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Compensators for dose and scatter management in cone-beam computed tomography S. A. Graham Ontario Cancer Institute, Princess Margaret Hospital, Toronto, Ontario, M5G 2M9, Canada and Department of Medical Biophysics, University of Toronto, Ontario, M5G 2M9, Canada D. J. Moseley Radiation Medicine Program, Princess Margaret Hospital, Toronto, Ontario, M5G 2M9, Canada and Department of Radiation Oncology, University of Toronto, Toronto, Ontario, M5G 2M9, Canada J. H. Siewerdsen Ontario Cancer Institute, Princess Margaret Hospital, Toronto, Ontario, M5G 2M9, Canada and Department of Medical Biophysics, and Department of Radiation Oncology, University of Toronto, Toronto, Ontario, M5G 2M9, Canada D. A. Jaffray Radiation Medicine Program, Princess Margaret Hospital, Toronto, Ontario, M5G 2M9, Canada, Department of Radiation Oncology, and Department of Medical Biophysics, University of Toronto, Toronto, Ontario, M5G 2M9, Canada; and Ontario Cancer Institute, Princess Margaret Hospital, Toronto, Ontario, M5G 2M9, Canada 共Received 22 May 2006; revised 19 April 2007; accepted for publication 20 April 2007; published 11 June 2007兲 The ability of compensators 共e.g., bow-tie filters兲 designed for kV cone-beam computed tomography 共CT兲 to reduce both scatter reaching the detector and dose to the patient is investigated. Scattered x rays reaching the detector are widely recognized as one of the most significant challenges to cone-beam CT imaging performance. With cone-beam CT gaining popularity as a method of guiding treatments in radiation therapy, any methods that have the potential to reduce the dose to patients and/or improve image quality should be investigated. Simple compensators with a design that could realistically be implemented on a cone-beam CT imaging system have been constructed to determine the magnitude of reduction of scatter and/or dose for various cone-beam CT imaging conditions. Depending on the situation, the compensators were shown to reduce x-ray scatter at the detector and dose to the patient by more than a factor of 2. Further optimization of the compensators is a possibility to achieve greater reductions in both scatter and dose. © 2007 American Association of Physicists in Medicine. 关DOI: 10.1118/1.2740466兴 Key words: x-ray scatter, dose, compensator, bow-tie filter, cone-beam CT I. INTRODUCTION Standard practice in external beam radiation therapy relies on imaging techniques that lack the soft-tissue detectability necessary for precisely and accurately locating soft-tissue targets. Treatment setup relies on a combination of skin marks and/or megavoltage radiographs 共portal images兲 that can guide treatment based largely on the location of bony anatomy. Image guidance is an important development for more precise treatment delivery, with numerous techniques developed to accomplish visualization of soft-tissue structures of interest with the patient in treatment position.1 A recently developed technology for soft-tissue visualization and image-guided radiation therapy is kV cone-beam CT 共CBCT兲.2 CBCT systems employ conventional x-ray tubes and x-ray detectors such as image intensifier systems3,4 or flatpanel detectors.2,5 In this work, recently developed flat-panel detectors 共FPDs兲 are used to generate two-dimensional radiographic projection data that can be processed to form a volumetric reconstruction with sub-millimeter spatial resolution. FPDs provide digital images read out at frame rates up to 30 2691 Med. Phys. 34 „7…, July 2007 frames per second, providing sufficient numbers of projections for reconstruction with a single rotation of the sourcedetector pair. The use of CBCT for image guidance is gaining widespread popularity with a number of manufacturers offering CBCT imaging platforms integrated with linear accelerators. This allows radiography, fluoroscopy, and CBCT images to be acquired in the treatment position. Although the large area of FPDs is advantageous for CBCT acquisitions of clinically relevant fields of view, the FPDs currently employed in CBCT systems suffer from limited dynamic range. To achieve reasonable signal levels through the midline of many patients it is necessary to deliver a high exposure per projection. It is often the case that the techniques used overwhelm the signal range of the detector at the periphery of the patient, leading to a loss of information in projections and artifacts in reconstruction due to the truncation of anatomy. Another issue arising in CBCT imaging is the large quantity of scattered radiation generated in the patient and reaching the detector, to be hereafter referred to simply as scatter, due to the large projection field sizes 共⬃25⫻ 25 cm2兲 used in patient imaging.6 In some imaging geometries, the scatter 0094-2405/2007/34„7…/2691/13/$23.00 © 2007 Am. Assoc. Phys. Med. 2691 2692 Graham et al.: Compensators for cone-beam CT fluence exceeds the primary fluence at the detector. With such a large fraction of the detected fluence arising from scatter, there can be significant artifacts in the resulting reconstructions. The presence of scatter reduces contrast and also contributes additional x-ray quantum noise and induces localized artifacts such as streaking and reduced attenuation estimation at the center of the object 共known as a cupping artifact兲.6 Scattering within the patient also contributes additional dose to the patient, which does not necessarily contribute to corresponding improvements in image quality. The harmful effects of scatter in cone-beam CT imaging and the rapid deployment of these systems make this an important area of research. The development of robust correction schemes7–11 for removal