluoroscopy units and
30 cm for mobile fluoroscopy units.46 Figure 7 also shows the
relative response of the MOSkin and Markus detectors, normalized to their ratio at 1.2 mGy/s dose rate. MOSkin detector
shows a very low dependence on exposure dose rate (<1%).
3.G. Frame rate dependence
Table II shows the different frame rates available on the
Philips Allura Xper FD20/20® system, grouped under cardiac
and vascular applications. Note that when switching from the
high frame rates to the low frame rates, the system automatically adjusted the tube potential (filtration) and dose rate.
Figure 8 shows that MOSkin detector has low dependence
on frame rate and dose rate variations (coefficient of variation
= ±3%). This variation is within the uncertainty due to the
energy dependence of MOSkin detector.
3.H. Angular dependence
The angular dependence of the MOSkin detector is presented in Fig. 9. The variation in the readings for the azimuth
axis was within ±5%. The lowest drop in sensitivity occurred
at −60◦ face-up and the maximum sensitivity was obtained
at 0◦ face-down. The MOSkin detector has a different WED
for face-down and face-up orientations (inherent anisotropy)
on the surface of the phantom. This detector has a thicker
build-up layer due to the silicon substrate in the case of facedown orientation [Figs. 1(c) and 1(d)], which produces more
secondary electrons and subsequently causes dose enhancement at the sensitive volume of the MOSkin. On the other
hand, photon attenuation in the silicon substrate exists as well,
with the response of the device affected by a combination of
the two processes. The MOSkin detector shows slightly higher
sensitivity in a face-down orientation relative to a face-up
orientation. As Fig. 9 shows, the MOSkin detector’s sensitivity
F. 9. The azimuth angular dependence of the MOSkin detector. The MOSkin response was normalized to the 0◦ value in face-up orientation.
Medical Physics, Vol. 42, No. 5, May 2015
2558
Safari et al.: Calibration of MOSkin for skin dose measurement
decreases with increasing exposure angle. This change in the
MOSkin detector sensitivity is due to the intrinsic angular
dependence of MOSkin and the effect of backscatter radiation.
Published reports on the angular dependence of the MOSkin
detector for a 6 MV x-ray field reported that this detector
has a very low intrinsic dependence on exposure angle31,47
and the minimum angular dependence for this detector was
measured for a dual MOSkin to within ±2.5% around the
azimuth axis under charged particle equilibrium conditions.42
For kilovoltage x-rays, the field contribution of the intrinsic
response to the skin dose measurements with a single face on
MOSkin is slightly larger in comparison to a dual MOSkin in
a 6 MV field but is still small enough for this application.
and N. Abdullah at the Medical Physics group of Malaysian
Nuclear Agency, Dr. N. M. Ung at the Clinical Oncology Unit,
University of Malaya, C. C. Lee and M. Mozaker, radiographers at the UMMC, and K. H. Lam at Philips Healthcare
Malaysia and Medical Physics Unit (UMMC) for the loan
of equipment for this research. The authors would like to
thank the reviewers, editor, and associate editor for their
time and effort in improving this manuscript. This study
was supported by the High Impact Research (HIR) grant,
UM.C/625/1/HIR/MOHE/MED/38, Account No. H-2000100-E000077, and PPP grant, PG035- 2013A from the University of Malaya.
a)Author
4. CONCLUSION
In this study, the MOSkin detector was characterized for its
diagnostic beam energies. The dose linearity, reproducibility,
and depth-dose measurements were performed. The detector’s
response with variations of beam energy, field size, sourcesurface distance, frame rate, and beam incidence angles was
also investigated. This work demonstrated that the MOSkin detector is suitable for monitoring the skin dose during interventional radiology procedures, taking into account the various
uncertainties and limitations of the detector.
The limitation of this system is that the MOSkin as a point
dosimetry system may not be placed at the location where the
skin experiences the most intense radiation dose. Hence, there
may be an underestimation of the MSD during IR procedures.
For this purpose, a 2D dosimeter, such as film, would still provide the best solution. Nevertheless, using multiple MOSkin
detectors simultaneously to measure dose at selected regions
of interest or radiosensitive organs, such as the eye lens, would
still be useful in providing insights into the radiation dose
received by these organs. Compared to other commercially
available dose tracking systems, which predict the 2D dose
distribution for patients via parameters from the fluoroscopic
machine, this system does not provide 2D dose distribution
information. Nevertheless, this system tracks the dose every
second using actual direct skin dose measurements. Similar
to other dose tracking systems, this system also enables interventional radiologists to balance the expected clinical benefits
and radiation risks of performing a procedure. Skin dose is
an important issue for patient radiation safety during interventional radiology procedures. The MOSkin detector, with
WED equal to 0.07 mm, is a suitable detector for in vivo skin
dose measurements. Future work would include the application of this detector system in clinical interventional radiological procedures. This detector can be used as an eye lens
dose tracking system during neuro-interventional procedures
without interfering with the treatment procedure, due to the
small physical size and transparency on the patient’s image.
