the inﬂuence of bowtie ﬁltration on cone-beam ct image...

The influence of bowtie filtration on cone-beam CT image qualityN. MailRadiation Medicine Program, Princess Margaret Hospital, Toronto, Ontario M5G 2M9, Canadaand Ontario Cancer Institute, University Health Network, Toronto, Ontario M5G 2M9, Canada

D. J. MoseleyRadiation Medicine Program, Princess Margaret Hospital, Toronto, Ontario M5G 2M9, Canada;Ontario Cancer Institute, University Health Network, Toronto, Ontario M5G 2M9, Canada;and Department of Radiation Oncology, University of Toronto, Toronto, Ontario M5G 2M9, Canada

J. H. Siewerdsen and D. A. Jaffraya�

Radiation Medicine Program, Princess Margaret Hospital, Toronto, Ontario M5G 2M9, Canada;Ontario Cancer Institute, University Health Network, Toronto, Ontario M5G 2M9, Canada;and Department of Radiation Oncology and Department of Medical Biophysics, University of Toronto,Toronto, Ontario M5G 2M9, Canada

�Received 20 February 2008; revised 10 October 2008; accepted for publication 11 October 2008;published 4 December 2008�

The large variation of x-ray fluence at the detector in cone-beam CT �CBCT� poses a significantchallenge to detectors’ limited dynamic range, resulting in the loss of skinline as well as reductionof CT number accuracy, contrast-to-noise ratio, and image uniformity. The authors investigate theperformance of a bowtie filter implemented in a system for image-guided radiation therapy �Elektaoncology system, XVI� as a compensator for improved image quality through fluence modulation,reduction in x-ray scatter, and reduction in patient dose. Dose measurements with and without thebowtie filter were performed on a CTDI Dose phantom and an empirical fit was made to calculatedose for any radial distance from the central axis of the phantom. Regardless of patient size, shape,anatomical site, and field of view, the bowtie filter results in an overall improvement in CT numberaccuracy, image uniformity, low-contrast detectability, and imaging dose. The implemented bowtiefilter offers a significant improvement in imaging performance and is compatible with the currentclinical system for image-guided radiation therapy. © 2009 American Association of Physicists inMedicine. �DOI: 10.1118/1.3017470�

Key words: bowtie filter, Elekta synergy CBCT system, image quality, dose

I. INTRODUCTIONCone-beam computed tomography �CBCT� systems are em-ployed in radiation therapy to provide information regardingdaily target and normal tissue localization during the courseof fractionated radiation treatment. CBCT allows radiation tobe directed at tumors with greater accuracy and precisionthan was previously possible. However, improving CBCTimage quality further will allow more precise dose deliveryto the tumor. Precision in tumor localization and reduction insetup error is based on image quality of CBCT system. How-ever, flat-panel detectors used in CBCT have limited dy-namic range compared to CT. These limitations are exposedby the large variation of x-ray fluence at the detector acrossthe imaged field-of-view. It is often the case that the tech-niques used overwhelm the signal range of the detector at theperiphery of the patient, leading to a loss of information inprojections and artifacts in reconstruction due to the trunca-tion of anatomy. As a result, effects include the loss of skin-line and reduction in CT number accuracy, contrast, and im-age uniformity. The purpose of this study is to investigate therelative advantages and limitations of a bowtie filter for im-proving image quality of a commercially used CBCT ElektaSynergy System. The implementation of a bowtie filter inCBCT offers the method of addressing the issue of x-ray flux
variation across the detector.
22 Med. Phys. 36 „1…, January 2009 0094-2405/2009/36„

Previously, beam shaping x-ray filters or compensatorshave been used in computed tomography �CT� to modify thedistribution of x-ray flux across the field of view. Compen-sator filters have been used in CT for beam hardening1–5 andreduction6–8 in dose to the patient. These compensators havebeen used in CBCT for scatter9–11 and heel12 compensation.Initially, compensators have been used to improve the detect-ability in film radiographs with the delivery of more uniformfluence at the film.13

Bowtie filters potentially play a larger role in scatter,beam hardening, uniform fluence across the detector, anddose reduction. Quantitative investigations are performed onthe Elekta Synergy radiotherapy image-guidance system.These investigations provides quantitative evidence ofbowtie filter for improving skinline, image uniformity, CTnumber accuracy, contrast, and reduction in streak artifactsand patient dose.

II. MATERIALS AND METHODS

II.A. Cone-beam CT system

Investigations of bowtie filter use in CBCT were per-formed on Elekta oncology system shown in Fig. 1�a�. Thex-ray source uses a rotating anode x-ray tube �DunleeD604,
IL� with maximum potential of 150 kVp, 14 deg tungsten
221…/22/11/$25.00 © 2009 Am. Assoc. Phys. Med.

http://dx.doi.org/10.1118/1.3017470

http://dx.doi.org/10.1118/1.3017470

23 Mail et al.: The influence of bowtie filtration on cone-beam CT image quality 23

target, and 0.8 mm focal spot. The detector used was theRID1640-Al1 �Perkin Elmer, Wiesbaden, Germany� indirect-detection flat-panel imager with a 1024�1024 array of0.4 mm�0.4 mm pixels and a 0.55 mm thick CsI:Tl x-rayconverter. The geometry of the system was configured tohave a source-to-isocenter distance of 100 cm and source-to-detector distance of 153.6 cm.

The x-ray tube and flat-panel imager were placed orthogo-nally to the treatment head and its EPID imaging device. ThekV system shares a common axis of rotation with the MVtreatment source. The kV beam is 425�425 mm2 incidenton the flat-panel detector. Images can be acquired with threedifferent fields-of-view �FOV�; small, medium, and large.

