Deformable Scintillation Dosimeter II: Real-Time Simultaneous Measurements of Dose and Tracking of Deformation Vector Fields
DDeformable Scintillation Dosimeter II: Real-TimeSimultaneous Measurements of Dose and Trackingof Deformation Vector Fields
E Cloutier , , L Beaulieu , ,L Archambault , ‡ Physics, physical engineering and optics department and CancerResearch Center, Universite Laval, Quebec, Canada. CHU de Quebec - Universit´e Laval, CHU de Quebec, Quebec,Canada
Abstract.
Anatomical motion and deformation pose challenges to the understandingof the delivered dose distribution during radiotherapy treatments. Hence, deformableimage registration (DIR) algorithms are increasingly used to map contours and dosedistributions from one image set to another. However, the lack of validation toolsslows their clinical adoption, despite their commercial availability. This work presentsa novel water-equivalent deformable dosimeter that simultaneously measures the dosedistribution and tracks deformation vector fields (DVF). The dosimeter in made of anarray of 19 scintillating fiber detectors embedded in a cylindrical elastomer matrix. Itis imaged by two pairs of stereoscopic cameras tracking the position and angulation ofthe scintillators, while measuring the dose. The resulting system provides a precisionof 0.3 mm on DVF measurements. The dosimeter was irradiated with 5 ×
3, 4 × × ‡ Present address: Department of Physics, Universit´e Laval, Quebec, QC BS8 1TS, Canada. a r X i v : . [ phy s i c s . m e d - ph ] J a n eformable Scintillation Dosimeter II : Dose and DVF measurements
1. Introduction
Advances in modern radiotherapy treatment techniques have led to the advent ofcomplex personalized treatment plans aimed at maximizing the dose delivered to thetumor while minimizing the dose delivered to surrounding tissues. Treatments plansare personalized to the patient’s anatomy, resulting in dose gradients close to thetarget. However, over the course of treatments, the patient’s anatomy may be deformedand/or change in volume. These anatomical variations challenge the understandingof the cumulative dose delivered throughout the course of radiotherapy treatments[1]. Hence, deformable image registration (DIR) algorithms are increasingly used inthe clinics to either map organ contours or dose distribution from one image set toanother[2]. However, in low contrast tissues, the high number of degrees of freedomof these algorithms can lead to inaccuracies in the computed deformation vector field(DVF) [3–5]. Using those DVFs would result in incorrect voxel pairing, leading toerrors in dose accumulation. Thus, the American Association of Physicist in MedicineTask Group 132 on the use of image registration algorithm in radiotherapy (TG-132)recommends that end-to-end tests should be performed using quality assurance (QA)phantoms prior to the implementation of these systems in the clinics [6]. In spiteof the these recommendations, the definition of a patient-specific gold standard DIRvalidation tool remains an open issue [7]. Amongst the proposed validation tools,physical phantoms benefit from their ability to test the entire registration process, fromthe image acquisition to the registration itself.Deformable dosimetric gels have shown potential in measuring three-dimensionaldose distributions delivered to deformable targets [8–11]. These water-equivalent gelsdemonstrated robust reproducibility and spatial resolution up to 1 mm [12]. However,they are integrating dosimeters and thus can only provide information on the cumulativedose deposited. Some anthropomorphic phantoms were also developed using landmarksto measure solely the deformation, not the dose [13, 14]. Some deformable phantomswere further developed with enclosures to insert ion chambers, radiochromic films orMOSFETs, for dose measurements [15–17]. However, the non-water equivalence of thesedosimeters limits the practical number of simultaneous measurement points as somedetectors can disturb the dose deposition pattern. Moreover, the contrast associatedwith these detectors may bias DIR validation in homogeneous mediums since it couldbe interpreted as fiducial markers in the images by the algorithms.On the other hand, work on volumetric scintillation detectors has shown thefeasibility of real-time dose measurements over whole 2D and 3D volumes [18–24].Those systems provide millimeter resolution and water-equivalent measurements, butwas limited to fixed measurements. As scintillators possess essential dosimetric qualities[25], they may constitute an ideal candidate for the sensitive volume of a volumetricdeformable dosimeter [26]. Such a dosimeter could be suited for both the challenges ofmotion management and advanced radiotherapy modalities.This work presents the development of a novel scintillator-based deformable eformable Scintillation Dosimeter II : Dose and DVF measurements (a)(b) (c) CCD1CCD2SCMOS1SCMOS2 Dosimeter
Figure 1: Representation of the developed dosimeter and its experimental set-up. (a)Clear deformable elastomer matrix. (b) Dosimeter composed of the elastomer matrixwith 19 scintillating fibers embedded. (c) Experimental irradiation setupdetector that simultaneously measures the dose distribution and tracks deformationvector fields at 19 positions.