ACKNOWLEDGMENTS
The authors acknowledge and thank the personnel at the
following institutions for their support in this study: M. J. Isa
Medical Physics, Vol. 42, No. 5, May 2015
2558
to whom correspondence should be addressed. Electronic mail:
ngkh@um.edu.my
1United Nations Scientific Committee on the Effects of Atomic Radiation,
UNSCEAR 2008 Report to the General Assembly (United Nations, New
York, NY, 2010).
2T. B. Shope, “Radiation-induced skin injuries from fluoroscopy,” Radiographics 16(5), 1195–1199 (1996).
3E. Vano, L. Arranz, J. M. Sastre, C. Moro, A. Ledo, M. T. Garate, and
I. Minguez, “Dosimetric and radiation protection considerations based on
some cases of patient skin injuries in interventional cardiology,” Br. J.
Radiol. 71(845), 510–516 (1998).
4L. K. Wagner, M. D. McNeese, M. V. Marx, and E. L. Siegel, “Severe skin
reactions from interventional fluoroscopy: Case report and review of the
literature,” Radiology 213(3), 773–776 (1999).
5T. H. Frazier, J. B. Richardson, V. C. Fabre, and J. P. Callen, “Fluoroscopyinduced chronic radiation skin injury: A disease perhaps often overlooked,”
Arch. Dermatol. 143(5), 637–640 (2007).
6S. Balter, J. W. Hopewell, D. L. Miller, L. K. Wagner, and M. J. Zelefsky,
“Fluoroscopically guided interventional procedures: A review of radiation
effects on patients’ skin and hair,” Radiology 254(2), 326–341 (2010).
7A. Spiker, Z. Zinn, W. H. Carter, R. Powers, and R. Kovach, “Fluoroscopyinduced chronic radiation dermatitis,” Am. J. Cardiol. 110(12), 1861–1863
(2012).
8Food and Drug Administration, Public Health Advisory: Avoidance of
Serious X-ray-Induced Skin Injuries to Patients During FluoroscopicallyGuided Procedures (Center for Devices and Radiological Health, Rockville,
MD, 1994).
9World Health Organization and Institut für Strahlenhygiene des Bundesgesundheitsamtes, Efficacy and Radiation Safety in Interventional Radiology
(World Health Organization, Switzerland, Geneva, 2000).
10J. Valentin, “Avoidance of radiation injuries from medical interventional
procedures,” Ann. ICRP 30(2), 7–67 (2000).
11International Atomic Energy Agency, “Patient dose optimization in fluoroscopically guided interventional procedures,” IAEA-TECDOC-1641: IAEA
(2010).
12J. T. Cusma, M. R. Bell, M. A. Wondrow, J. P. Taubel, and D. R. Holmes, Jr.,
“Real-time measurement of radiation exposure to patients during diagnostic
coronary angiography and percutaneous interventional procedures,” J. Am.
Coll. Cardiol. 33(2), 427–435 (1999).
13J. W. Jaco and D. L. Miller, “Measuring and monitoring radiation dose during fluoroscopically guided procedures,” Tech. Vasc. Interv. Radiol. 13(3),
188–193 (2010).
14S. Balter, D. W. Fletcher, H. M. Kuan, D. Miller, D. Richter, H. Seissl,
and T. B. Shope, “Techniques to estimate radiation dose to skin during
fluoroscopically guided procedures,” AAPM Summer School Proceedings,
Madison, WI (2002).
15E. Vano, L. Gonzalez, J. I. Ten, J. M. Fernandez, E. Guibelalde, and C.
Macaya, “Skin dose and dose-area product values for interventional cardiology procedures,” Br. J. Radiol. 74(877), 48–55 (2001).
16S. Balter, “Methods for measuring fluoroscopic skin dose,” Pediatr. Radiol.
36(Suppl. 2), 136–140 (2006).
17P. B. Johnson, D. Borrego, S. Balter, K. Johnson, D. Siragusa, and W. E.