(a)

85mm

117mm

-5

5

15

25

35

-10

Thickness(mm)

!"#$%

!&'%(#

)*!+(%,#%-

../ 1 23456 (7

3456(7

8&94:(7

;<:(!==$7>#!-

8>9;<:(!==$7>#!-

FIG. 1. �a� A photograph of the cone-beam CT Elekta Synergy CBCT system�c� Schematic of the CBCT offset geometry and medium FOV after 180 de

FIG. 2. �a� Photograph of the Elekta designed bowtie filter. �b� Bowti

Medical Physics, Vol. 36, No. 1, January 2009

The difference between the three is the edge to the kV centralbeam, which is 138 mm for the small FOV, 213 mm for themedium FOV, and 262 mm for the large field of view. Dur-ing 360 deg rotations the system acquires approximately 650projection images. The size of the flat-panel detector is 410�410 mm2 and is located at a fixed source to detector dis-tance, resulting in limited FOVs along the x and z direction.An offset scanning geometry was used, in which the imagerwas shifted laterally 10 cm and a corresponding asymmetricbeam, defined by an M20 collimator, was utilized to scan alarger FOV. This offset geometry �Fig. 1�b�� was used toprovide the optimum image sets, in terms of quality andreconstructed FOV. The Feldkamp algorithm for 3D filtered

15 40 65 90 115 140

Length (mm)

(b)

!"#$%

!&'%(#

)*!+(%,#%-

../ 1 23456 (7

3456(7

8(9

4:(7

;<:(!==$7>#!-

A schematic of the CBCT offset geometry at angle 0 deg and medium FOV.tion.

. �b�g rota

e thickness measured with needle gauge as a function of length.


back-projection algorithm was used for CBCT reconstruc-tion, with a Hann reconstruction filter. The total recon-structed volume used in this work was 410�410�264 mm3 with voxel size of 1�1�1 mm3.

II.B. Bowtie filter

An aluminum bowtie filter �Elekta, Filter Cassette Assem-bly, F1� was employed for all the tests described here. ThisF1 filter was inserted in the filter tray 30 cm from the x-raysource on an Elekta Synergy XVI unit. For the medium FOV,a cassette of M20 collimator was inserted between the x-raytube and F1 filter shown in Fig. 2�a�. There are various pos-sible designs for bowtie filters implemented on Elekta Syn-ergy XVI systems. Bowtie filters could potentially be opti-mized based on a variety of imaging tasks while accountingfor tradeoffs between image quality and patient dose. Thedesign of the bowtie filter used here is based on patient size,dose, nominal x-ray energy, and material.

The bowtie filter used in this study �Fig. 2�a�� has nomodulation in the z direction. The filter used for this workwas built based on the objective of achieving uniform flu-ence through a cylindrical water phantom of diameter300 mm measured at 120 kVp. The manufacturing materialwas aluminum of size 85�135 mm2. The bowtie filter thick-ness measured with a needle gauge as a function of length isshown in Fig. 2�b�. The filter was rigidly mounted in thecassette with a 1 mm thick acrylic covering protecting itfrom physical damage or scratching.

II.C. Dosimetry measurements

Dose measurements were made with and without a bowtie

!"

##$ &&

'$$ &&

("

!"#%#&'()*!+,*)-.!,)*#

1

1

1

1

1

1

FIG. 3. �a� A schematic diagram of the CTDI body phantom. �b� Percentage othe phantom center. Symbol represents measured data and solid line theoret

filter, such that the scanning geometry, phantom size, beam


quality, and mAs/projections were the same for both cases.The phantom �shown in Fig. 3�a�� was a PMMA CTDI do-simetry phantom with diameter of 320 mm and thickness of150 mm �RTI Electronics, Mölndal, Sweden�. Two low-density polyethylene �LDPE� blocks �32�32�15 cm3�were placed superior and inferior to the acrylic phantom tosimulate scattered dose. The center of the phantom wasplaced at the machine isocenter position. This specially de-signed phantom allows for accurate positioning of radiationdetectors. A 0.6 cc farmer-type ion chamber �NE 2571,Nuclear Associates� and Fluke 3504 electrometer at atmo-spheric pressure 756.2 �mmHg� and room temperature22 °C was used for dose measurements. The dose was mea-sured at five different radial positions of the phantom, start-ing from the phantom center to the skin. A total of fiveCBCT scans were obtained, with and without a bowtie filterat 120 kVp and 1.6 mAs per projection ��660 projections�.The ratio of doses with and without bowtie was modeled forall the radial positions of the phantom’s center to periphery.This model contained four fitting coefficients and is veryuseful to compute the dose reduction attributed to the bowtieat any radial position of the CTDI phantom.

II.D. Scatter-to-primary ratio measurements

Scatter-to-primary ratio was measured for imaging condi-tions with and without the bowtie filter using a beam-blockermethod.9 The SPR can be defined as

SPR =S

P, �1�

where S is the energy integrated signal of the scattered ra-2

0 1 0 2 3 4 51 50 52 5389&*9: &*")9+(# ;%!< (#+)#% =(<>

)"*$+$,''- .$,$$'/+01,$0.$,234+00,#5.$,306+#,-3.$,#1

'$ " 2'"

!"# $"$

+−+=

π

e reduction with bowtie filter is plotted as a function of radial distance fromt. �c� Circular �Catphan� phantom with an irregular annulus �human torso�.

