Beam generation and planar imaging at energies below 2.40 MeV with carbon and aluminum linear accelerator targets

Beam generation and planar imaging at energies below 2.40 MeV with
carbon and aluminum linear accelerator targets
David Parsons
Department of Physics and Atmospheric Science, Dalhousie University, 5820 University Avenue, Halifax,
Nova Scotia, B3H 1V7, Canada
James L. Robar
Department of Radiation Oncology, Dalhousie University, 5820 University Avenue, Halifax, Nova Scotia,
B3H 1V7, Canada
(Received 20 February 2012; revised 6 June 2012; accepted for publication 7 June 2012; published 5
July 2012)
Purpose: Recent work has demonstrated improvement of image quality with low-Z linear accelerator
targets and energies as low as 3.5 MV. In this paper, the authors lower the incident electron beam
energy between 1.90 and 2.35 MeV and assess the improvement of megavoltage planar image quality
with the use of carbon and aluminum linear accelerator targets.
Methods: The bending magnet shunt current was adjusted in a Varian linear accelerator to allow
selection of mean electron energy between 1.90 and 2.35 MeV. Linac set points were altered to
increase beam current to allow experimental imaging in a practical time frame. Electron energy was
determined through comparison of measured and Monte Carlo modeled depth dose curves. Planar
image CNR and spatial resolution measurements were performed to quantify the improvement of
image quality. Magnitudes of improvement are explained with reference to Monte Carlo generated
energy spectra.
Results: After modifications to the linac, beam current was increased by a factor greater than
four and incident electron energy was determined to have an adjustable range from 1.90 MeV to
2.35 MeV. CNR of cortical bone was increased by a factor ranging from 6.2 to 7.4 and 3.7 to 4.3
for thin and thick phantoms, respectively, compared to a 6 MV therapeutic beam for both aluminum
and carbon targets. Spatial resolution was degraded slightly, with a relative change of 3% and 10% at
0.20 lp/mm and 0.40 lp/mm, respectively, when reducing energy from 2.35 to 1.90 MV. The percent-
age of diagnostic x-rays for the beams examined here, ranges from 46% to 54%.
Conclusion: It is possible to produce a large fraction of diagnostic energy x-rays by lowering the
beam energy below 2.35 MV. By lowering the beam energy to 1.90 MV or 2.35 MV, CNR improves
by factors ranging from 3.7 to 7.4 compared to a 6 MV therapy beam, with only a slight degradation of
spatial resolution when lowering the energy from 2.35 MV to 1.90 MV. © 2012 American Association
of Physicists in Medicine. [http://dx.doi.org/10.1118/1.4730503]
Key words: low-Z target, carbon, aluminum, planar imaging
I. INTRODUCTION
In the past, megavoltage portal imaging has been used to
align a patient’s position for radiation therapy. These images
have inherently poor contrast due primarily to the Compton
dominant interaction of photons in the megavoltage energy
range. The introduction of a diagnostic quality x-ray tube
and detector, orthogonal to the therapy beam line1
has
largely replaced portal imaging as a means for image guided
radiation therapy. While the use of this technology for planar
imaging and cone beam computed tomography (kV CBCT)
has greatly improved available image quality, it also involves
greater cost, increased quality assurance, and does not allow
for beam’s eye viewing.
Several groups have shown that the removal of the high
atomic number (Z) target and flattening filter from the therapy
beam line and the introduction of a low-Z target can greatly
enhance image quality compared to standard megavoltage
therapy beams.2–6
Most recently, Faddegon et al.7
used a
1.32 cm thick carbon target placed in the high-energy electron
scattering port of a high-energy linac with 4.2 MeV electrons
incident on the target. This work showed that soft tissue CNR
was improved up to a factor of three in CBCT images com-
pared to a 7.0 MV treatment beam line. Similarly, Roberts
et al.8
used a 2 cm thick carbon target (ρ = 1.89 g/cm3
)
placed in the high-energy collimator port of high-energy
linac operating in 4 MeV electron mode with the primary
and secondary scatter foils removed. This work showed that
the megavoltage planar image contrast of dense bone in
water was improved by a factor of 4.62 for thin phantoms
(5.8 cm thick) and 1.3 for thicker phantoms (25.8 cm).