Bolch, “Skin dose mapping for fluoroscopically guided interventions,” Med.
Phys. 38(10), 5490–5499 (2011).
18Y. Khodadadegan, M. Zhang, W. Pavlicek, R. G. Paden, B. Chong, B. A.
Schueler, K. A. Fetterly, S. G. Langer, and T. Wu, “Automatic monitoring
2559
Safari et al.: Calibration of MOSkin for skin dose measurement
of localized skin dose with fluoroscopic and interventional procedures,” J.
Digital Imaging 24(4), 626–639 (2011).
19S. L. Dong, T. C. Chu, G. Y. Lan, T. H. Wu, Y. C. Lin, and J. S. Lee, “Characterization of high-sensitivity metal oxide semiconductor field effect transistor dosimeters system and LiF: Mg,Cu, P thermoluminescence dosimeters for use in diagnostic radiology,” Appl. Radiat. Isot. 57(6), 883–891
(2002).
20S. Delle Canne, A. Carosi, A. Bufacchi, T. Malatesta, R. Capperella,
R. Fragomeni, N. Adorante, S. Bianchi, and L. Begnozzi, “Use of
GAFCHROMIC XR type R films for skin-dose measurements in interventional radiology: Validation of a dosimetric procedure on a sample of
patients undergone interventional cardiology,” Phys. Med. 22(3), 105–110
(2006).
21L. D’Ercole, L. Mantovani, F. Z. Thyrion, M. Bocchiola, A. Azzaretti,
F. Di Maria, C. M. Saluzzo, P. Quaretti, G. Rodolico, P. Scagnelli, and
L. Andreucci, “A study on maximum skin dose in cerebral embolization
procedures,” AJNR, Am. J. Neuroradiol. 28(3), 503–507 (2007).
22B. A. Schueler, D. F. Kallmes, and H. J. Cloft, “3D cerebral angiography:
Radiation dose comparison with digital subtraction angiography,” AJNR,
Am. J. Neuroradiol. 26(8), 1898–1901 (2005).
23T. Moritake, Y. Matsumaru, T. Takigawa, K. Nishizawa, A. Matsumura,
and K. Tsuboi, “Dose measurement on both patients and operators during
neurointerventional procedures using photoluminescence glass dosimeters,”
AJNR, Am. J. Neuroradiol. 29(10), 1910–1917 (2008).
24M. Hayakawa, T. Moritake, F. Kataoka, T. Takigawa, Y. Koguchi, Y.
Miyamoto, K. Akahane, and Y. Matsumaru, “Direct measurement of patient’s entrance skin dose during neurointerventional procedure to avoid
further radiation-induced skin injuries,” Clin. Neurol. Neurosurg. 112(6),
530–536 (2010).
25T. Moritake, M. Hayakawa, Y. Matsumaru, T. Takigawa, Y. Koguchi, Y.
Miyamoto, Y. Mizuno, K. Chida, K. Akahane, and K. Tsuboi, “Precise
mapping system of entrance skin dose during endovascular embolization
for cerebral aneurysm,” Radiat. Meas. 46(12), 2103–2106 (2011).
26D. J. Peet and M. D. Pryor, “Evaluation of a MOSFET radiation sensor for
the measurement of entrance surface dose in diagnostic radiology,” Br. J.
Radiol. 72(858), 562–568 (1999).
27D. Glennie, B. L. Connolly, and C. Gordon, “Entrance skin dose measured
with MOSFETs in children undergoing interventional radiology procedures,” Pediatr. Radiol. 38(11), 1180–1187 (2008).
28D. D’Alessio, C. Giliberti, A. Soriani, L. Carpanese, G. Pizzi, G. Vallati, and
L. Strigari, “Dose evaluation for skin and organ in hepatocellular carcinoma
during angiographic procedure,” J. Exp. Clin. Cancer Res. 32 (2013).
29L. K. Wagner and J. J. Pollock, “Real-time portal monitoring to estimate
dose to skin of patients from high dose fluoroscopy,” Br. J. Radiol. 72(861),
846–855 (1999).
30ICRP, “The biological basis for dose limitation in the skin. A report of a
task group of committee 1 of the international commission on radiological
protection,” Ann. ICRP 22(2), 1–104 (1991).
31I. S. Kwan, A. B. Rosenfeld, Z. Y. Qi, D. Wilkinson, M. L. Lerch, D. L.
Cutajar, M. Safavi-Naeni, M. Butson, J. Bucci, and Y. Chin, “Skin dosimetry
with new MOSFET detectors,” Radiat. Meas. 43(2), 929–932 (2008).
