Dosimetric characterization of an image-guided stereotactic small animal irradiator
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PHYSICS IN MEDICINE AND BIOLOGY
Phys. Med. Biol. 56 (2011) 2585–2599
Dosimetric characterization of an image-guided
stereotactic small animal irradiator
R Pidikiti, S Stojadinovic, M Speiser, K H Song, F Hager, D Saha and
T D Solberg1
Department of Radiation Oncology, University of Texas Southwestern Medical Center, Dallas,
Received 23 November 2010, in final form 2 March 2011
Published 28 March 2011
Online at stacks.iop.org/PMB/56/2585
Small animal irradiation provides an important tool used by preclinical studies
to assess and optimize new treatment strategies such as stereotactic ablative
radiotherapy. Characterization of radiation beams that are clinically and
geometrically scaled for the small animal model is uniquely challenging
for orthovoltage energies and minute field sizes.
a commercial x-ray device (XRAD 320, Precision x-ray, Inc.)
a custom collimation system to produce 1–10 mm diameter beams and
a 50 mm reference beam. Absolute calibrations were performed using
the AAPM TG-61 methodology.Beam’s half-value layer (HVL) and
timer error were measured with an ionization chamber.
dose (PDD), output factors (OFs) and off-axis ratios were measured using
radiochromic film, a diode and a pinpoint ionization chamber at 19.76
and 24.76 cm source-to-surface distance (SSD). PDD measurements were
also compared with Monte Carlo (MC) simulations.
absolute calibrations for the reference 50 mm diameter collimator at 19.76
cm SSD were measured as 20.96 and 20.79 Gy min−1, respectively, agreeing
within 0.8%.The HVL at 250 kVp and 15 mAs was measured to be
0.45 mm Cu.The reference field PDD MC simulation results agree
with measured data within 3.5%. PDD data demonstrate typical increased
penetration with increasing field size and SSD. For collimators larger than
5 mm in diameter, OFs measured using film, an ion chamber and a diode were
within 3% agreement.
The irradiator employs
In-air and in-water
1Author to whom any correspondence should be addressed.
0031-9155/11/082585+15$33.00© 2011 Institute of Physics and Engineering in MedicinePrinted in the UK2585
2586 R Pidikiti et al
Modern radiotherapy combines advanced imaging technologies, computerized treatment
planning systems and high-energy medical accelerators to optimize radiation dose to a target
quantifiable improvements in treatment outcomes of patients treated with radiation therapy.
Moreover, the technological progress is enabling a fundamental paradigm shift in the practice
of radiation therapy. The new paradigm is a departure from small doses traditionally delivered
over many weeks, i.e. conventionally-fractionated radiotherapy (CFRT), in favor of large,
ablative doses delivered in a few fractions in 2 weeks or less. Stereotactic body radiation
therapy (SBRT), also referred to as stereotactic ablative radiotherapy (SAbR), utilizes a hypo-
fractionated dose regimen where the total prescribed dose is delivered in five fractions or
less (Blomgren et al 1995, Timmerman et al 2003). SBRT has shown the potential for a
dramatic clinical impact on the ability to achieve local control of cancer in multiple organ
sites (Timmerman et al 2003, 2007, Uematsu et al 2001, Rusthoven et al 2009b). This
success, particularly evident in early stage lung cancer patients unfit for surgery, has allowed
an absolute reduction in the rate of in-field recurrence on the order of 50% compared to CFRT
(Cheung et al 2000, Qiao et al 2003). Additionally, 80–95% local control rates for stage I
lung cancer have been reported in multiple prospective studies (Hara et al 2006, Hiraoka and
Nagata 2004, Timmerman et al 2006, Xia et al 2006). In a national prospective trial, local
control in excess of 97% was reported for patients with inoperable stage I lung cancer treated
using SBRT (Timmerman et al 2010). Similar success has been reported for SBRT of lung
and liver metastases (Herfarth et al 2001, 2004, Rusthoven et al 2009a, 2009b). These rates
rival the best surgical series.