Spatial resolution was improved when compared to the 6 MV
therapy beam. Orton and Robar9
modeled 7.0 MeV electrons
incident on an aluminum and beryllium low-Z targets in a
high-energy linac, with a QC3 phantom and 25-layer aS500
detector, demonstrating a contrast enhancement factor rang-
ing from 1.6 to 2.8. Sawkey et al.10
showed a 25% increase in
contrast when decreasing the electron energy from 6 MeV to
4 MeV on a diamond target in a high-energy linac, with no
4568 Med. Phys. 39 (7), July 2012 © 2012 Am. Assoc. Phys. Med. 45680094-2405/2012/39(7)/4568/11/$30.00

4569 D. Parsons and J. L. Robar: Beam generation below 2.40 MeV with low-z targets 4569
appreciable change in spatial resolution. Similarly Robar
et al.11
showed CNR increases by factors ranging from of 1.2
to 1.7 for 7.0 MeV incident on an Al target compared to a
6 MV therapeutic beam, and when incident electron energy
was decreased to 3.5 MeV, this factor increased, ranging from
2.7 to 3.8. Connell and Robar12
showed that spatial resolution
is improved with certain combinations of low-Z material
and thickness compared to a 6 MV therapy beam. A slight
decrease in spatial resolution was observed when decreasing
the incident electron from 7.0 to 4.5 MeV (Ref. 12).
It is clear that by lowering the incident electron energy in
combination with a low-Z target, for a given imaging dose,
CNR can be greatly enhanced. To date, the lowest energy ex-
amined with a low-Z target beam has been 3.5 MeV (Ref. 11);
here we address the issues of (i) the amount by which the inci-
dent electron energy can be lowered in a standard high energy
linac, and (ii) the resultant gains in planar image CNR. We
demonstrate practical limitations with regard to the reduction
of beam current with lowered energy,13
and report on depth
dose, imaging CNR, and energy spectral characteristics, for
beams generated with carbon and aluminum targets, with in-
cident electron beams with energies below 2.4 MeV.
II. MATERIALS AND METHODS
II.A. Beam generation and low-Z targets
The 6 MV and low-Z target beams were generated us-
ing a Varian 2100EX linear accelerator (Varian Medical, Inc.,
Palo Alto, CA). Two separate low-Z targets were investigated
(i) 6.7 mm thick aluminum (ρ = 2.69 g/cm3
) and (ii) 7.6 mm
thick carbon (ρ = 1.88 g/cm3
). The aluminum target was orig-
inally designed and used by Robar et al.11
previously for a
3.5 MV beam. The thickness of the carbon target was calcu-
lated to coincide with the continuous slowing down approx-
imation range for a 2.5 MeV electron. While an increase in
low-Z target thickness slightly degrades resolution,12
a “full-
thickness” target alleviates the need for a polystyrene filter to
absorb transmitted electrons and increases the production of
lower energy photons.2
As described previously,9,11,12
when
used experimentally, the carousel is operated in manual mode
and the appropriate target is rotated into the beam line via
a rotary switch. The linac is operated in electron mode and
equipped with a dedicated 4 MeV program board used for
research purposes only. The bending magnet was adjusted
to tune the beam energy down from the nominal operating
point of 4.5 MeV to the lower energies used. Complete de-
sign and installation of similar targets used in this study has
been published recently9,12
and will not be explained in full
here. Briefly, targets were mounted to the beam side of the
carousel and vertically offset 0.9 cm from the beryllium exit
window of the primary collimation vacuum. By keeping the
target as close as possible to the beryllium exit window, the
amount of electron scattering in air is minimized.8,9,12
II.B. Beam tuning to maximize current
Adjustment of the bending magnet current can be used to
select the mean energy of electrons exiting the bending mag-
net, with electrons outside the mean energy being stopped
within the linac head. Mean energy selection is accomplished
by an energy slit within the bending magnet that allows elec-
tron energy selection within 3% of the nominal energy. When
using this technique, the number of electrons within the se-
lected energy window decreases when the selected energy dif-
fers from the nominal energy for the selected mode. This is
due to the irregular electron spectra exiting the waveguide.