/

602

604

670

673

621

622

f dosical fi

diation measured as the average of a 10�10 pixel area on


the FPI, and P is the signal from the primary radiation. Thesame CBCT system/geometry was used as discussed in Sec.II A. The SPR measurements were performed at the centerand edge of the projection image for the circular and irregu-lar �posterior-anterior �PA� view� phantoms. A 10�10�4 mm3 lead block was placed on the central axis before thephantom on the side facing the source. The scatter account-ing for shadowing a part of the phantom by the blocker wasnot included for this SPR data. Twenty �20� images of thecircular and irregular phantoms were acquired with and with-out the bowtie filter at the tube voltage of 120 kVp and ex-posures 1.0 and 1.6 mAs, respectively. The SPR measure-ments were performed on an irregular phantom at equivalentpatient dose techniques �0.1 mGy/exposure� for with andwithout the bowtie filter. This required an exposure of1.6 mAs/projection �bowtie filter� as compared to1.0 mAs/projection �without bowtie filter� at 120 kVp. Thesame procedure was used for the SPR measurements at thephantom’s edge except the block was moved from the centerto the edge of the phantom’s projection.

II.E. CBCT image quality: Phantom study

The Catphan 500 �The Phantom Laboratory, Salem, NY�was employed in these investigations. Experiments were per-formed with the Catphan 500 phantom alone �referred to as“circular phantom”� and with the Catphan 500 phantom in-serted in an irregular annulus �referred to as “irregular phan-tom;” Fig. 3�c�� shaped to reflect a humanoid torso. Twotypes of configurations were used to evaluate overall CBCTimage quality with and without a bowtie filter. The phantomwas positioned at the center of the imaging field of view withthe help of room lasers. A photograph of the cone-beam CTsystem is shown in Fig. 1�a�. Image acquisition proceededwith gantry rotation over 360 deg for all the cases duringwhich approximately 660 planar images were acquired at themedium FOV. Images of the circular phantom, with andwithout bowtie filter, were acquired under the same imagingconditions at 120 kVp and tube current 40 mA and exposuretime 20 ms per projection. Similarly, images of the irregularphantom with and without bowtie were acquired under thesame imaging conditions at 120 kVp and tube current 40 mAand exposure time 40 ms, respectively. The CBCT recon-struction algorithm includes a scatter correction technique asdescribed by Boellaard et al.14 The irregular phantom wasscanned using a conventional helical CT scanner �DiscoveryST 16 slice, GE Healthcare, Milwaukee, WI� at 120 kVp and300 mAs. The total number of CT reconstructed slices were264 with each slice thickness of 1 mm and voxel size of 1�1�1 mm3. The CT number for both CBCT and CT im-ages was defined as CT#= ��object−�water� /�water��1000 HU, where �object and �water are the linear attenua-tion coefficients for the object and water, respectively. Quan-titative comparison of image quality based on skinline recov-ery, CT number accuracy, CT number uniformity, spatialresolution, and contrast-to-noise ratio were evaluated using
the circular phantom tests performed in acquisitions with and

without the bowtie filter. Imaging metrics were also per-formed on images generated on a conventional CT scanner.

II.E.1. Uniformity

The uniformity measurements were performed on the cir-cular and irregular phantoms. The circular phantom containsa uniform, water-equivalent portion �CTP486� with CT num-ber �HU� within −10 and +10 HU at standard scanning pro-tocols. The CBCT images were analyzed in Matlab. FiveROIs of size 10�10 mm2 were selected in CBCT images ofthe phantom—one at the center and four at peripheral posi-tions symmetrically around the center, each within the cen-tral axial plane. The spatial nonuniformity �SNU� was de-fined as

SNU =VVmax − VVmin

1000� 100, �2�

where VVmax and VVmin are the maximum and minimummean voxel values, respectively, within each ROI.

II.E.2. Skinline artifacts/Reduction in CT number atskin zone

In CBCT images, the reduction in CT number �RCTN� atthe skin zone is attributed to the detector saturation near theobject periphery. Measurements were performed on circularand irregular phantoms. The reduction in CT number �de-noted �CT#� was measured using 3 ROIs at skin depths 5,10, and 20 mm. The size of each ROI was 4�4 mm2. Theaverage reduction in CT �RCTN� was calculated as

RCTN =HUCT − HUCBCT

HUCT

� 100, �3�

where HUCT and HUCBCT are the mean voxel values for agiven ROI measured in helical CT and CBCT images, re-spectively.

II.E.3. CT number accuracy and linearity

The CT number accuracy and linearity measurementswere performed on circular and irregular phantoms. Thephantom contains seven different cylindrical contrast inserts�12 mm diameter, CTP 404� with known electron densitiesand CT numbers. The mean HU value was measured for eachcontrast insert. The CBCT images were converted into polarcoordinates for convenient analysis. A 2D ROI size was ap-proximately 4�4 mm2 covering approximately 16 voxelsand was confined to within the contrast inserts. For conven-tional CT images, the 2D ROI size was 4�4 mm2. Thevoxel values in helical CT images of the phantom defined thetrue CT number.

II.E.4. Contrast-to-noise ratio

CNR measurements were performed using the circularand irregular phantoms �CTP 404 module�, including poly-styrene �PS, −100 HU� and low-density polyethylene�LDPE, −35 HU� inserts of 12 mm diameter. Contrast-to-
noise ratio �CNR� measurements were performed on circular


and irregular phantoms. For each individual phantom, theimaging conditions including kVp, exposure time, and tubecurrent were the same, with and without the bowtie filter.ROIs of size 4�4 mm2 ��4�4 voxels� within each insertwere used to measure the mean and standard deviation�noise� in HU value. The mean HU value was measured foreach contrast insert. The CNR �signal difference divided byaverage noise� values was calculated as

CNR = 2 ��HU�LDPE� − HU�PS��

Noise�LDPE� + Noise�PS�, �4�

where HU�LDPE� and HU�PS� are the mean voxel values inLDPE and PS, respectively, and Noise�LDPE� andNoise�PS� are the standard deviation in voxel values inLDPE and PS, respectively.