In the field of medical and molecular imaging, the trend of swift clinical implementation
of technological innovation has been successfully followed by corresponding development of
small animal imaging systems. Every anatomical and functional imaging modality in humans
is at present commercially available for small animal use as well. Small animal imaging
systems provide new insights on cancer detection and characterization and continue to have a
noteworthy impact in the fundamental understanding of tumor kinetics, the effects of tumor
growth on the local environment and response to therapies (Vallabhajosula 2007, Bahri et al
2008, Wessels et al 2007). In contrast to sophisticated imaging systems, conformal therapy
irradiation systems for small animals are uncommon and presently under development by just
a few groups in North America. While there have been various approaches in design and
implementation, a common theme is to incorporate image guidance with conformal, small
field radiotherapy to enable preclinical studies. Notable efforts in small animal irradiator
development have been presented by research teams from Washington University School of
Medicine (Stojadinovic et al 2006, 2007, Kiehl et al 2008), Johns Hopkins University (Deng
et al 2007, Matinfar et al 2007, Matinfar et al 2009, Wong et al 2008), Stanford University
(Bazalova et al 2009, Graves et al 2007, Motomura et al 2010, Zhou et al 2010), Princess
Margaret Hospital (Lindsay et al 2008, Lindsay et al 2009, Chow and Leung 2007, Clarkson
et al 2011), University of Arkansas Medical Sciences (Moros et al 2009, Sharma et al 2009)
and University of Texas Southwestern (Cho et al 2010, Saha et al 2010, Song et al 2010).
irradiator developed to facilitate preclinical studies of SBRT techniques at the University of
Texas Southwestern Medical Center. A preclinical SBRT micro-irradiator requires significant
SBRT system. These essential capabilities include a high dose rate (∼10 Gy min−1), allowing
Small animal stereotactic irradiator 2587
Figure 1. The tungsten–copper alloy collimation system: (A) collimator inserts; (B) an insert
holder. The collimators are 25.4 mm thick with six standard collimator diameters: 1.0, 2.0, 3.5,
5.0, 7.5 and 10.0 mm. The holder itself can also be used as a 20.0 mm diameter collimator. In
addition, a reference 50.0 mm diameter collimator is used for calibration purposes and collimators
of specific shape have been manufactured for specific applications. The collimation system is
mounted into a bracket just beneath the x-ray tube.
must be capable of delivering appropriately-scaled radiation fields commensurate with small
animal anatomy. Applications of these miniaturized radiation fields necessitate robust image
guidance and sub-millimeter target localization in order to effectively deliver a potentially
ablative dose while minimizing normal tissue toxicity. Our device meets these essential
requirements, thus making SBRT technology accessible for small animals and providing
radiation delivery in a manner analogous to that employed in clinical SBRT delivery.
2. Materials and methods
The irradiator consists of three main components: (1) a small field collimation system, (2)
an image guidance system and (3) a commercial x-ray source (X-RAD 320, Precision x-ray
Inc., North Branford, CT) which provides a range of photon energies for both therapeutic and
2.1. Collimation system
The collimation assembly, shown in figure 1, consists of a disk-shaped collimator holder and
interchangeable collimators of various apertures. The collimators are 25.4 mm thick, which
of a high density (14.7 g cm−3) tungsten and copper alloy and feature six standard collimator
apertures: 1.0,2.0,3.5,5.0,7.5and10.0mmindiameter; additionally,collimatorsofaspecific
shape have been constructed for specific applications. The collimator holder itself, made of
brass (8.4 g cm−3), can be used to produce a 20.0 mm diameter beam. The collimation system
is mounted into a bracket just beneath the x-ray tube. An additional reference collimator,
made of brass with 50 mm diameter aperture, was used as the standard for absolute calibration
purposes. Absolute calibration was conducted in accordance with the recommendations of the
American Association of Physicists in Medicine (AAPM) Task Group 61 (TG-61) (Ma et al
2588R Pidikiti et al
Figure 2. The x-ray image guidance system inside the irradiator enclosure.