Although the electron energy spectrum exiting the waveguide
is often approximated as Gaussian,14–16
in fact it can be highly
asymmetric,14,15
with a sharp decrease in electron population
away from the nominal energy. An example of this can be seen
in the work by Wei et al.,14
which gives electron spectra re-
constructed from depth dose curves. The results demonstrate
that low-energy electrons, e.g., ranging from 1 to 3 MeV, are
present in a 6 MeV spectrum; however, the relative fluence
of these electrons is approximately one hundredth of that at
the nominal energy. As a result, any lowering of the incident
electron energy results in a considerably lower beam current
exiting the bending magnet and, for our application, signifi-
cantly protracted image acquisition times. A solution, without
significant alteration to the waveguide, is to increase the total
current exiting the waveguide. To observe slight changes in
beam current, a 4 MeV tantalum electron scattering foil was
inserted into the beam line via a rotatory switch controlling
the carousel. The linac was operated in 4 MeV electron mode
at a MU rate of 1000 MU/min with the dose servos turned off.
In maximizing beam current, we optimized the combination
of two parameters: gun high voltage (HV) and grid voltage,
over the full ranges of each parameter recommended by the
manufacturer for same operation. Course setting of both pa-
rameters is done inside the gun deck of the linear accelerator,
while the grid voltage may also be fine-tuned from outside the
room with the beam on, using the “gun-I” potentiometer. The
process of maximizing beam current thus involved (i) setting
the first gun HV value in the range, (ii) setting the first grid
voltage coarse value on the gun deck, (iii) fine-tuning the grid
voltage with the beam on, and (iv) recording the maximum
achievable dose rate. For the same gun HV setting, the next
coarse grid voltage value was then set on the gun deck, and
fine-tuned. After exploring the full allowable range of grid
voltage, the next gun HV setting was chosen on the gun deck.
This was repeated to produce a two dimensional matrix of
measured dose rate values, from which the optimal combina-
tion of the two parameters was chosen. Finally, for this com-
bination, the solenoid and buncher cavity steering set points
were tuned on the program board to further tune beam current.
We reiterate that, to avoid potential damage, the examined
range did not exceed the manufactures recommendations. Fi-
nally, once the final tuning was complete, the scattering foil
was replaced by either aluminum or carbon targets for x-ray
production.
II.C. Electron energy determination
Over the low electron energy range produced here, the re-
lationship between bending magnet and electron energy was
not known precisely based on manufacturer specifications.
Medical Physics, Vol. 39, No. 7, July 2012

To determine incident electron energy, the 4 MeV tantalum
electron scatting foil described previously was rotated into the
beam line. Percent depth dose (PDD) curves were measured
in a 50 × 50 × 50 cm3
scanning water tank (Scandtronix,
Uppsala, Sweden) with a p-type silicon electron field diode
(EFD, Scandtronix, Uppsala, Sweden) at a source to surface
distance (SSD) of 100 cm and with a 10 × 10 cm2
electron ap-
plicator. Measurements were taken at 0.2 mm intervals. These
curves were then matched to Monte Carlo PDD curves of
known electron energy. BEAMnrc (Ref. 17) and DOSXYZnrc
(Ref. 18) were used to simulate the linac and water phantom.
All simulations were run in accordance with the physical
setup using validated models provided by Orton and Robar.9
Parallel beams of mono-energetic electrons were used with
a spot size of 0.1 cm. Global electron (ECUT) and photon
(PCUT) cut-off energies of 0.521 MeV and of 0.010 MeV,
respectively, were used. The incident electron energy was
varied in 0.025 MeV increments between 1.700 MeV and
2.500 MeV and plotted against the measured data in MATLAB
(Mathworks, Natick, MA).
II.D. Photon measurements
Given that low-Z target photon beams have not been pro-
duced previously using the low electron energies considered
here, it was of interest to examine depth dose characteristics.
Photon depth dose measurements were acquired using a 50
× 50 × 50 cm3
water tank with 0.015 cm3
(PTW N31014,
Freiburg, Germany) and 0.125 cm3
(PTW N31010, Freiburg,
Germany) cylindrical ion chambers. For all measurements,
an SSD of 100 cm and a field size of 10 × 10 cm2
were used.