Medical Physics, Vol. 42, No. 5, May 2015
2559
32H. Alnawaf, M. Butson, and P. K. Yu, “Measurement and effects of MOSKIN
detectors on skin dose during high energy radiotherapy treatment,” Australas. Phys. Eng. Sci. Med. 35(3), 321–328 (2012).
33Z. Y. Qi, X. W. Deng, S. M. Huang, L. Zhang, Z. C. He, X. A. Li, I. Kwan,
M. Lerch, D. Cutajar, P. Metcalfe, and A. Rosenfeld, “In vivo verification
of superficial dose for head and neck treatments using intensity-modulated
techniques,” Med. Phys. 36(1), 59–70 (2009).
34I. S. Kwan, D. Wilkinson, D. Cutajar, M. Lerch, A. Rosenfeld, A. Howie, J.
Bucci, Y. Chin, and V. L. Perevertaylo, “The effect of rectal heterogeneity
on wall dose in high dose rate brachytherapy,” Med. Phys. 36(1), 224–232
(2009).
35Z. Y. Qi, X. W. Deng, S. M. Huang, J. Lu, M. Lerch, D. Cutajar, and A.
Rosenfeld, “Verification of the plan dosimetry for high dose rate brachytherapy using metal-oxide-semiconductor field effect transistor detectors,”
Med. Phys. 34(6), 2007–2013 (2007).
36International Electrotechnical Commission, “Radiation conditions for use
in the determination of characteristics,” IEC 61267 (2005).
37National Institute of Standards and Technology, X-ray form factor, attenuation, and scattering tables, 2010, available at http://physics.nist.gov/
PhysRefData/FFast/html/form.html, accessed on 12 January 2015.
38R. Ramani, S. Russell, and P. O’Brien, “Clinical dosimetry using MOSFETs,” Int. J. Radiat. Oncol., Biol., Phys. 37(4), 959–964 (1997).
39T. Cheung, M. J. Butson, and P. K. Yu, “Effects of temperature variation on
MOSFET dosimetry,” Phys. Med. Biol. 49(13), N191–N196 (2004).
40P. P. Dendy and B. Heaton, Physics for Diagnostic Radiology (CRC, Boca
Raton, FL, 2011).
41C. Lian, J. Wong, A. Young, D. Cutajar, M. Petasecca, M. Lerch, and
A. Rosenfeld, “Measurement of multi-slice computed tomography dose
profile with the dose magnifying glass and the MOSkin radiation dosimeter,”
Radiat. Meas. 55, 51–55 (2013).
42N. Hardcastle, D. L. Cutajar, P. E. Metcalfe, M. L. Lerch, V. L. Perevertaylo,
W. A. Tome, and A. B. Rosenfeld, “In vivo real-time rectal wall dosimetry
for prostate radiotherapy,” Phys. Med. Biol. 55(13), 3859–3871 (2010).
43W. L. Jong, J. H. Wong, N. M. Ung, K. H. Ng, G. F. Ho, D. L. Cutajar, and
A. B. Rosenfeld, “Characterization of MOSkin detector for in vivo skin dose
measurement during megavoltage radiotherapy,” J. Appl. Clin. Med. Phys.
15(5), 120–132 (2014).
44T. Kron, L. Duggan, T. Smith, A. Rosenfeld, M. Butson, G. Kaplan, S.
Howlett, and K. Hyodo, “Dose response of various radiation detectors to
synchrotron radiation,” Phys. Med. Biol. 43(11), 3235–3259 (1998).
45C. Lian, M. Othman, D. Cutajar, M. Butson, S. Guatelli, and A. Rosenfeld,
“Monte Carlo study of the energy response and depth dose water equivalence
of the MOSkin radiation dosimeter at clinical kilovoltage photon energies,”
Australas. Phys. Eng. Sci. Med. 34(2), 273–279 (2011).
46M. A. S. Sherer, P. J. Visconti, and E. R. Ritenour, Radiation Protection
in Medical Radiography (Elsevier Health Sciences, Maryland Heights,
MO, 2013).
47Z. Y. Qi, X. W. Deng, S. M. Huang, A. Shiu, M. Lerch, P. Metcalfe,
A. Rosenfeld, and T. Kron, “Real-time in vivo dosimetry with MOSFET
detectors in serial tomotherapy for head and neck cancer patients,” Int. J.
Radiat. Oncol., Biol., Phys. 80(5), 1581–1588 (2011).