2.2. Image-guidance system
The image guidance system (figure 2) consists of a motorized XY linear translation stage (XY-
6060, Danaher Motion Inc., Washington, DC), a manual Z-directional stage (GDL, Parker
Company, Cleveland, OH), a motion controller (R364, Linengineering, Santa Clara, CA), a
digital imaging panel (PaxScan 1313, Varian Inc., Palo Alto, CA) and an animal platform.
1313 panel is equipped with a cesium iodide (CsI) conversion screen and amorphous silicon
(a-Si) receptor. The panel has an active area of 13 × 13 cm2and supports two pixel matrices,
1024 × 1024 (for a pixel pitch of 127 μm) and 512 × 512 (for a pixel pitch of 254 μm).
The frame rate varies depending on the binning mode, e.g. up to 10 frames per second (fps)
Small animal stereotactic irradiator 2589
for 1024 × 1024 mode and 30 fps for the 512 × 512 mode. The animal support platform
is made of a 2 mm acrylic sheet and has an attachment for isoflurane gas used for animal
anesthesia. The characteristics of the image-guidance system are described in more detail in
another article (Song et al 2010).
2.3. X-ray system
The radiation source is a commercial x-ray device (XRAD 320) that employs the MXR-321
x-ray tube (Comet AG, Liebefeld-Berne, Switzerland) to produce both imaging and treatment
ranges from 0.1 to 45 mA. This capability allows both diagnostic (30 kVp) and therapeutic
(250 kVp) configurations. These therapeutic energies provide adequate penetration for small
animal SBRT. Exposure times can be set from 0.1 to 99.9 min in 0.1 min increments. A
single rectangular focal spot of 8 mm in largest dimension allows greater dose output required
of high dose, small animal irradiation. The target consists of a tungsten anode with a 30◦
target angle; the tube has an inherent filtration of 3 mm beryllium. An integrated filter holder
allows the user to add beam-conditioning filters to achieve desired half-value layer (HVL)
specifications. To absorb additional low energy photons, the therapeutic beams described in
this manuscript contain an added 1.65 mm aluminum of beam filtration. The MXR-321 x-ray
tube is located inside the XRAD 320’s shielded enclosure. The enclosure is equipped with an
adjustable specimen shelf, a 15 × 20 cm2lead glass viewing window and a duct for accessory
cables, including a gas anesthesia tube.
2.4. Beam quality and collimator transmission
As outlined by the TG-61 protocol, beam quality specification for orthovoltage energies is
separated in two main stages. The first is to obtain the air-kerma calibration factor NKfrom a
national standards lab in terms of both tube potential and HVL. The second stage is to measure
the absorbed dose in the user’s beam. The air-kerma calibration coefficient for a PTW
N31013 ionization chamber (PTW-New York Corporation, Hicksville, NY) was obtained
for beam quality M250 from the Accredited Dosimetry Calibration Laboratory (ADCL) at
the University of Wisconsin. The HVL measurement was performed in-air using a 20 mm
diameter collimator with the calibrated chamber placed 50 cm from the collimator opening.
As the irradiator is enclosed in shielding, 50 cm was also the maximum distance possible for
narrow beam geometry. Thin sheets of copper were used as the attenuating material.
The transmission through the collimator was measured in air, at 19.76 cm SSD, using
a PinPoint N31014 ionization chamber (PTW-New York Corporation, Hicksville, NY) by
comparing the central axis readings for a reference 50 mm diameter collimator to readings of
a totally blocked field.
2.5. Irradiator calibration
The absolute dose calibration of the irradiator was performed in accordance with the
were performed at a 19.76 cm source-to-surface distance (SSD) for the reference 50 mm
diameter collimator. The ADCL calibrated PTW N31013 ionization chamber was used to
measure the (percent) depth dose in a water phantom, as the in-water calibration requires the
knowledge of the PDD value at reference depth of 2 cm. The PDD data points were measured
for 30 s exposure time at depths of 0–150 mm, using 1 mm increments for the first 10 mm and
then 5 mm steps for the remaining depths.
2590 R Pidikiti et al
An independent output verification of the irradiator was done using lithium fluoride TLDs
provided by the Radiation Dosimetry Services (MD Anderson Cancer Center, Huston, TX).