The 0.015 cm3
ion chamber was used for measurements
shallower than 1.0 cm below the surface to provide resolution
in the buildup region, while the 0.125 cm3
ion chamber was
used for improved signal-to-noise characteristics at greater
depths. These two sets of data were concatenated at a depth
of 1.0 cm and the resultant curve was normalized to the max-
imum dose. Photon depth dose measurements were acquired
for both carbon and aluminum targets, for the minimum and
maximum energies examined in this investigation, which,
using the method described above, were determined to be
1.90 MV and 2.35 MV.
II.E. Photon Monte Carlo simulations
In order to study changes in photon fluence with energy
and target material, BEAMnrc (Ref. 17) was used to simulate
the linac beam-line. The linac was modeled in accordance
with the physical setup using validated models provided by
Orton and Robar.9
1.1 × 108
incident electron histories were
run for the aluminum and carbon targets, with energies of 1.90
and 2.35 MeV. Linac collimation was set to a field size of 10
× 10 cm2
(specified at isocenter). Selective bremsstrahlung
splitting was used, and as recommended by Rogers et al.,19
an effective splitting field size of 20 × 20 cm2
was used at an
SSD of 100 cm with minimum and maximum bremsstrahlung
splitting numbers of 100 and 1000, respectively. PCUT and
ECUT were set to 0.010 MeV and 0.700 MeV, respectively.
For confirmation of the modeled spectra, percent depth doses
were run for the aluminum target at 1.90 MV and 2.35 MV
in DOSXYZnrc (Ref. 18) and compared to corresponding
measured results. To determine the spectral distributions,
BEAMdp (Ref. 20) was used to analyze the phase space data.
To determine the source of the photons and their resultant
spectral contribution, the LATCH feature was utilized to
identify the origin of generated photons. In addition to sim-
ulating the two homogeneous carbon or aluminum targets, a
carbon target with various thickness of copper (on the exit
side) was modeled to examine the requirements for filtration
of very low energy (e.g., <25 keV) photons that would
contribute to superficial patient dose but not image formation.
II.F. Imaging system
An aS1000/IAS3 imaging system (Varian Medical Sys-
tems, Inc.) was used for all planar imaging. The system was
set up as a stand-alone configuration separate from the clini-
cal imaging system on the treatment unit. The aS1000 panel
has an active area of 30 × 40 cm2
and includes a 1.0 mm Cu
buildup plate, a Gd2O2S:Tb scintillating phosphor layer and a
1024 × 768 array of photodiodes switched by thin-film tran-
sistors deposited on a glass substrate. The setup for imaging
was as follows: The gantry and couch were rotated to 90◦
. The
imaging panel was placed in a stand located on the couch with
fine adjustments screws and leveled to ensure orthogonality
of the panel and the beam axis. For all imaging, a source-to-
detector distance (SDD) of 140 cm was used. The number of
monitor units (MU) per exposure was varied to set the dose
to the phantom, where the IAS3 allows cMU control of this
parameter. Prior to imaging, flood field and dark field cali-
brations were performed where the former has the effect of
correcting for the forward peaked nature of the low-Z target,
flattening filter free beam.
II.G. Image quality: Contrast to noise ratio
To observe changes in CNR with target material and
energy, a contrast phantom was constructed for planar
imaging. The phantom, shown in Fig. 1, consisted of an
18 × 18 × 3 cm3
block of polystyrene with six circular holes
of 2.5 cm diameter for tissue equivalent materials (Gammex,
Middleton, WI). To measure dose, a channel was added along
the phantom midline to accommodate a 0.125 cm3
cylindrical
ion chamber (PTW N31010, Freiburg, Germany). The cham-
ber was calibrated for the low-Z target beams by measuring
the dose per MU in solid water (Gammex, Middleton, WI)
with thermoluminescent dosimeters (TLDs) (TLD-800,
Saint-Gobain Crystals, Hiram, OH) at an SSD of 95 cm
depth of 5 cm with a 10 × 10 cm2
(defined at isocenter).
Manganese doped lithium tetraborate (Li2B4O3:Mn) TLDs
were used due to their minimal energy dependence21
(0.9 at
30 keV/60
Co). The TLDs were then removed and replaced
by the ionization chamber in the same location, to obtain a
conversion from collected charge to cGy. This was done for
the 2.35 MV carbon and aluminum target beams. During
imaging the phantom was oriented as shown in Fig. 1, such

FIG. 1. Left figure shows a schematic drawing of the contrast phantom. Shown are tissue equivalent inserts with electron densities relative to water. Right figure
shows the experimental setup for planar imaging.
that the circular cross sections of material regions faced the
beam and the phantom was centered on the linac isocenter.