II.E.5. Spatial resolution and modulation transferfunction

The MTF measurements were performed on circular andirregular phantoms using the CTP528 module, which con-tains a radial high contrast spatial resolution bar pattern rang-ing from 1 through 21 line pair per cm. A bar pattern may beconsidered resolved if the bars can be perceived with somediscernible spacing or lowering density among them. Theradial design eliminates the possibility of streaking artifactsfrom other test objects. A circular pattern of pixels �passingthrough the bar patterns� was taken to obtain square-waveresponse function �SWRF�. The amplitude response at vari-ous spatial frequencies was analyzed between 1 and21 lp /cm. The MTF of the system was determined by decon-volving out the SWRF.15

TABLE I. The scatter, primary, and scatter-to-primarPhan500� phantom using lead strip of size 10�10�

Scanningtechniques

Scatter atphantom

center�ADU�

Primary atphantom

center�ADU�

Sppe

�

Bowtie120 kVp /0.8 mAs

302 801

Without bowtie120 kVp /1.6 mAs

630 1212

TABLE II. The scatter, primary, and scatter to prim�Cat-Irreg� phantom using lead strip of size 10�10�

Scanningtechniques

Scatter atphantom

center�ADU�

Primary atphantom

center�ADU�

Sppe

�

Bowtie120 kVp /1.6 mAs

1011 2280


2125 3221


1328 2059


II.F. Patient study

CBCT image quality measurements including uniformityand skinline reconstruction were performed on ten gyneco-logical patients. But only one patient is reported here be-cause there were no significant differences from patient topatient. Images were acquired with and without a bowtiefilter at 120 kVp and 1.6 mAs ��650 projections acquiredacross 360 deg. The same CBCT system geometry and pro-cedure was used as explained in the phantom study section.CBCT reconstructions with and without bowtie filter wereanalyzed quantitatively �difference images� and qualitatively�examination by a radiation oncologist�. Of specific interestin these studies were the gynecological patient contrast andthe reduction in CT number �cupping artifact� from the skin-line to the center of the image.

III. RESULTS AND DISCUSSIONS

III.A. Dose measurements for kV CBCT

The uncertainty of the ion chamber measurements wasdetermined �standard deviation/mean value� to be with in�0.55%. The dose reduction �define in methods� withbowtie filter as a function of radial distance from the phan-tom center is shown in Fig. 3�b�. It shows the achievabledecrease in dose at the center and periphery of the bodyphantom with the bowtie used in this study. Dose reductionsby 25% and 43% were found at phantom center and skindepth of 1 cm, respectively. It may be noted that the increasein dose reduction was consistent and followed a Lorentzianpeak function �LPF�. A theoretical fit represented by a solidline is shown in Fig. 3�b�. Dose reduction with a bowtie as a

os behind the center and edge of the circular �Cat-m3.

atmry�

Primary atphantomperiphery

�ADU�

Scatter-to-primary ratio

at center

Scatter-to-primary ratio

at edge

6594 0.377�0.02 0.092�0.02

12 730 0.519�0.03 0.102�0.02

ratios behind the center and edge of the irregularmm3.

atmry�

Primary atphantomperiphery

�ADU�

Scatter toprimaryratio atcenter

Scatter toprimary

ratio at edge

20 427 0.443�0.03 0.124�0.01

31 921 0.661�0.03 0.133�0.01

1937 0.645�0.025 0.132�0.01

y rati3.5 m

catterhantoripheADU

602

1301

ary3.5

catterhantoripheADU

2539

4240

2625


function of the radial distance, X, from the center of thedosimetry phantom was calculated using an LPF function:

D�X� = D0 +2AB

4��X − C� + B2 , �5�

where D0 is the intercept and A, B, and C are the fittingcoefficients. The values of the intercept and fitting coeffi-cients are listed in Fig. 3�b�. With the help of Eq. �5�, thedose reduction with the bowtie filter can be calculated forany radial distance within the body phantom. Also this em-pirical fit will assist in the design for an optimum bowtiefilter.

III.B. Scatter-to-primary ratio

The SPR at the center of each phantom is higher than theSPR measured at the edges. But the absolute scatter signal ishigher at the edges of both phantoms including circular andirregular. The SPR values for the circular and irregular phan-toms for with and without bowtie filter are listed in Tables Iand II, respectively. It shows that the scatter signal at thephantom’s edge is higher than the primary signal at the cen-ter for the case without the bowtie filter. The absolute scattersignal at the edges of both phantoms is almost two times lessthan without the bowtie filter. This high absolute scatter sig-nal strongly affects the SPR behind the center of the phantomand hence image quality. Also, the SPR values at the centerand edge are higher for the irregular phantom because of thesize and excessive fluence at the edge that creates high scat-ter and affects the SPR at the center �throughout the image�.

"

!" )"

!"""

#""

$""

%""

&""

"

&""

!" '" !'" &'" ('" %'")*+,- /012,3

45

678

9:

;*<=>0< 2>;<*,?>;<*,

("

FIG. 4. Trans-axial images of uniform water portion of a circular phantom�Catphan� acquired �a� without �window level −130 to 150 HU and �b� withbowtie filter �window level: −130 to 150 HU�. �c� Profiles through the uni-form water portion of the circular phantom with and without bowtie; theprofile position is indicated by a dotted line in image.

The bowtie filter shapes and modulates the beam to reduce


the unwanted signal at the object’s periphery and hence im-proves the SPR behind the phantom’s center and preventsdetector saturation at the skin zone.