TLD measurements were done in air, that is, on top of a 3 mm thick plastic platform, at the
reference SSD of 19.76 cm. The one requirement was not to exceed 4 Gy since the TLDs’
response to radiation would not be in the linear region to any further extent.
2.6. Beam characterization
A Monte Carlo (MC) model of the irradiator beam was generated and subsequently validated
using larger field PDD measurements. The MXR-321 beam spectrum for 250 kVp
was calculated using the SpekCalc program described by Poludniowski and Evans (2007,
Poludniowski 2007). The resulting energy spectrum, the custom beam collimation system
and experimental geometries were all subsequently modeled using the MCNPX program (Los
Alamos National Laboratories, NM). All simulations were conducted to achieve estimated
statistical errors of less than 2%.
The Hurter–Driffield (HD) curve, i.e. the sensitometric orthovoltage calibration curve,
was established by exposing radiochromic films (Gafchromic EBT2, International Specialty
Products, Wayne, NJ) using the reference collimator to a known dose at the surface of
a 6 cm thick solid water block. In turn, the orthovoltage HD calibration curve for all
collimators was used to convert the measured optical densities to dose, i.e. to establish
the PDD behavior relative to field size. Generally an exposure time of 30 s was used for
Gafchromic EBT2 films placed between thin solid water slabs (Gammex RMI, Middleton,
WI) at 19.76 and 24.76 cm SSD. The irradiated area was at least 5 mm from any film edge
to avoid optical density measurement artifacts observed near film edges (Butson et al 1998).
The films were irradiated simultaneously perpendicular to beam’s direction. Starting from
top to bottom, the solid water phantom consisted of eight 1 mm slabs, nine 2 mm slabs, one
3 mm slab, four 5 mm slabs and six 10 mm slabs. This setup resulted in a total of 29 films.
The measurement depths were corrected for the average film thickness of 0.29 mm.
using an Epson 1000 XL scanner (Epson America Inc., Long Beach, CA) in 48 bit red–green–
blue (RGB) mode, 16 bits per color, at 150 dpi resolution and saved as tagged image file
format (tiff) files. The scanned films were imported into and analyzed with a custom written
MatLab R2010a (MathWorks, Natick, MA) code. Only the red channel data were used in the
analysis. The regions of interest selected at the central portion of the exposed films ranged
from 20 × 20 pixels for 10 mm diameter collimator to 4 × 4 pixels for the smallest 1 mm
diameter collimator. The net optical density was computed as
?Iexp(D) − Ibckg
where Iexp(D) and Iunexpare the readings for the exposed and unexposed films, respectively,
while Ibckgis the reading for an opaque film, that is, the zero light transmitted intensity value.
Standard deviations in the net optical density were computed following Devic et al (2005):
The same experimental setup used for obtaining the PDD curves was used to obtain beam
profiles and off-axis ratios (OARs) and to characterize the penumbra. Beam profiles were
measured in a solid water phantom at an SSD of 19.76 cm and at a depth of 1.46 cm. The
penumbra, defined as the distance between the 80% and 20% isodose lines, is an important
netOD(D) = −log10
Small animal stereotactic irradiator2591
part for treatment planning considerations, especially for stereotactic situations when beams
are directed from several different directions to optimize treatment.
Theoutputfactors(OFs), i.e.totalscatterfactors, weremeasuredusingGafchromicEBT2
collimator were determined by converting the measured optical densities to dose utilizing the
orthovoltage calibration curve. For ‘larger’ collimator diameters, i.e. >5 mm, the OFs were
also measured in a water phantom using a PinPoint N31014 ionization chamber and an SFD
stereotactic diode (IBA Dosimetry America, Bartlett, TN).
accounted for by the machine timer mechanism during the x-ray delivery, i.e. the end effect,
and by the nonlinearity of beam intensity as the function of the tube current at constant tube
voltage. The end effect was assessed in a series of exposures in a water phantom by measuring
the charge collected by a PTW N31013 ionization chamber in the increments of 6 to 120 s.