of scatter in projections is an active area of research. In addition to postprocessing algorithms applied to data to reduce scatter effects it is also possible to use more mechanical methods to reduce scatter. This can be, for example, in the form of larger air gaps between the patient and detector,12–15 or by implementing scatter rejecting grids.16–18 Each of these approaches has limitations. The use of a grid is known to reduce the primary fluence at the detector, despite having delivered the dose to the patient, whereas the postprocessing methods are limited to artifact reduction and they cannot address the additional x-ray quantum noise induced by the scatter fluence. While the scatter correction and rejection methods mentioned above may be capable of reducing the effects of scatter in the reconstructed volumes, it would be beneficial to use scatter rejection techniques, which also reduce the dose to the patient. The use of a smaller longitudinal field-of-view 共FOVz兲 by reducing the collimator opening during imaging would reduce scattered radiation, but would also limit the total volume that could be reconstructed. The implementation of compensating filters19 共a subset of which is known as bow-tie filters兲 in CBCT offers an alternative method of addressing the issue of scatter. The scattered radiation in CBCT has been shown in a qualitative fashion to be reduced when using compensators.20,21 These compensating filters can also provide a method of both reducing scattered dose within the patient and relaxing the dynamic range requirements for the FPD. Compensators have existed in some form since it was recognized that the detectability in film radiographs would be improved if a filter were used to deliver a more uniform fluence at the film.22 Filters could be designed to account for more subtle changes in patient anatomy to further increase image quality.23 Furthermore, the use of appropriate filters in radiography has been shown to reduce patient dose.24 Compensation schemes were implemented early on in computed tomography scanning, with the EMI scanner using a water bag compensator.25,26 CT scanners later used metal “dodgers” made of low atomic number material to closely mimic water-based compensation.27 The purpose of compensation in CT is both to accommodate the dynamic range of the detectors and preferentially harden the x-ray beam.19–21,27 Patient dose is also reduced when introducing a compensating filter.21,28–30 Despite scatMedical Physics, Vol. 34, No. 7, July 2007 2692 FIG. 1. Schematic of the bench-top cone-beam CT system used for investigations into the dose and scatter reduction capabilities of compensating filters. Parameters that were varied in order to alter the scatter and dose included the compensator, z collimation, phantom size, and the air gap between the phantom and detector. The geometry of the system is also summarized in Table I. ter not playing as large a role in conventional CT imaging as in CBCT imaging, scatter induced artifacts have been reported as being reduced by use of a bow-tie filter.31 For compensator use in multi-row CT or CBCT it is necessary to manufacture 2D shaped filters32 to modulate the x-ray fluence across the imager. Filters of this type can be expected to provide similar benefits as bow-tie filters provide in conventional CT, though they may potentially play a larger role in scatter and dose reduction when applied to CBCT imaging. Although qualitatively it is recognized that compensators offer benefits to CBCT, it is necessary to quantify the magnitude of effect obtained under realistic conditions. The investigations are performed assuming a cone-beam CT imaging geometry consistent with current radiotherapy imageguidance systems. These investigations will provide quantitative evidence of the value of using compensators for the purpose of improving cone-beam CT image quality, reducing the magnitude of x-ray scatter at the detector, and reducing patient dose. II. MATERIALS AND METHODS A. Experimental cone-beam CT bench Investigations of compensating filters for CBCT were performed on a bench-top CBCT system 共Fig. 1兲. The x-ray source was a Rad-94 rotating anode x-ray tube 共Varian Medical Systems兲 in a Varian Sapphire housing with a maximum potential of 150 kVp, 14° tungsten rhenium molybdenum graphite target, and 0.4 and 0.8 mm focal spot sizes. The geometry of the system was configured to have a source-toisocenter distance 共dSA兲 of approximately 100 cm, and a source-to-detector distance 共dSD兲 of approximately 155 cm, which is consistent with the geometry of clinical CBCT sys- 2693 Graham et al.: Compensators for cone-beam CT 2693 TABLE I. Nominal imaging geometry and acquisition/reconstruction parameters. Imaging system geometry Focal spot to isocenter distance 共dSA兲 Nominal focal spot to detector distance 共dSD兲 Nominal isocenter to detector distance 共dAD兲 Focal spot to compensator distance 共dcomp兲 Axial field of view 共FOVx兲 Maximum longitudinal field of view 共FOVz兲 Minimum longitudinal field of view 共FOVz兲 100 155 55 9.7 25.6 19.2 2 cm cm cm cm cm cm cm Image acquisition Tube potential Tube current Exposure time Head phantom Body phantom 20.5 cm diameter watercylinder Frame rate Number of exposures Nominal beam filtration 10 ms 20 ms 4 ms 1 s/projection 320 4 mm Al, 0.13 mm Cu Image processing and reconstruction Defect pixel removal Reconstruction filter Maximum reconstruction volume Voxel dimension 3 ⫻ 3 median filter Hamming 25.6⫻ 25.6⫻ 19.2 cm3 0.05⫻ 0.05⫻ 0.05 cm3 120 kVp 100 mA tems used for image guidance during radiation therapy. Object rotation during CBCT acquisition was performed with a direct-drive rotation stage 共Dynaserv DM1060B兲. The detector used was a 14 