2. Methods
The dosimeter consists of 19 scintillators embedded in a clear, water-equivalentelastomer (figure 1b). The elastomer (Clearflex30: Smooth-On, Macongie, USA) wascast in a silicone cylindrical mold (diameter: 6 cm, thickness: 1.2 cm) and the compoundwas degassed to ensure an optimal transparency of the bulk. Physical properties of theelastomer are listed in table 1.Table 1: Physical properties of the clear plastic matrix provided by the manufacturer.
Density Refractive index Tensile strength Elongation at break Shore hardness[g/cm ] [-] [psi] [%] [A]1.03 1.486 725 675 30 After pouring the gel, 19 polyethylene terephthalate (PET) tubes (Nordson medical,Salem, USA) were inserted in the elastomer guided by a 3D printed template. Oncethe elastomer set, the holder was removed, leaving an array of 19 hollow tubes in thecylindrical gel matrix, as can be seen on figure 1a. The hollow tubes have an internal eformable Scintillation Dosimeter II : Dose and DVF measurements ± ± ± AB CD E
A: BCF-60 scintillator - 1.00B: PET tubing 1.1 2.16C: Polyester tubing 2.16 2.25D: PET tubing 2.44 2.69E : Elastomer - 60A 1 cm vertical compression was applied to the dosimeter in the antero-posteriordirection. The dosimeter was inserted between two plastic plates distant by 6 cm (fixed)and 5 cm (deformed state). The plates were brought closer with two tighten nylon screws(figure 1c).
The dosimeter was simultaneously imaged by 4 cameras as depicted on figure 1c.As scintillating fibers emit light in proportion to the dose deposited in their volume,collecting this signal provides information on the dose delivered as well as the scintillatorslocation in the phantom. The cameras were arranged to form two facing stereoscopicpairs. Therefore, the setup enables the 3D position tracking of both ends of eachscintillator. All cameras were coupled to 12 mm focal length lenses (F/ eformable Scintillation Dosimeter II : Dose and DVF measurements ( θ ( i,j ) ) fit [27]. The stereoscopic pair was calibrated using a(15 ×
10) grid chessboard pattern and a calibration algorithm inspired by Zhang fromthe OpenCV python library version 3.4.2 [28, 29]. The scintillation signal was correctedaccording to their angle and distance from the CCD’s sensor center (figure 2). A detaileddescription of this process is provided in the companion paper [26]. The cameras wereshielded with lead blocks to reduce noise from stray radiation.
The dosimeter was irradiated with a 6 MV, 600 cGy/min photon beam (Clinac iX,Varian, Palo Alto, USA). The signal-to-noise ratio (SNR) and signal-to-backgroundratio (SBR) of the detector were studied while varying the dose delivered and the doserate. Signal-to-noise ratio describes the system’s sensitivity and was defined as the ratioof the mean pixel value to its standard deviation for each scintillation spot [30]. Signal-to-background was defined as the ratio of the signal to the standard deviation of thebackground and describes the signal’s detectability.
SN R ave = µ s σ s , SN R spot = √ nSN R ave , SBR = µ spot σ bg (1)Different instantaneous dose rates were achieved by varying the distance between thedetector and the irradiation source, keeping the delivered monitor units and linacsettings constant.Each fiber was dose-calibrated by irradiating the phantom with a 6 × field sizeand monitor units (MU) ranging from 3 to 10 MU. The phantom was centered at the eformable Scintillation Dosimeter II : Dose and DVF measurements Probe Reader
Figure 3: Picture illustrating the Hyperscint customized probe in the deformabledosimeter and its reader. Dose measurements were performed at the location of thefive encercled scintillators.isocenter of the linac. Reference dose calculation was performed using a treatmentplanning system (Raystation; RaySearch laboratories, Stockholm, Sweden). Dosecalculations were performed with a 1 mm dose grid. These measurements enabledthe light to dose conversion and assessed the linearity of the detector. Then, thedeveloped dosimeter was used to measure the dose distribution and the deformationvector field resulting from a deformation. The dosimeter was imaged and irradiatedin both states, i.e. fixed and deformed, with 5 ×