Images were acquired with dose values at the phantom center
ranging from approximately 0.01 to 0.16 cGy. This low dose
range was chosen to match a typical anticipated dose per
projection used in CBCT, given that this is a likely application
of the low-Z target beam.11
CNR was calculated as
CNR =
|Pmaterial − Ppolystyrene|
σ2
material + σ2
polystyrene
, (1)
where Pmaterial is the average pixel value within the insert,
Ppolystyrene is the average value of the surrounding polystyrene,
σmaterial is the average noise within the material and σpolystyrene
is the average noise in the surrounding polystyrene. Pixel
values in the surrounding polystyrene were measured within
two concentric circles with radii of 1.8 and 2.1 cm (measured
at isocenter). Error bars were found by calculating the mean
and standard deviation of CNR measured in multiple images.
While the 3 cm thick phantom in Fig. 1 allows sensitivity
in measuring changes of CNR with beam parameters, it is
not representative of typical patient separation. Thus, the
experiment was repeated with the phantom padded by 5 cm
of solid water on the source side and 7 cm on the detector
side, for a total phantom thickness of 15 cm.
II.H. Image quality: Spatial resolution
Images of the QC3 phantom22
were taken at photon ener-
gies of 1.90, 2.15, and 2.35 MV/carbon and a 6 MV thera-
peutic beam. These images were analyzed according to the
method outlined by Rajapakshe et al.,22
where the relative
modulation transfer function (RMTF) is defined as
RMTF =
M(f )
M(f1)
, (2)
where M( f ) is the output modulation of the line pair in a re-
gion of interest and M( f1) is the output modulation for the
lowest frequency line pair region. As suggested by Droege
and Morin,23
the output modulation is obtained by using the
relationship between the signal amplitude and its variance is
given by
M2
(f ) = σ2
m(f ) − σ2
(f ), (3)
where σm
2
( f ) is the measured total variance within a region of
interest and σ2
( f ) is the variance due to random noise within
the region of interest. The variance due to random noise as
calculated by Rajapakshe et al.22
is given by
σ2
(f ) =
σ2
sub
2
, (4)
where σ2
( f ) represents the variance within the region of in-
terest of two subtracted images. Error bars were found by cal-
culating the mean and standard deviation of RMTF measured
in multiple image sets.
II.I. Qualitative planar imaging
Planar imaging of a sheep head was done to observe the
qualitative effects on imaging with changes in target compo-
sition and energy. For these images the CNR phantom was
replaced with an adult sheep head, all other setup parameters
were unchanged. An approximate imaging dose of 0.14 cGy
was delivered to the center of the head, located on isocen-
ter. Images were acquired at 2.35 MV with both carbon and
aluminum target beams and a 6 MV therapeutic beam. The re-
sultant images were analyzed in MATLAB and were compared
with identical gray level window settings.
III. RESULTS AND DISCUSSION
III.A. Beam current maximization
After adjusting the bending magnet shunt current only
with the scattering foil in place, the electron MU rate was
approximately 600 MU/min at 2.35 MeV. As shown in
Fig. 2, upon altering the voltage of the electron gun and grid
to the optimal setting the MU rate increased to approximately

FIG. 2. MU rate in relation to electron energy for both electrons and photons. The upper plot shows observed MU rate with tantalum scattering foil in place.
The lower plot shows observed MU rate with aluminum and carbon targets in place. Error bars are generated from the observed variation in MU rates.
2800 MU/min. The MU rate gradually decreases as the inci-
dent electron energy is decreased to 1.90 MeV, below which a
sharp drop was observed from approximately 1400 MU/min
to 20 MU/min. Due to this, it does not appear possible to oper-
ate at a lower energy with this current setup. With the carbon
and aluminum targets in place, a MU rate was observed rang-
ing from 5–9 MU/min and 3–6 MU/min, respectively, before
adjustments this rate was 0–1 MU/min. With these improve-
ments in beam current, it is now possible to acquire multiple
low-Z target images per second at doses typical of CBCT pro-
jections. It should be noted that the monitor chamber was not
calibrated and that these are only relative MU rate measure-
ments. However, with the carbon target in place, a dose rate of
0.7 cGy/min was measured for a 10 × 10 cm2
ﬁeld at a depth
of 5 cm and an SSD of 95 cm was measured with a 0.125 cm3
cylindrical ion chamber (PTW N31010, Freiburg, Germany).