bit indirect-detection flat-panel detector 共PaxScan 4030A, Varian Medical Systems兲 with a 600 m thick CsI:Tl scintillating layer. The pixel array was a 2048 ⫻ 1536 共40⫻ 30 cm2兲 amorphous silicon flat panel matrix with a pixel pitch of 194 m. This detector provides a 25.6⫻ 19.2 cm2 maximum FOV at isocenter when using a centered detector. Table I summarizes the relevant parameters of the bench-top CBCT system. B. Compensating filters There are numerous possible designs for compensators implemented on CBCT imaging systems. Compensators could potentially be optimized based on a variety of imaging tasks while taking account of the tradeoffs between image quality and dose to the patient. Patient size, scatter, dose, and exposure dependent detective quantum efficiency 共DQE兲 of the detector are examples of quantities that could be considered in the design of modulating filters for cone-beam CT. Simple representative compensators were used in this work to examine the dose and scatter effects of modulating the input fluence. Although compensators could be manufactured to have variable modulation across the two-dimensional cone-beam projection images, the compensators used in this study have constant modulation in the z direction. The two bow-tie filters constructed for this work were built based on the objective of achieving uniform fluence through a cylindrical head phantom with a diameter of 16 cm Medical Physics, Vol. 34, No. 7, July 2007 FIG. 2. Thickness profile designed for the 8:1 and 4:1 modulation factor compensators, shown with machining tolerances for the 8:1 filter. Measurements of the compensator thickness performed after the completion of machining, shown here only for the 8:1 compensator, indicate that the 4:1 compensator corresponded well with the designed thickness of 0.1 mm, while the 8:1 compensator was measured to be 0.14± 0.03 mm at the center. This discrepancy in thickness at the center was corrected for in the measurements shown in this work. The width of the compensators continues as a flat function out to ±62.5 mm, but the axes of this figure have been adjusted to emphasize the shape of the machined section of the copper plates. measured at 120 kVp. The shape of the filters 共Fig. 2兲 was then smoothed using an unweighted sliding average to remove discontinuous first derivatives that are difficult to machine and that could induce artifacts under conditions of imperfect geometric calibration. The compensators were manufactured from 99.9% pure copper plates, as this material made it possible to machine to the desired shapes and would allow thin compensators that could be placed between the x-ray tube and collimator assembly at approximately 10 cm from the focal spot. At this position it is expected that very few x rays scattered from the compensators will reach the detector because of the large 共145 cm兲 air gap. The bowtie filters were manufactured from 125⫻ 110 mm2 sheets of Cu approximately 2.4 and 1.6 mm in thickness. The plates were machined to the desired shape down to a copper thickness of approximately 0.1 mm at the center 共central axis of the beam兲. The compensators were mounted on 2.5 mm thick acrylic sheets for mechanical support. The chosen thicknesses and profiles provide compensators with modulation factors of approximately 8:1 and 4:1, where the modulation factor was defined as the ratio between the measured detector signal at the center of a projection, through the thinnest area of the compensators, to that at the periphery of a projection 共in plateau of compensator profile兲 where the beam is attenuated by 0.1, 1.6, or 2.4 mm of copper for the 1:1, 4:1, and 8:1 compensators, respectively. The modulation factor was defined only in terms of the measured detector signal in flood field images where the compensators are in place but no object is present in the beam. The added filtration used on the CBCT system is commonly 4 mm of aluminum and 0.13 mm of copper 共i.e., no compensator, referred to here as the 1:1 compensator兲. When placing the 4:1 or 8:1 compen- 2694 Graham et al.: Compensators for cone-beam CT 2694 FIG. 3. 共a兲 SPR and 共b兲 the corresponding SRF value at the center of the projection in a 16 cm diameter acrylic phantom 共head兲 and 共c兲 SPR and 共d兲 SRF for a 32 cm diameter acrylic phantom 共body兲 for varying longitudinal FOV and the three compensators employed in this study. The value of dAD was held at 55 cm for all scans 共giving a dAG of 39 cm for the body phantom and 47 cm for the head phantom兲. Measurements are also shown for the beam-blocker method for the 1:1 共open circle兲 and 8:1 共open square兲 compensators. sator in the beam, the 0.13 mm copper filter was removed to obtain a primary fluence at the center of the beam that is nearly equivalent in all cases. C. Scatter-to-primary ratio measurements The SPR at the center of the detector was measured for the 8:1, 4:1, and 1:1 compensators. The SPR is defined as SPR = S , P 共1兲 where S is the energy integrated signal of the scattered radiation measured as the average of a 100⫻ 20 pixel2 area on the FPD, and P is the signal due to the primary radiation. The SPR was measured at the location of the central axis in the projection images for both an acrylic 16 cm cylindrical head phantom and the 16 cm head phantom placed inside of an acrylic annulus with an outside diameter of 32 cm. Acrylic has a Compton scattering contribution to the mass attenuation coefficient of approximately 2% to 3% higher than that of tissue 共water兲 in the relevant energy range 共for example, acrylic has values of 0.173 cm2 g−1 at 60 keV and 0.158 cm2 g−1 at 100 keV, while water has values of 0.177 and 0.163 cm2 g−1 at the same energies兲,33 and we believe this is sufficiently