The 6 s increment is the shortest possible irradiator time increment. The data were analyzed
using the graphical extrapolation method for the ion chamber readings as a function of timer
settings. The timer error is the zero exposure represented by the intercept of the regression
line on the time axis.
To determine beam intensity as a function of tube current, radiochromic films were
exposed at the surface of a solid water phantom using the reference collimator. These films
were exposed at 250 kVp tube voltage for 24 s and varying tube currents from 1 to 15 mAs in
1 mAs increments.
3.1. Beam quality and collimator transmission
The HVL measurement was performed in-air utilizing a 20 mm diameter collimator with a
PTW N31013 ionization chamber placed 50 cm from the collimator opening. The resulting
HVL, at a tube potential of 250 kVp and tube current of 15 mAs, was measured to be
0.45 mm Cu.
The transmission through the collimator, measured in air using a PinPoint N31014
ionization chamber, was 0.7%.
3.2. Irradiator calibration
output of the system was also independently verified; the ratio of stated versus measured dose
was1.01, usinglithiumfluorideTLDmeasurements readbytheRadiationDosimetryServices
(MD Anderson Cancer Center, Huston, Texas). The in-air method resulted in a dose rate of
20.96 Gy min−1, while the in-water calibration measurement at 2 cm depth calculated at the
surface yielded a dose rate of 20.79 Gy min−1, i.e. an agreement within 0.8%. Noticeably,
the agreement between in-air and in-water calibrations is very dependent on an accurate PDD
measurement. The PDD curve for the reference collimator was determined using the ADCL
calibrated PTW N31013 ionization chamber in a water phantom.
3.3. Beam characterization
the reference collimator to the corresponding measured PDDs using a calibrated ion chamber
2592 R Pidikiti et al
Figure 3. The percent depth dose (PDD) for the reference 50 mm diameter collimator measured
using an ion chamber (IC) and an EBT2 Gafchromic film compared to Monte Carlo (MC)
simulation. The average percent agreement calculated at 1 mm depth increments was 1.0%
with largest outliers at 3.5% difference.
in a water phantom and a film in a solid water phantom. The PDDs from the various methods
agree remarkably well. The average percent agreement calculated at 1 mm depth increments
was 1.0% with largest outliers at 3.5% difference (figure 3).
The sensitometric orthovoltage calibration curve, established by utilizing the reference
collimator to expose EBT2 films to a known dose at the surface of a 6 cm thick solid water
block is shown in figure 4. The error bars in figure 4 correspond to the uncertainty associated
with the film scans and represent two standard deviations or 95.4% confidence level.
Depth dose characteristics, for example, 52.3% and 23.6% at depths of 22 and
50 mm, respectively, for the 5 mm diameter collimator at 19.76 cm SSD, provide more than
adequate penetrations for targeting deep-seated targets in small animals, particularly when
using multiple beam directions. As expected for orthovoltage energies, the maximum dose
point is at the surface. The PDD data show the familiar relationship of increasing penetration
with increasing field size as shown in figure 5. Likewise, the increase in SSD produced an
increase in PDD, apparent in the comparison of the PDDs at 19.76 cm SSD in figure 5 with
those at 24.76 cm shown in figure 6.
The total scatter factors for all collimators were measured with Gafchromic EBT2 film at
the surface of a solid water slab and compared to SFD stereotactic diode and PinPoint N31014
ionization chamber measurements at the surface in a water phantom. The results, shown in
table 1, are within 3% agreement for collimators greater than 5 mm in diameter. For smaller
collimators, however, the detector size excluding the film is comparable to the collimator
diameter, rendering the measurements unreliable for these field sizes.