III.B. Incident electron energy
Figure 3 shows the experimentally measured and Monte
Carlo modeled electron PDD curves for energies ranging
from 1.90 MeV to 2.35 MeV. The measured and modeled
data show excellent agreement in the dose build-up and
fall-off region of the PDD, with a slight variation where
the central axis depth dose curve meets the bremsstrahlung
background. The depth of the 90% PDD (R90) ranges from
2.9 mm to 4.3 mm for 1.90 MeV and 2.35 MeV, respectively.
Beyond the buildup region, the dose quickly decreases with
R50 ranging from 4.2 mm to 6.1 mm for 1.90 MeV and
2.35 MeV, respectively, and the maximum range is less than
9.0 mm for the 2.35 MeV electron beam. To our knowledge,
these are the lowest energy beams used in conjunction with a
low-Z target produced with a megavoltage Clinac.8,10,11
FIG. 3. Experimental (curves) and Monte Carlo (points) electron percent depth dose curves ranging from 1.90 MeV to 2.35 MeV.

FIG. 4. Percent depth doses for both aluminum and carbon targets are shown as well as corresponding Monte Carlo modeled depth dose curves for the aluminum
target.
III.C. Photon beam characteristics
Figure 4 shows measured percent depth dose curves nor-
malized to maximum dose for 1.90 and 2.35 MeV electron
beams incident on the aluminum and carbon targets. Also
shown are modeled depth dose curves for the aluminum tar-
get. Because of the low dose rates produced, relatively few
depths were measured compared to the number of modeled
depths. As expected, as energy was lowered the dose at a
given depth was decreased. At 2.35 MV, we observe dmax at
0.6 ± 0.1 cm and 0.4 ± 0.1 cm for the aluminum and carbon
targets, respectively. No distinguishable shift in dmax toward
the surface is observed when decreasing the incident electron
energy from 2.35 MV to 1.90 MV. For comparison, Faddegon
et al.7
and Sawkey et al.10
reported a dmax between 0.9 cm and
1.1 cm for 4.2 MV/carbon beams, respectively. Compared to
the 4.2 MV/carbon beam generated by Faddegon et al.,7
the
2.35 MV/carbon beam is significantly less penetrating with
a change in dmax of approximately 0.5 cm, toward the surface
and a decrease in the percent dose at 10 cm of 25%. Compared
to the 4 MV/aluminum investigated by Orton and Robar,9
the
percentage of photons within the diagnostic energy range was
enhanced by approximately 9% and 13% when decreasing the
incident electron energy to 2.35 and 1.90 MeV, respectively,
on an aluminum target. At a depth of 10 cm, a decrease in
percent dose of 20% and 22% was observed when changing
the target from aluminum to carbon at 2.35 MV and 1.90 MV,
respectively. These observations indicate that the aluminum
target produces a harder beam compared to the carbon target.
This is confirmed when examining the spectral distributions
shown in Fig. 5, which shows that the mode of the spectrum
is shifted from 20 keV to 50 keV when replacing aluminum
with carbon, and significantly more photons exist below
40 keV (25% for carbon compared to 6% with aluminum).
This is caused by the reduced hardening of the beam through
photoelectric absorption by the lower-Z target material.
Figure 6 illustrates the sources of photon generation within
the linac. Approximately 12%–17% of photons are generated
within the beryllium exit window of the primary collimator,
with 87%–83% being generated in the aluminum and car-
bon targets, respectively, at 2.35 MeV. The amount of photons
generated in the exit window was relatively minor compared
to that report by Roberts et al.,8
in which approximately 71%
of the beam was generated in the nickel exit window and 28%
in the carbon target.
The considerable population of very low energy photons
from the carbon target may increase patient dose without con-
tributing to image formation. Figure 6 shows 2.35 MV/carbon
spectra for various thicknesses of copper on the exit side of
the target. The percentage of photons below 25 keV can be
reduced by half with as little as 25 μm of copper.