close to demonstrate the effect the compensators would have on scatter in patient imaging. The projection images used in the measurement of the SPR were an averaged set of 30 in-air exposures 共flood field images兲 with dark fields subtracted. Projections were acquired at 120 kVp Medical Physics, Vol. 34, No. 7, July 2007 with 1 mAs/projection for the head phantom and 2 mAs/projection for the body phantom. The determination of the value of SPR in the images was accomplished by varying the longitudinal field of view 共FOVz兲 from approximately 18.5 cm at isocenter down to the minimum field of view achievable with the collimator assembly, which is approximately 2 cm 共±0.05 cm兲. The axial field of view 共FOVx兲 was fixed at the maximum detectable 共25.6 cm兲. The total signal in the projections was assumed to be the sum of the primary and scattered radiation signals. By plotting the total signal measured in images against FOVz, a curve could be found which was employed to extrapolate the total signal to a FOVz of zero. For the 16 cm phantom the data were fit to a square root function, while the 32 cm data were found to be more linear with field size. At the point of zero FOVz the signal was attributed completely to the primary fluence and was assumed to be constant in all of the projections. Subtracting the primary signal estimate in all projections permitted the determination of the scatter from the phantom at the center of the projection images. Additional measurements of the SPR were performed with a variable air gap 共dAG兲 between the phantom and detector, while maintaining the FOVz at 18.5 cm to compare the effects of the modulating filters with alternative imaging geometries. A secondary check of some of the SPR measurements was also performed using a beam-blocker method.34 This was done in order to compare the SPR values obtained by varying the longitudinal FOV with a second method of determin- 2695 Graham et al.: Compensators for cone-beam CT 2695 FIG. 4. Variation in SPR for different spacing between the phantoms and flat-panel detector. The 共a兲 SPR and 共b兲 SRF results are shown for the head phantom along with the SPR and SRF shown respectively in 共c兲 and 共d兲 for the body phantom. ing SPR. The lead blocker had approximately an 8 mm width and 1 cm thickness. Measurements were made at a single FOVz of 18 cm using the 16 and 32 cm phantoms and the 1:1 and 8:1 compensators. A correction accounting for shadowing of part of the phantom by the blocker was accomplished by performing measurements with blockers of varying widths 共5 mm up to 14 mm兲 and extrapolating to a blocker size of zero. In order to quantify the benefits offered by the compensating filters, the scatter reduction factor 共SRF兲 was defined as the ratio 共for a given phantom, air gap, and FOVz兲 of the SPR measured with a compensating filter divided by the SPR corresponding to a flat filter 共i.e., 1:1 case兲. The SRF shows the magnitude of scatter reduction offered by the compensators as well as the trends in scatter reduction when altering FOVz and air gap. An SRF of one would indicate no change in scatter magnitude, less than one would indicate scatter removal, and greater than one would signify elevated scatter. D. Dose measurements Dose was measured in the acrylic 16 cm head and 32 cm body phantoms exposed at 120 kVp with 1 mAs/projection for the head phantom and 2 mAs/projection for the body phantom. The total dose was found based on an acquisition of 320 projections across 360°. Measurements were perMedical Physics, Vol. 34, No. 7, July 2007 formed with a 0.6 cc Farmer ionization chamber 共NE 2571, S/N 1700兲 and a Keithley 35040 electrometer 共S/N 62968兲 with a 300 V bias. The charge integrated by the electrometer was converted to dose using an air kerma calibration factor 共Nk兲 for the specific ionization chamber/electrometer pair determined from a standard ionization chamber/electrometer pair calibrated by the National Research Council 共Ottawa, Ontario, Canada兲. The Nk corresponding to 120 kVp was 420 mGy/ 10−8 C. The dose was measured for variations in compensator modulation, longitudinal field of view 共FOVz兲, and position of the dosimeter in the phantom. The dose reduction factor 共DRF兲 was defined similarly to the SRF and was used to quantify the dose reductions achieved with the compensators in place. Slight differences in filter thickness along the central axis for the three choices of compensator were expected to play a small role in creating the differences in measured dose. Measurements of the copper thickness after machining was complete indicated that the 4:1 compensator was 0.10± 0.03 mm at the center, while the 8:1 compensator was 0.14± 0.03 mm at the center. To provide the most accurate comparison of the 1:1, 4:1, and 8:1 compensators the differences caused by the differing primary fluence along the central axis were corrected out of the measured doses. This was performed by measuring the detector signal under narrow beam geometry 2696 Graham et al.: Compensators for cone-beam CT 2696 that of the 1:1 filter. Appropriate adjustments were made to the doses measured with the 8:1 filter to account for this systematic discrepancy. E. Reconstructed images FIG. 5. SPR measured as a function of modulation factor for the 16 cm 共head兲 and 32 cm 共body兲 phantoms. 