III.C. CBCT image quality: Phantom study

III.C.1. Uniformity and skinline

Trans-axial CBCT images of a uniform water-equivalentportion of the circular phantom acquired without and withbowtie filter are shown in Figs. 4�a� and 4�b�, respectively.Profiles through the images are shown in Fig. 4�c�; the pro-file position is indicated by a dotted line in the image. Image�Fig. 4�a�� and its profile both show severe reduction in CTnumber near the skin zone, which reduces as a function ofskin depth. This is due to the signal range of the detector�saturation� at the periphery of the phantom, leading to a lossof information in projections that creates a skinline artifact inthe reconstruction. Approximately 15% of average reductionin CT number was measured at 1 cm skin depth. The CBCTimage with the bowtie filter �Fig. 4�b�� and its profile �Fig.4�c�� demonstrate significant improvement in the recovery ofCT number at the skin zone compared to the nominal case.Almost 100% of the CT number is restored at 1 cm skindepth. The bowtie filter improves image uniformity by �i�flattening of the fluence profile at the detector �thereby re-ducing the dynamic range requirement and preventing detec-tor saturation at the skinline� and �ii� attenuates more fluenceat the edge of the field which reduces the primary beam onthe periphery of the phantom, which, in turn reduces scatterthroughout the image. The reduction in scatter throughoutthe image has the greatest advantage in the center of theimage where the bowtie has minimally affected the primaryfluence. The mean spatial non-uniformity, as defined in Eq.�2�, was 2.1% and 9.8% for image acquired with and withouta bowtie filter, respectively. The higher non-uniformity num-ber as defined in Eq. �2� means poorer uniformity.

Trans-axial CBCT images of irregular phantom acquiredwithout and with bowtie filter are shown in Figs. 5�a� and5�b�, respectively, at the same window/level. The differenceimage �image without bowtie subtracted from the image ac-quired with bowtie� is shown in Fig. 5�c�. Profiles throughthe reconstructed slices of CBCT images are compared withhelical CT in Fig. 5�d�, with the bowtie filter case providingan improvement in edge definition. The images and profilesshow a severe reduction of CT number near the skin zone inthe nominal CBCT image. Since the irregular phantom islarger in size than the circular phantom, a higher exposure isrequired for sufficient signal behind the center of the object,resulting in increased pixel saturation near the phantom pe-riphery, which is the main source of CT number reduction atthe skinline. Even with the bowtie filter case, there is stillsome missing skin on the right shoulder of the image profilein Fig. 5�d�. This is due to the detector lag and ghostingeffects. As the object is irregularly shaped, the projected lo-cation of skinline relative to previous projections causes anaccumulated lag signal that degrades the reconstruction andresults in reduction in CT number in the skinline zone. The
gantry rotates clockwise around an irregular phantom, which


!"" !"" """""!""!""&"")(* )+*),*

)%*

(""

&""

!""

"

!""

&""

&&" &%" &$" &#" ("" (&")>+*<*>/ 7112:

45

678

9:

;*<=>0< 2>;<*,?>;<*,45

!&""

'""

$""

(""

"

(""

!" '" !'" &'" ('"

)>+*<*>/ 711:

456

789

:

;*<=>0< 2>;<*,?>;<*,45

)-*

).*

"

#

!$

&%

(&

%"

;*<=>0< 2>;<*, ?>;<*,5,@=/*A0,+ BCC-*,E

F45G

7H:

"2" @1 +I*/ E,C<=!@1 +I*/ E,C<=& @1 +I*/ E,C<=

FIG. 5. Trans-axial CBCT images of irregular �Cat-Irreg� phantom �a� without �window level: −350 to 500 HU� and �b� with bowtie filter �window level:−350 to 500 HU�. �c� The difference of image �b� and �a�. �d� Profiles through the uniform water portion of the irregular phantom. �e� The right shoulder of
�d� is magnified to see the missing skinline more clearly. �f� Reduction in CT number �RCTN� at several skin depths for with and without bowtie filter.
()* (+*

(,*

"""

"

"""

!"""

" "" !"" !"" &"" &"" ("" ("")*+,- /012,3

45

678

9:

;*<=>0< 2>;<*,2>;<*,45

(-*

!"""

#"

%"

"

%"

#"

!&"

>E,B- 45 /012,3

A33>

3B4

5 4

?45

789

: ?>;<*,(%* D*<=>0< 2>;<*,

!!"" $"" !"" %"" '""

FIG. 6. Trans-axial images of circular phantom with several contrast inserts acquired �a� without �window level −800 to 500 HU� and �b� with bowtie filter�window level −800 to 500 HU�. �c� The same image shown in �a�, but transformed into polar coordinate �window level −800 to 500 HU�. �d� Profilesthrough the image �c� with several inserts are compared with CT and without bowtie filter. �e� The difference between the measured CT and CBCT numberplotted as a function of Ideal CT number for several inserts. The empty circle and filled square symbols represent CT number error with and without bowtie.
A gray solid line represents a linear fit to bowtie data.


causes lag on the right side of the image. Mail et al.16 quan-titatively explained this lag effect in CBCT images of irregu-lar phantom acquired with and without bowtie filter. Evenwithout a bowtie filter, the reduction in CT number at theskinline is worse on the right shoulder than the left in Fig.5�d�, due to the lag and ghosting.16 This clearly explains thebowtie case where the reduction in CT number at the skinlineon the right shoulder in Fig. 5�d� is due to lag and ghosting16

�not due to the compatibility of the bowtie filter for an ir-regular shaped phantom�. The shoulder on the right hand sideof Fig. 5�d� is magnified in Fig. 5�e� to illustrate the reducedsignal at the periphery. The mean reduction in the CT num-ber at the skin region, as defined in Eq. �3�, is quantitativelyshown in Fig. 5�f� for three different skin depths for bothcases. The mean signal losses for skin depths of 5, 10, and20 mm were found 35.1%, 25.15%, and 20.12%, respec-tively, in cases without bowtie filter images. These numbersin bowtie cases were measured 13.8%, 12.5%, and 11%.