III.D. Quantitative image quality
Figure 7 shows planar CNR images of the cortical bone
insert with 6 MV therapeutic, and 2.35 MV carbon and
aluminum beams at 0.01, 0.05, and 0.10 cGy for a thin (3 cm)

FIG. 5. Monte Carlo generated spectral distributions for carbon and aluminum targets over the diagnostic energy range (25–150 keV), for the maximum and
minimum energies used.
and a thick (15 cm) phantom. These images show that CNR
is improved with increased dose and a decreased phantom
thickness and also highlight the difference in image quality
between a 6 MV therapeutic beam and the low-Z target
beams. Figure 7 also shows nonuniformity in the central
region of the detector when imaged with the 6 MV thera-
peutic beam, as previously reported by Orton and Robar.9
Figure 8 shows the corresponding thin phantom CNR results
for cortical bone, inner bone, brain, and breast inserts as a
function of dose at the center of the phantom. At approxi-
mately 0.05 cGy, CNR values between a 6 MV therapeutic
beam and the low-Z target beams were increased by a factor
ranging from 6.2 to 7.4, from 8.5 to 9.7, from 6.7 to 7.5, and
from 2.2 to 2.7 for cortical bone, inner bone, brain, and breast,
respectively. It is notable that even for this thin phantom, the
CNR values for the 6 MV beam do not exceed 1.0, except
for cortical bone. Figure 9 shows the corresponding thick
phantom CNR results. At approximately 0.05 cGy, CNR
values between a 6 MV therapeutic beam and the low-Z
target beams were increased by a factor ranging from 3.7
to 4.3, from 5.0 to 6.0, and from 7.2 to 10.0 for cortical
bone, inner bone, and brain, respectively, with no significant
improvement for breast material for the thick phantom. No
measureable improvement in CNR was seen with reduction
of energy from 2.35 MV to 1.90 MV, or between target
materials. Ideally an increase in CNR should be observed
for the lowest energy and atomic number material.2,6,9,11
However, Fig. 5 illustrates that the relative increase in photon
fluence for carbon is primarily for the spectral component
below approximately 50 keV. In this very low energy range,
a significant proportion of photons will be absorbed within
the phantom or within the copper layer in detector, thereby
not contributing to imaging data.5
Previous measurements of
CNR (Refs. 2, 4–6, 8, and 11) with low-Z target beams have
involved a variety of materials and phantom geometries, and
therefore it is in general difficult to compare directly to CNR
values reported in the work of others. However, extrapolating
the reduction of measured CNR with the five-fold increase
in phantom thickness, we would expect that the gains in
planar CNR reported here are improved upon the 2.7–3.8
reported by Robar et al.11
for a 22.5 cm thick water phantom.
Although this direct comparison is difficult, it is intuitive
that increasing the proportion of photons in the diagnostic
energy range from 37% reported by Orton and Robar9
for
a 4.0 MV/aluminum to approximately 50% should increase
the CNR characteristics. It is clear from work reported here
and the reports by Ostapiak et al.,4
Roberts et al.,8
and
Robar et al.11
that image quality is degraded with increasing
phantom thickness and therefore is not ideally suited to large
thicknesses such as those seen in the pelvic region.
Figure 10 shows RMTF curves of the QC3 phantom, for
the carbon target at 1.90, 2.15, and 2.35 MV and for the 6 MV
therapeutic beam. Measured spatial resolution is consistent
with Connell and Robar,12
Roberts et al.,8
and Sawkey et al.10
with a spatial frequency of 0.40 lp/mm being distinguishable.
Slight degradation of the spatial resolution is apparent with

FIG. 6. Left ﬁgure shows Monte Carlo generated spectral distributions for the beryllium exit window and the carbon and aluminum targets, at 2.35 MV. Right
ﬁgure shows Monte Carlo generated spectral distributions at 2.35 MV for the carbon target with various thickness of copper (on the exit side) and the percentage
of photons with energy less than 25 keV.
the low energy carbon target beams, compared to 6 MV.
The frequency where RMTF equals 0.5, or f50, decreases by
approximately 0.1 lp/mm, from 0.42 to 0.32 lp/mm. With
beams of 4.5 MV/tungsten and 7.0 MV/beryllium, Connell
and Robar12
showed a relative reduction of f50 by 10.4%
–15.5% compared to a 6 MV therapeutic beam. As reported
by Roberts et al.8
and Connell and Robar,12
the loss of spatial
resolution results from larger angle electron scatter in the
FIG. 7. Planar images of the cortical bone insert at 2.35 MV with both aluminum and carbon targets and 6 MV therapeutic beam at 0.01, 0.05, and 0.10 cGy
for both thin (3 cm) and thick (15 cm) phantoms.