共2 ⫻ 2 cm2 field at isocenter兲 with each of the compensators in place and no object in the beam. Ideally, the measured signal through each of the compensators would be identical. The measured signal through the 1:1 filter was compared to the signal through the center of the 4:1 and 8:1 compensators and deviations from the 1:1 signal were used to correct the measured doses. These measurements showed that the transmission through the center of the 4:1 compensator matched 共within 2%兲 that of the 1:1 filter. The transmission through the 8:1 filter was found to be approximately 7% lower than The influence of the compensators on CBCT image quality was assessed using a simple cylindrical water phantom for examination of cupping artifacts and noise. The acrylic wall of the phantom had an outside diameter of 20.5 cm and a thickness of 0.5 cm. The height of the phantom was approximately 28 cm. Projection images were acquired for FOVz values of 2 and 18.5 cm at 120 kVp and 0.4 mAs/projection with 320 projections being acquired through 360°. The raw projection data acquired at 2048 pixels by 1536 pixels were down-sampled to 512⫻ 384, gain and offset corrected using floods taken with the proper compensator in place, and median filtered to remove defective pixels from the images. Volumes of 25.6⫻ 25.6⫻ 1.3 cm3 共0.05⫻ 0.05⫻ 0.05 cm3 voxels兲 were reconstructed using the Feldkamp algorithm35 for each case 共FOVz = 2 cm, 18.5 cm; compensation: 1:1, 4:1, 8:1兲 so that comparisons could be made in terms of apparent cupping, noise, and any compensator-induced artifacts. A summary of the reconstruction parameters is given in Table I. Images of the phantom were also performed with two liver-equivalent tissue inserts 共GAMMEX rmi, w e = 1.07兲 suspended in the phantom for the purpose of contrast-tonoise ratio 共CNR兲 measurements. The CNR was defined as FIG. 6. Dose measured in the 16 cm acrylic head phantom at 共a兲 the center and 共c兲 the periphery of the phantom. The compensators reduce the dose in both locations, as demonstrated by the DRF for the 共b兲 center and 共d兲 peripheral doses. Medical Physics, Vol. 34, No. 7, July 2007 2697 Graham et al.: Compensators for cone-beam CT 2697 FIG. 7. Dose measured in the 32 cm acrylic body phantom at 共a兲 the center of the phantom and 共c兲 the periphery of the phantom, shown with the DRF for 共b兲 the center of the phantom and 共d兲 the periphery of the phantom. the difference between the mean attenuation value in the insert and in the surrounding water 共both measured in a 100 voxel ROI兲, divided by the standard deviation of the attenuation value in the water. The images acquired for the CNR measurements were measured at 4 mAs/projection, and an additional 3 mm of copper filtration added to the beam. This was necessary to reduce beam-hardening artifacts present in images. All other parameters for the CNR measurements were the same as those given above. III. RESULTS A. Scatter-to-primary ratio measurements The SPR at the center of each phantom was found to decrease as the longitudinal field of view was reduced, and the modulation factor was increased 共Fig. 3兲. The SPR in the head phantom reached a value greater than 100% for the highest FOVz for the 1:1 modulation factor filter. With the 4:1 and 8:1 compensators the magnitude of the scatter signal does not reach the same level as the primary signal for even the highest field of view. The SRF showed a constant 20% reduction in scatter for all FOVs when applying the 4:1 compensator, and a greater than 40% reduction for the 8:1 compensator. Beam-blocker measurements agreed well with the extrapolation measurements giving SPR values of 1.02± 0.03 for the 1:1 compensator and 0.54± 0.03 for the 8:1 compensator at a FOVz of 18 cm, compared with 1.01± 0.03 and 0.60± 0.02 for the 1:1 and 8:1 compensators, respectively, at 18.5 cm using the extrapolation method. Medical Physics, Vol. 34, No. 7, July 2007 For the body phantom similar effects were noticed, though the levels of scatter were much higher. The scatter signal equalled that of the primary signal when applying flat filtration even at relatively small FOVz with an SPR of 100% at a longitudinal field of view of approximately 3.7 cm. The same SPR when using the 4:1 and 8:1 compensators could be achieved with a FOVz of 6.5 and 9.8 cm, respectively, demonstrating the large gains that can be made through the implementation of compensation schemes in cone-beam CT. The reduction of scatter was more pronounced for the body phantom than the head phantom, with SRF values of approximately 0.6 and 0.4 for the 4:1 and 8:1 compensators, respectively. Beam-blocker measurements gave values of 4.4± 0.4 and 2.1± 0.2 for the 1:1 and 8:1 compensators. With FOVz fixed, compensating filters reduced the SPR regardless of the air gap 共Fig. 4兲 for the range of geometries examined. Although the primary signal in each pixel decreased as the detector moved away from the phantom and x-ray tube, the scatter fell off more quickly, resulting in a decreasing SPR. This is consistent with the well-documented role of air gaps in reducing scatter at the detector.13,15 While the smallest air gap caused a large value of SPR when using a flat filter, the application of the 8:1 compensator reduced the SPR such that the SPR at the smallest air gap for the 8:1 compensator was less than that at the highest air gap for the 1:1 case for both the head and body phantoms. SRF values remained close to the values found at the nominal distance between the phantom and the detector 共dAD = 55 cm兲. There 2698 Graham et al.: Compensators for cone-beam CT 2698 FIG. 8. Dose as a function of the distance from the center of the phantom in both the 共a兲 head and 共c兲 body phantom. The DRF is shown for the 共b兲 head and 共d兲 body phantoms. was a slight decrease in SRF as the air gap was increased, though the slope was not significantly different from zero. Figure 5 summarizes the SPR as a function of the modulation factor. The SPR results for an 18.5 cm longitudinal field of view at isocenter and a dSD of 155 cm at the center of the two phantoms clearly show a substantial decrease in SPR as the