OARs were measured using Gafchromic EBT2 film in solid water. Measured penumbra
are shown for a depth of 1.46 cm as a function of collimator diameter for 19.76 cm and
24.76 cm SSD in table 2. Corresponding OARs at 1.46 cm depth are shown in figures 7
Small animal stereotactic irradiator 2593
Figure 4. The sensitometric orthovoltage calibration curve; the error bars show 95.4% confidence
Figure 5. The percent depth dose at 19.76 cm SSD shown for 1 to 10 mm collimators measured
with a Gafchromic EBT2 film.
and 8. The intrinsic magnification factor of 1.44 for this setup is the ratio between source-
to-detector distance (SDD = SSD + depth = 21.22 cm) and the source-to-collimator distance
(SCD = 14.76 cm). The profiles in figures 7 and 8 reflect the divergence of the beam from
a large asymmetric focal spot which creates a beam that is slightly elliptical in shape; thus,
2594R Pidikiti et al
Figure 6. The percent depth dose at 24.76 cm SSD shown for 1–10 mm collimators measured
with a Gafchromic EBT2 film.
a Gafchromic EBT2 film and compared to an SFD stereotactic diode and a PinPoint N31014
The total scatter factors relative to the reference 50 mm diameter measured with
Total scatter factors
SFD Stereotactic Diode
PinPoint N31014 IC
in-plane and cross-plane profiles differ slightly from each other. The elliptical shape of the
beam in orthogonal planes is not the artifact of the beam concentricity, i.e. the manufacturer’s
specification of ±0.5 mm for the focal spot alignment with the central axis of the applicator
Asageneralnote, theXRAD320irradiatordoesnothaveaflatteningfilter. Consequently
the beams produced are not intended to be flat. This is very similar to the beams produced
by units such as a CyberKnife, Gamma Knife and dedicated stereotactic radiosurgery and/or
flattening-filter free linacs. Moreover, our largest field size is 10 mm in diameter, for which
typical flatness and symmetry specifications do not apply. For treatment planning purpose
such beams are scanned and modeled as measured.
Small animal stereotactic irradiator2595
Figure 7. Off-axis ratio as a function of the distance along the x-axis.
Table 2. The 80−20% beam penumbra along the in-plane axis and the cross-plane axis at 1.46 cm
depth for 19.76 and 24.76 cm SSD.
Penumbra (mm) at 19.76 cm SSD
In-plane, y axisCollimator diameter (mm) Cross-plane, x axis
Penumbra (mm) at 24.76 cm SSD
As a result of the time delay required to switch the beam ‘ON’ and ‘OFF’ in an x-ray
unit, the timer typically does not accurately indicate the exposure time. Hence, it needs to be
corrected by a small increment known as the timer error or the end effect. Using the graphical
extrapolation method, the end effect was determined to be 3 s as represented by the intercept
of the regression line on the time axis in figure 9.
The beam intensity as a function of tube current at constant tube voltage produced an
anticipated linear relationship as shown in figure 10. Consequently, doubling the current at
2596R Pidikiti et al
Figure 8. Off-axis ratio as a function of the distance along the y-axis
Figure 9. The end effect of x-ray unit was determined graphically as the intercept of the regression
line on the time axis.
constant exposure time has the same effect as doubling the exposure time at constant tube
current. A high dose rate of over 20 Gy min−1and minimum timer setting of 6 s would enable
only relatively coarse units of dose that can be delivered during experimental procedures. The
ability to change mAs is essential to obtain finer dose increments.
Small animal stereotactic irradiator 2597
Figure 10. The linear relationship between x-ray beam intensity and the tube current.
Animal irradiation systems facilitate scientific testing of biomedical hypotheses in vitro (cell)
novel protocols for human cancer treatments. A small animal irradiator system is an essential
element for quantitative molecular-imaging studies and radiobiological experiments of tumor
and healthy tissue response to radiation. We have developed and characterized an image-
guided SBRT irradiator for small animals. The irradiator mimics the clinical application of
SBRT providing high dose rate, sharp beam profiles and more than adequate beam penetration
depths for targeting deep-seated tumors in small animal models. We have characterized
and leakage. These results have been cross-compared either with a benchmarked Monte Carlo
model or additional measurement techniques. These measured parameters are the foundation
for simple treatment planning calculations and advanced beam modeling. The irradiator has
already become an essential tool for scientists at UT Southwestern in conducting a broad
spectrum of preclinical studies.
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