FIG. 8. Contrast-to-noise ratio results of the thin (3 cm) phantom for the four materials used at 1.90, 2.15, and 2.35 MV with both aluminum and carbon targets.
For comparison, a 6 MV therapy beam is also shown.
target when decreasing incident electron energy. In our data,
small decreases of RMTF (by 3% and 10% at 0.20 lp/mm
and 0.40 lp/mm, respectively) are observed with a reduction
of beam quality from 2.35 to 1.90 MV. These are minor
changes in RMTF compared to that observed by Connell
and Robar;12
however, the work by this group involved a
comparatively larger reduction of beam quality from 7 to
3.5 MV.
FIG. 9. Contrast-to-noise ratio results of the thick (15 cm) phantom for the four materials used at 1.90, 2.15, and 2.35 MV with both aluminum and carbon
targets. For comparison, a 6 MV therapy beam is also shown.

FIG. 10. Relative modulation transfer function for 1.90, 2.15, and 2.35 MV/carbon beams and a 6 MV therapeutic beam.
FIG. 11. Planar sagittal images of a sheep head at 2.35 MV with both carbon
and aluminum targets and a 6 MV therapeutic beam.
III.E. Qualitative image quality
Figure 11 shows planar sagittal images of a sheep head at
2.35 MV with both carbon and aluminium targets and a 6 MV
therapeutic beam. Relatively little variation in image quality
is observed between the aluminium and carbon targets, which
is consistent with the observations above. Images with the
2.35 MV beams show signiﬁcant enhancement in image
quality compared to a 6 MV therapeutic beam in both bone
and soft tissues. Roberts et al.8
showed similar gains in planar
image quality between their low-Z target beam line and a
6 MV therapeutic beam in planar images of a head and neck
phantom.
In summary, given (i) that the beam current is maximized
at 2.35 MV, thus shortening imaging times, (ii) that there is
no measurable improvement in CNR compared to the lowest
achievable energy, and (iii) that there is not signiﬁcant conse-
quence with regard to variation of spatial resolution, 2.35 MV
appears to be the optimal selection for low-Z target imaging
over the range explored here.
IV. CONCLUSIONS
In this work, we have shown that it is possible to decrease
the incident electron energy produced by a Varian 2100EX
linac to values ranging from 1.90 MeV to 2.35 MeV. In
electron mode, beam current can be increased from approx-
imately 600 MU/min to 2800 MU/min by altering set points
of the electron gun, grid, solenoid, and bunching magnets,
thereby deceasing image acquisition times. While it is
possible to lower the incident electron energy further, below

1.90 MeV, the beam current falls sharply to an impractical
value. Lowering the incident electron energy and increasing
the beam current has optimized the photon beams for imag-
ing, yielding approximately 50% of generated photons within
the diagnostic energy domain. This allowed an approximate
increase in CNR of cortical bone ranging from 6.2 to 7.4 and
from 3.7 to 4.3 for a thin and a thick phantom, respectively,
compared to 6 MV therapy beam for both aluminum and
carbon targets. Slight relative decreases of 3% and 10% at
0.20 lp/mm and 0.40 lp/mm, respectively, were observed
in spatial resolution when lowering from 2.35 MV to
1.90 MV. Planar images of a sheep head showed that
2.35 MV carbon and aluminum beams produced enhance-
ments to image quality compared to a 6 MV therapeutic
beam, in both bone and soft tissues. Based on the observations
and constraints encountered in this study, we suggest that
2.35 MV with a carbon target should be the normal operating
energy for further low-Z imaging with this linac. Removal
of the copper layer within the detector7,9
and detection
with a higher-efficiency receptor8
should allow for further
increases in image quality without further alteration of the
beam.
ACKNOWLEDGMENTS
The authors are grateful for support provided by Varian
Medical Systems for the funding of this project and the Nat-
ural Sciences and Engineering Research Council of Canada
for their additional financial support. The authors would like
to thank Robert Moran for his vital electronic and technical
support, Ian Porter for fabrication of apparatus, Dr. Robin
Kelly and Dr. Mammo Yewondwossen for their knowledge
and teaching of techniques utilized in this project.
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