modulation factor of the compensator increases. The effect is larger in the body phantom where the SPR is much greater. B. Dose measurements The dose at the center and periphery of the 16 cm acrylic head phantom was shown to vary with FOVz and compensation 共Fig. 6兲. At the center of the phantom the primary intensity of the beam was approximately constant, so the reduction in dose when increasing the compensator modulation was due to the reduction in scatter. The dose at the center of the phantom for the 8:1 compensator was approximately 75% of the dose measured at the center of the phantom when using a flat filter. The compensators further reduced the dose at the periphery of the phantom, since both the primary and scatter were reduced. When using the 8:1 compensator instead of the flat filter the dose was reduced to approximately 55% of the original value at the periphery. The reduction in dose measured at the periphery of the phantom when employing compensating filters came largely from the reduction in primary x rays. The DRF appears constant across all fields-of-view for both positions in the phantom, except for a Medical Physics, Vol. 34, No. 7, July 2007 suspicious drop in the DRF when the dose was measured with the lowest FOVz at the periphery of the phantom. Similar trends in dose reduction were also seen when measuring dose in the body phantom while varying the compensation scheme and FOVz 共Fig. 7兲. The dose decreased to about 60% and 40% of the value measured with the 1:1 modulation filter at the center and periphery, respectively, when applying the 8:1 filter. There are large reductions in dose at the periphery of the phantom because the compensators used in this study were designed according to the fluence through the 16 cm acrylic head phantom and could not be placed closer to the x-ray tube to adjust the magnification of the compensator shape. Thus the fluence pattern applied to the body phantom reduced the primary fluence to a larger degree than a filter that was designed to accommodate the exact shape of this larger phantom. Despite this large reduction in dose to the periphery of the phantom, it was seen that, unlike the case of the head phantom, the dose at the periphery of the phantom for the 8:1 compensator was larger than the dose at the center. The dose as a function of distance from the center of the phantom for the different compensators was measured in the head and body phantoms 共Fig. 8兲. In the head phantom the dose increased towards the outside of the phantom in the 1:1 case, was nearly constant for the 4:1 case, and was reduced for the 8:1 case. In the body phantom, the dose increased when moving away from the center of the phantom up to a point where the dose began to drop. This drop is attributed to the 2699 Graham et al.: Compensators for cone-beam CT 2699 C. Image reconstructions FIG. 9. For both the head and body phantoms the dose is shown to vary with varying modulation factors. The dose shown in 共a兲 is measured at the center of the phantoms where the primary intensity of the beam is approximately the same for all modulations. In 共b兲 the dose is shown at a depth of 1 cm in the two phantoms where both the scatter and primary fluences are altered by the compensators during a CBCT scan. edge of the phantom moving outside the field-of-view at approximately 12.8 cm from the center of the phantom. The peak dose was found at different distances from the center of the phantom because of the changes in dose due to scatter. The DRF shows, as expected, that the dose reducing capabilities of the compensators increased when moving away from the center of the phantom—this is a combination of the reduced primary and the reduction in primary-induced scatter fluence within the phantom. Figure 9 shows the achievable decrease in dose at the center and periphery of the head and body phantoms with the compensators used in this study. Similar trends were seen in both phantoms when varying the modulation of the compensator. The decrease in dose measured at the center of the phantom was due only to the decreased scatter dose arising from the modified primary fluence pattern offered by the compensators. The decrease in dose shown at the periphery of the phantoms was due to a combination of the decrease in scatter, along with the decrease in primary fluence caused by the shape of the filter. Medical Physics, Vol. 34, No. 7, July 2007 Reconstructions were performed on projection images of a cylindrical water phantom for a FOVz of 2 cm 共low scatter, Fig. 10兲 and 18.5 cm 共high scatter, Fig. 11兲. In the low scatter case the 1:1 compensator gives an image with a very uniform reconstructed attenuation value. The 4:1 image is similar in signal uniformity to the 1:1 case, though some reduction in the attenuation value is seen at the edges of the phantom. Large discrepancies are seen with the 8:1 compensator image where a considerable “cap” is seen in the image with a peak in reconstructed attenuation values at the center of the image. The cause of the nonuniformity in the reconstruction is hypothesized to be the change in the energy spectrum of the x-ray beam when passing through the compensators. The shape of the compensators results in a spectrum with a higher mean energy as the beam moves away from the central axis. The harder x-ray spectrum directed at the outside of the cylinder, caused by the thicker copper on the compensator, results in a larger percentage of the x rays passing through the cylinder resulting in a perceived reduction in x-ray attenuation at the periphery of the cylinder. A signalto-noise 共SNR兲 analysis of the images was performed in 10 voxel by 10 voxel regions near the center and edges of reconstructed