The spatial non-uniformities of the irregular phantomwithout and with bowtie filter �Figs. 5�a� and 5�b�� werecalculated to be 4.1% and 5.7%, respectively. Profiles in Fig.5�d� show that the CT and CBCT images with the bowtiefilter have similar uniformity within the circular phantomregion.

III.C.2. CT number accuracy and linearity

The central slice images of a circular phantom without

(" )"

*" !"

!"""

%""

&""

#""

!%""

&"""

!"""

4,B

+03,

E4

567

89

:2

FIG. 7. Trans-axial CBCT images of irregular �Cat�Irreg� with several inser�window level −350 to 500 HU�. �c� The difference of image �b� and �a�.−700 to 300 HU�. �e� The same image in �b� but converted into polar coordartifacts �window level −700 to 300 HU�. �f� The measured CT number plo

and with bowtie filter are shown in Figs. 6�a� and 6�b�, re-


spectively. For convenient analysis, the image in Fig. 6�b�was transformed into polar coordinate in Fig. 6�c�. A profilethrough several inserts, indicated by a dotted line in image, isshown in Fig. 6�d�. The dotted, solid black, and gray solidlines represent CT and CBCT with and without bowtie filterimage profiles, respectively. This signal profile shows thatCBCT image acquired with a bowtie filter is comparable tohelical CT. The comparison of CT number accuracy and lin-earity for circular phantom, without and with the bowtie fil-ter, is shown in Fig. 6�e�, where the difference of measuredCT number to CBCT is plotted versus the ideal CT number.The CBCT with bowtie filter demonstrated a clear improve-ment in CT number accuracy and an increase in image con-trast compared to the nominal cases. For the nominal case,inaccuracy in CT number is related to the assumption that thedetected signal is all primary fluence. Scattered fluenceclearly violates this assumption. Reducing scatter will lead tomore accurate CT numbers by allowing accurate inversion of

TABLE III. Comparison of the CT number accuracy and linearity using threedifferent techniques in terms of slope, intercept �at ideal CT# water�, and R2

values.

Techniques Slope R2 Intercept

CT 0.98 0.995 4Bowtie 0.85 0.99 43Without bowtie 0.71 0.98 −3

!"" !"" """""#$$#%$&$$&%$ '"

%"" &"" #"" !%"" &""">E,B- 45 6789:

5 +@B/*<=>0< 2>;<*,>;<*,*/,B3 745 +@B/:*/,B3 7;*<=>0< 2>;<*,:*/,B3 7?>;<*,:

+"

without bowtie �window level −350 to 500 HU� and �b� with bowtie filterhe same image in �a� but converted into polar coordinates �window levels to make it simple for analysis such as CT number accuracy and streakings a function of ideal CT number for several inserts.

4;?BBB

ts, �a��d� Tinatetted a


Beer’s law. It should be noted that other invalid assumptions�e.g., beam hardening� can lead to inaccuracies in CT num-ber as well. The improvement in CT number accuracy withthe bowtie filter is primarily attributed to the scatter reduc-tion �with improved beam-hardening and non-uniformityalso contributing to improved CT number accuracy� through-out the image and particularly near the phantom’s periphery.This reduction in scatter reduces the scatter to primary ratioand hence improves CT number accuracy. The CT numberaccuracies with and without a bowtie filter were found to be�21 and �68 HU, respectively.

Central slice images of the irregular phantom with andwithout a bowtie filter are shown in Figs. 7�a� and 7�b�,respectively. These images were then transformed into polarcoordinates to make the image analysis simple. The trans-formed polar coordinate images are shown in Figs. 7�d� and7�e�. In Fig. 7�f�, the measured CT number for Catphanphantom within irregular annulus is plotted versus the idealCT number. The measurements performed with a bowtie fil-ter showed greater linear CT number accuracy than nominal.The mean CT number obtained from a conventional CT im-age showed reasonably good agreement with the expectedCT numbers. Detailed comparisons of the CT number accu-racy and linearity in terms of slope, R2 values, and interceptsare shown in Table III. The slope of the fitting line and in-tercept values are reasonably improved. The CT number ac-curacy for the helical CT and CBCT acquired with a bowtieis higher than the nominal CBCT images. This is because ofthe influence of scatter radiation in nominal CBCT, which ismore severe as compared to that of a fan beam geometry. Ingeneral, x-ray scatter reduces image contrast, increases im-age noise, and may introduce reconstruction error intoCBCT.

III.C.3. Contrast-to-noise ratio and streak artifacts

CNR measurements were performed on Catphan and ir-regular phantoms �contrast inserts, LDPE and Polystyrene�.For Catphan phantom, the CNR, as defined in Eq. �4�, wasmeasured as 4.51�0.43 and 3.58�0.35, with and without abowtie filter, respectively �an improvement by a factor of 1.3at fixed mAs�. The improvement is likely due primarily toreduced scatter �with improved beam-hardening and non-uniformity also contributing to improved CNR�. The CNRvalues for the irregular phantom were found to be4.22�0.41 and 2.24�0.32, with and without bowtie cases,respectively �at equivalent mAs�. The improvement in CNRis found to be much greater for the irregular than circularphantom. A primarily possible approach is that for the with-out bowtie case, the amount of scatter in the irregular phan-tom is higher than in the circular phantom, particularly at thephantom’s peripheries. The CNR in the nominal case is sig-nificantly reduced for the irregular phantom than for the cir-cular phantom, while this reduction is very small in the caseof the bowtie filter. This is because of the scatter compensa-
tion. Comparing images acquired at equivalent patient dose

��0.1 mGy/exposure�, the comparison of CNR improveseven further, giving a CNR of 4.22�0.41 for the bowtie,compared to 1.61�0.3 without a bowtie.