images. This analysis demonstrates that the compensators provide a more uniform pattern of noise across the images, with SNR values at the edges of the phantom of 72± 4 for the 1:1 case, 54± 5 for the 4:1 compensator, and 41± 4 for the 8:1, and SNR values near the center of the phantom of 33± 3, 36± 5, and 37± 5 for the 1:1, 4:1, and 8:1 compensators, respectively. Increasing the FOVz up to 18.5 cm significantly increases the percentage of scatter present in the system. In the image acquired with the 1:1 compensator a severe cupping artifact caused by the increased scatter is now seen. The reduction in scatter afforded by the 4:1 compensator provides a substantial reduction in the cupping due to scatter. An additional benefit is that the 4:1 image is acquired using an estimated 15% to 20% decrease in dose compared to the 1:1 case, though the SNRs for these two images are not significantly different 共72± 4 at the edge and 42± 3 at the center of the 1:1 image, compared to 71± 5 at the edge and 48± 4 at the center of the 4:1 reconstruction兲. In the 8:1 case the change in the spectrum across the beam continues to cause an artifact in reconstruction. In fact, the capping artifact in the 8:1 image has increased in the presence of higher amounts of scattered x rays. This effect is hypothesized to be due to the change in the distribution of scattered x rays relative to primary x rays across the phantom in the presence of the 8:1 compensator. The SNR at the center of the 8:1 image is comparable to the 1:1 and 4:1 images with a value of 42± 6. At the edge of the 8:1 reconstruction the SNR dropped to 42± 4. With the large errors caused by beam hardening in the 8:1 images it is difficult to assess any improvements in image quality afforded by the compensating filters. To reduce these errors images were acquired with an additional 3 mm of copper filtration to considerably harden the x-ray beam. Although this may not be a realistic approach for everyday 2700 Graham et al.: Compensators for cone-beam CT 2700 FIG. 10. Reconstructions of the cylindrical water phantom with a FOVz of approximately 2 cm for the 1:1, 4:1, and 8:1 compensating filters. imaging because of the heating of the x-ray tube, it does provide some idea of the benefits of compensation. Figure 12 shows images of the phantom with liver inserts in place for the 1:1 and 8:1 compensators, as well as profiles across the images. It is evident from these profiles that the cupping artifact due to scatter is greatly reduced by the compensator. With these imaging conditions the degree of cupping, given by6 tcup = edge − center ⫻ 100 % , edge 共2兲 was found to be 12.5% in 1:1 images and −1% in 8:1 images 共negative because of the presence of beam hardening capping artifacts兲 acquired with a FOVz of 18.5 cm. With a 2 cm FOVz the cupping was 3% in the 1:1 case and −0.6% in the 8:1 case. Measurements of the CNR when imaging with a 2 cm FOVz give values at the center of the phantom of 2.6± 0.2 and 2.4± 0.2 for the 1:1 and 8:1 compensators, respectively, and at the edge of the phantom of 3.1± 0.2 for the 1:1 compensator and 2.7± 0.2 for the 8:1 compensator. So, for small longitudinal FOVs compensators do not offer any improvements in image quality. If the FOV is increased, there is a larger amount of scatter, and the CNR in the 1:1 case drops to 1.9± 0.2 at the center of the phantom and 2.5± 0.1 at the periphery of the phantom. The CNR for the 8:1 compensator was measured as 2.4± 0.1 at the center and 2.3± 0.2 at the periphery. The 8:1 compensator does not imMedical Physics, Vol. 34, No. 7, July 2007 prove the CNR at the edge of the phantom compared to the 1:1 compensator when imaging under high scatter conditions, but this is expected since the 8:1 compensator greatly reduces the primary radiation at the periphery of the phantom. The improvement in CNR from the 8:1 compensator is seen at the center of the phantom where the CNR increased from 1.9 up to 2.4 when compared to the 1:1 filter. IV. DISCUSSION The influence of modulating filters on both dose and scatter in cone-beam CT was evaluated. The application of compensating filters for all choices of phantom, air gap, and field of view resulted in decreased scatter and dose, with the compensators removing over half the scatter and reducing the dose at the center of the phantom by nearly a factor of 2 in some cases. One known limitation in this work is the presence of extra-focal radiation, which cannot be separated from x rays scattered from the imaged objects. Potential future investigations will look to Monte Carlo methods for separating these effects and verifying the trends seen for the DRF and SRF. As scatter is likely the largest factor in the reduction of image quality in CBCT many approaches are being investigated to address it. This management may take the form of algorithms applied after imaging, or as physical modifications to the imaging system that reduce the amount of scat- 2701 Graham et al.: Compensators for cone-beam CT 2701 FIG. 11. Reconstructions of the cylindrical water phantom with a FOVz of approximately 18.5 cm for the 1:1, 4:1, and 8:1 compensating filters. tered radiation reaching the detector. Scatter correction algorithms could rely on Monte Carlo simulation methods or estimations of the scatter based on experimentally acquired scatter fluences.36 Some of the physical modifications that could be made to a cone-beam CT imaging system for the purpose of reducing scatter include adjusting the air-gap between the patient and detector, reducing the longitudinal field-of-view employed during imaging to only image the clinically relevant anatomy, utilizing anti-scatter grids to reject scatter, and, as evidenced by this work, application of compensating filters. This work gives indication of the magnitude of scatter reduction that can be accomplished with simple compensators, which could be combined with other methods of scatter reduction to achieve more quantitatively accurate and artifact-free reconstructions. An added benefit of compensators is that, unlike most of the methods for removing the effects of scatter in CBCT imaging, compensators also have the ability to modulate the primary and scatter dose delivered during image acquisition. Although the compensators are shaped to account for the shape of a head phantom, and not for the body phantom, the decrease in image quality associated with the decrease in primary fluence at the periphery of the phantom would not be a problem if, for example, the anatomy at the center of the patient volume is the most clinically relevant. Similar to region of interest imaging, this is one example of how comMedical Physics, Vol. 34, No. 7, July 2007 pensators could be employed while taking into account the various factors that they are capable of influencing. Images acquired of a cylindrical water phantom have shown that for compensators with considerable modulation factors there is a possibility of introducing considerable artifacts into the images due to the spectral hardening of the x-ray beam. Although the compensators may reduce scatterinduced artifacts, the distortion of the x-ray energy spectrum across the FOV introduced by the compensators may be more of a problem than the cupping from scatter. Beam hardening corrections have been investigated for CBCT systems,37,38 though beam hardening corrections generally correct spectral changes caused only by the patient, and not by compensating filters that are not present in the reconstructed images. Implementation of beam hardening corrections for cone-beam CT that can account for the large spectral changes introduced by compensators with considerable modulation factors may be necessary in order to use compensators with larger modulation factors. Further investigation into the selection of materials and compensation methods that minimize the spectral perturbations caused by the compensator is required. Despite the fact that the compensators employed here induce artifacts, there is still great potential for the use of compensators to reduce both the scatter at the detector and the patient dose. Compensators with lower modulation factors than those discussed in this work are a 2702 Graham et al.: Compensators for cone-beam CT 2702 tered x rays. The challenge and next logical step of this work is to seek the appropriate compromises that still satisfy the clinical requirements. V. CONCLUSIONS In this study, two representative compensators have been used to demonstrate some of the capabilities of modulating filters in cone-beam CT. Large reductions in scatter and patient dose could be made, as well as increased uniformity in reconstructed images, a decrease in cupping artifacts in reconstructed volumes, and an increase in CNR at the center phantoms imaged under high-scatter conditions. All of these results are expected to be similar to those that could be achieved with compensators of different shape or material. One complicating factor that requires more investigation is the introduction of beam-hardening artifacts into reconstructed images. One possibility for the design of a compensator would be to optimize the compensator such that it introduced the least amount of spectral hardening artifacts. An approach that would potentially be more beneficial to patients would be that the attenuation profile of compensating filters could be optimized to account for numerous properties of an imaging system. Depending on the desired imaging task, various aspects of the cone-beam CT imaging system, such as the magnitude of scatter, dose, and distribution of primary fluence, can be influenced by compensating filters. Further investigations into the possibility of creating optimal fluence patterns with compensators to be utilized during CBCT imaging continue to be an active area of research. ACKNOWLEDGMENTS The authors wish to acknowledge the radiation physics machine shop at Princess Margaret Hospital for construction of the compensators, and Bronwyn Hyland for assistance in editing the manuscript. This project was supported by the Ontario Graduate Scholarship program, the National Institutes of Health/National Institutes of Aging 共R21/R33AG19381兲, and the National Institutes of Health/National Institute of Biomedical Imaging and Bioengineering 共8R01EB002470-04兲. FIG. 12. Reconstructions of the cylindrical water phantom with a FOVz of approximately 18.5 cm for the 1:1 and 8:1 compensators. Liver equivalent soft-tissue inserts have been placed at two locations inside the phantom to allow measurement of the contrast-to-noise ratio. possible solution leading to a compromise between beamhardening artifacts and the removal of scattered x rays from the system. Cone-beam CT is used in a number of clinical systems, including linear accelerator and C-arm based CBCT systems. While the performance of these systems is sufficient to be valued, image quality is still recognized as an area requiring improvement. The results of this work demonstrate that compensators, if applied carefully, can play a role in reducing patient dose and increasing CNR through reduction of scatMedical Physics, Vol. 34, No. 7, July 2007 1 “Management of target localization uncertainties in external-beam therapy,” Seminars in Radiation Oncology, edited by G. S. Mageras, Vol. 15, pp. 133–216 共2005兲. 2 D. A. Jaffray, J. H. Siewerdsen, J. W. Wong, and A. A. 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