The central slices of irregular phantom converted into po-lar coordinates with and without the bowtie filter are shownin Figs. 7�e� and 7�d�, respectively. The streaking artifactsare more visible around the high-density inserts in the polarcoordinate image acquired without the bowtie filter. Thesestreaking artifacts appear due to scatter and beamhardening.17 Beam hardening and scatter effects are worsebecause of the large size and irregular shape of the phantomand the presence of high contrast objects. The image in Fig.7�e� acquired with a bowtie filter is almost free of streakingartifacts. The scatter compensation reduces streaking arti-facts and improves CNR. The bowtie filter reduces beamhardening effects by minimizing the range of variation inx-ray energies presented to the detector. For example, con-sidering the PA projection of the irregular phantom, the meanenergy of the x-ray beam is 75 keV at the center of thedetector �path length 320 mm acrylic� and 64 keV near theperiphery �path length 100 mm acrylic�. With a bowtie filter,however, the range in energy is reduced to 76 keV at thecenter and 76 keV at the periphery. Beam-hardening artifactsare therefore expected to be mitigated significantly by thebowtie filter.

III.C.4. Spatial resolution and modulation transferfunction

The spatial resolutions �for 50% MTF� of the circularphantom’s images acquired with and without a bowtie filterwere measured as 2.3�0.11 and 2.2�0.1 lp /cm, respec-tively. The MTF is essentially equivalent with or without abowtie filter, with the curves of Fig. 8 showing consistentshape between the two cases. For the irregular phantom, thespatial resolutions �for 50% MTF� were measured as2.1�0.08 and 2.0�0.06 lp /cm with and without a bowtie,respectively. The measured MTFs for with and without the

0

0.2

0.4

0.6

0.8

1

0 1 2 3 4 5 6 7 8 9 10 11

Line pairs (cm-1)

MTF

Bowtiewithout bowtie

FIG. 8. Measured MTF performed on irregular phantoms using the CTP528module �Catphan500� plotted as a function of line pairs with and withoutbowtie filter.

bowtie filter have the same values within the error bar.


III.D. Patient study

CBCT images of a gynecological patient acquired withoutand with a bowtie filter are shown with the same windowlevel in Figs. 9�a� and 9�b�, respectively. Profiles through theimages are shown in Fig. 9�c�. Moderate improvements inuniformity and skinline reconstruction are seen with thebowtie filter. Improvement in skinline reconstruction isclearly shown in Fig. 9�d�, which is the magnified portionindicated on the right of Fig. 9�c�. The black spots in theimages are a brachytherapy applicator used as a fiducialmarker. The image in Fig. 9�a� demonstrates less clarity dueto high scatter-to-primary ratio while the image in Fig. 9�b�demonstrates clarity due to reduction in scatter-to-primarywith the bowtie filter.

IV. DISCUSSIONS

A fixed, shaped filter has been used to selectively attenu-ate x rays at the periphery of the exposed field-of-view tocompensate for the patient profile. The filter reduces the dy-namic range demands on the detector during a CBCT scan.The resulting flux on the detector is relatively uniform acrossthe detector and does not result in saturation in regions be-yond the skinline. Further, the reduction of the excessivefluence near the skin zone with a bowtie filter reduces theabsolute scatter signal at the object’s edge. This reduction inthe absolute scatter signal improves the SPR behind the cen-ter of the phantom and hence image quality. The influence ofthe bowtie filter on both dose and image quality was evalu-ated. The application of the bowtie filter for circular andirregular phantoms resulted in an improvement in imagequality and reduction/alteration in the dose distribution

!" !"

("

!"""

$""

&""

&""

$""

" %" !"" !%" &"" &%" ("" (%)>)*<*>/ 711:

45

678

9:

;*<=>0< 2>;<*,?>;<*,

FIG. 9. CBCT images of gynecological patient �window level −500 to 500Profiles through the region of interest, indicated by a solid line. �d� A portionclearly at skinline zone.

within the patient. Figure 3�b� demonstrates that the applica-


tion of the bowtie filter results in a reduction in dose by 43%at the skinline and 26% at the phantom center. Assessment ofcontrast-to-noise performance demonstrated an improvementwhen the bowtie filter is employed. This was specificallyobserved in the large sized phantom �irregular phantom�where scatter is increased. Overall, the bowtie improves CTnumber accuracy and image uniformity. This improvement isattributed to the scatter reduction, beam hardening reduction,and flatness of fluence at the detector.

The selection/design of a bowtie filter is challenging andwill always be less than ideal. It is not likely to be optimumgiven that a subject being imaged is significantly less thanuniform, not exactly cylindrical in shape, and not necessarilycentrally located in the x-ray beam. In such cases, it is pos-sible for one or more disjointed regions of saturation to occuror conversely to over-filter the x-ray flux and unnecessarilycreate regions of very low signal. Regardless of compro-mises associated with patient size, shape, anatomical site,and field of view, the use of a simple bowtie filter results inan overall improvement in the quality of CBCT images.

V. CONCLUSION

The implemented bowtie filter shows reduction in scatter-to-primary ratio for the circular and irregular phantoms. Thebowtie filter shows an improvement in image quality includ-ing uniformity, CT number accuracy, contrast-to-noise ratio,and skinline reconstruction of the circular and irregularphantoms. The implementation of a bowtie filer also reducedstreaking artifacts in CBCT images of an irregular phantomcompared to nominal. The implemented bowtie filter demon-strated an improvement in skinline reconstruction and unifor-

""

)""

"""

<%"

%""

&%"

"

&<$ &'$ (!$ (($

)>)*<*>/ 711:

;*<=>0< 2>;<*,?>;<*,

""

�a� without bowtie and �b� with bowtie �window level −500 to 500 HU�.c�, indicated by a rectangular box, is magnified to illustrate signal loss more

)

" !

HU�of �

mity in images of gynecological patients. This compensator


is static and makes many compromises for anatomical imag-ing site, patient size, and imaged field of view. The idealcompensator would optimize the fluence profile to accountfor numerous properties of the patient and imaging system.The investigated performance of the bowtie filter on imagequality of CBCT will create confidence in clinical implemen-tation of bowtie filters.

a�Author to whom correspondence should be addressed. Address for cor-respondence: Radiation Physics, Radiation Medicine Program, PrincessMargaret Hospital, 610 University Avenue, Toronto, Ontario M5G 2M9,Canada. Telephone: 416-946-4501 X5384; Fax: 416-946-6566; Elec-tronic mail: [email protected]

1R. Ning, X. Tang, D. Conover, and R. Yu, “Flat panel detector-based coneorbit: Preliminary phantom study,” Med. Phys. 30, 1694–1705 �2003�.

2E. Seeram, Computed Tomography: Physical Principles, Clinical Appli-cations & Quality Control �Saunders, Philadelphia, 1994�.

3R. Ning, University of Rochester, “Apparatus and method for x-ray scat-ter reduction and correction for fan beam CT and cone beam volume CT,”U.S. Patent No. 6,618,466 �2003�.

4H. H. Barrett and W. Swindell, Radiological Imaging: The Theory ofImage Formation, detection, and Processing �Academic, New York,1981�.

5G. X. Ding, D. M. Duggan, and C. W. Coffey, “Characteristics of kilo-voltage x-ray beams used for cone-beam computed tomography in radia-tion therapy,” Phys. Med. Biol. 52, 1595–1615 �2007�.

6R. G. Walters and R. W. Carlson, Technicare Corporation, “Computerizedtomographic scanner with shaped radiation filter,” 1981. U.S. Patent No.
4288695.

7J. E. Tkaczyk, Y. Du, D. Walter, X. Wu, J. Li, and T. Toth, “Simulation ofCT dose and contrast-to-noise as a function of bowtie shape,” Proc. SPIE5368, 403–410 �2004�.

8S. Itoh, S. Koyama, M. Ikeda, M. Ozaki, A. Sawaki, S. Iwano, and T.Ishigaki, “Further reduction of radiation dose in helical CT for lung can-cer screening using small tube current and a newly designed filter,” J.Thorac. Imaging 16, 81–88 �2001�.

9S. A. Graham, D. J. Moseley, J. H. Siewerdsen, and D. A. Jaffray, “Com-pensator for dose and scatter management in cone-beam computed to-mography,” Med. Phys. 34�7�, 2691–2703 �2007�.

10G. H. Glover, “Compton scatter effects in CT reconstructions,” Med.Phys. 9, 860–867 �1982�.

11J. Hsieh, GE Medical Systems Global Technology Co., LLC, “Methodand apparatus for optimizing dosage to scan subject,” 2003, U.S. PatentNo. 6,647,095.

12C. D. Smith, “X-ray filter,” 1940, U.S. Patent No. 2,216,326.13S. Mori, M. Endo, K. Nishizawa, M. Ohno, H. Miyazaki, K. Tsujita, and

Y. Saito, “Prototype heel effect compensation filter cone-beam CT,” Phys.Med. Biol. 50, N359–N370 �2005�.

14R. Boellaard, M. Van Herk, H. Uiterwaal, and B. Mijnheer, “Two-dimensional exit dosimetry using a liquid-filled electronic portal imagingdevice and a convolution model,” Radiother. Oncol. 44, 149–157 �1997�.

15H. E. Johns and J. R. Cunningham, The Physics of Radiology, 4th ed.�Thomas, Springfield, IL, 1983�.

16N. Mail, D. J. Moseley, J. H. Siewerdsen, and D. A. Jaffray, “An empiri-cal method for lag correction in cone-beam ct,” Med. Phys. 35�11�, 5187–5196 �2008�.

17J. H. Siewerdsen and D. A. Jaffray, “Cone-beam computed tomographywith a flat-panel imager: Magnitude and effects of x-ray scatter,” Med.
Phys. 28, �2�, 220–231 �2001�.
http://dx.doi.org/10.1118/1.1582470

http://dx.doi.org/10.1088/0031-9155/52/6/004

http://dx.doi.org/10.1117/12.535161

http://dx.doi.org/10.1118/1.2740466

http://dx.doi.org/10.1118/1.595197

http://dx.doi.org/10.1118/1.595197

http://dx.doi.org/10.1088/0031-9155/50/22/N02

http://dx.doi.org/10.1088/0031-9155/50/22/N02

http://dx.doi.org/10.1016/S0167-8140(97)00073-X

http://dx.doi.org/10.1118/1.2977759

http://dx.doi.org/10.1118/1.1339879

http://dx.doi.org/10.1118/1.1339879

the inﬂuence of bowtie ﬁltration on cone-beam ct image...

Documents