3, 4 × × field sizes.Dose measurements were validated and compared using an independent scintillationdosimetry system (Hyperscint; MedScint Inc., Quebec city, Canada). Dose measurements previously described werereplicated using the Hyperscint scintillation dosimetry research platform. This providedan independent validation of the dose delivered at the location of five chosen scintillators(figure 3). A custom manufactured scintillating probe was inserted in the dosimeter atthe selected location (replacing the 1.2 cm long scintillator described in section 2.1). Thescintillator in the probe has a length and diameter of 1.2 cm and 1 mm respectively,resulting in the same sensitive volume as that of the scintillators used in the deformabledosimeter. The external diameter of the probe matched the internal diameter of theplastic tubing. However in this case, the scintillator was coupled to a 20 m long clearoptical fiber to guide the light to a photodetector, thus enabling traditional plasticscintillation dosimetry (PSD) measurements [25]. The system was calibrated at theisocenter of a 10x10 cm field, at a depth of 1.5 cm in a solid water phantom (SSD =98.5 cm). Cerenkov stem signal was corrected using the hyperspectral formalism [31,32]. The scintillation spectrum was measured from a kV irradiation. Cerenkov spectrumwas acquired from two MV measurements for which the dose at the scintillator was keptconstant: 1) minimal ( C min ), and 2) maximal ( C max ) clear fiber was irradiated in thebeam field [33]. The cerenkov spectrum results from the subtraction C max − C min . Figure4 summarizes the workflow of the experimental measurements. eformable Scintillation Dosimeter II : Dose and DVF measurements Camera calibrationDose measurement : • Gantry 0° - 5x3, 4x3, 3x3 cm Compression release Hyperscint dose measurement : • Gantry 0° - 5x3, 4x3, 3x3 cm Hyperscint dose calibration • Spectra acquisition (scintillation, Cerenkov) • Absolute dose calibration
Dosimeter compressionCT Scan Deformed stateFixed stateCT ScanDose measurement : • Gantry 0° - 5x3, 4x3, 3x3 cm Dosimeter compressionHyperscint dose measurement : • Gantry 0° - 5x3, 4x3, 3x3 cm Figure 4: Workflow of the dose measurements and tomographic images acquisition.
Deformation vector fields (DVF) were measured using the dosimeter by tracking thesurface centroid of each scintillating fibers, from both sides. Thus, 19 vectors aremeasured indicating the direction and magnitude of the fiber displacements between thefixed and deformed conditions. Stereo-vision enabled the detection of the 3D positionof both fiber ends in the two studied cases (fixed vs deformed). Angulation of the fiberswere extracted from the displacement differences measured by the facing stereoscopicpairs.The dosimeter was CT-scanned (Siemens Somatom Definition AS Open 64, SiemensHealthcare, Forchheim, Germany), for both conditions. The pitch, current, tube current-time and energy of the scanner were respectively set to 0.35, 60 mA, 1000 mAs and 120kVp. The CT images were further fed to a DIR algorithm and the computed DVF wasextracted.The B-Spline algorithm from Plastimatch [34] was used to compute the DVFdescribing the transformation mapping the fixed dosimeter state to its deformed state.The algorithm’s cost function is guided with image similarities using pixel’s meansquare error (MSE). The regularization term, i.e method to ensure physically realisticdeformation scenarios, was set to 0.005. The resulting deformation vector field, obtainedoptically and from the deformable image registration algorithm, were compared.Reproducibility of the deformation and hysteresis of the dosimeter were eformable Scintillation Dosimeter II : Dose and DVF measurements
3. Results
Calibration of the detector lead to an expected linear dose-light relationship ( R > >
5) and detectability (SBR >
2) thresholds for all the explored doses anddose rates (figure 5). Points and error-bars on figure 5 represent respectively the meanand standard deviation of the 19 fibers. Table 3 presents the position reproducibility of M e a n S N R [ - ] Dose at isocenter [cGy]10 M e a n S B R [ - ] Dose rate at isocenter [cGy/min]
Figure 5: Signal-to-noise ratio (SNR) and signal-to-background ratio (SBR) as afunction of the dose deposited and the dose rate at the isocenter. Dashed lines representcut-off values for accurate detectability. Error-bars indicate the range of values obtainedfor the 19 scintillating fibres rather than the uncertainty on the measure.the 19 scintillators in the fixed and deformed states. Variations in the position of thescintillators (mean ± standard deviation) are also listed. The higher variations wereobtained on the z (depth) axis, but remained under 0.3 mm : the precision of the 3Dtracking by the cameras. Hence, the deformation was reproducible and the elastomerdid not present hysteresis.To complete the dosimeter’s characterization, a mean density of 1.06 ± was extracted from the CT-scan images, which corroborates its water equivalence. Figure 6 presents the 3D deformation vector fields obtained from the scintillation signal(a, b) and from the Plastimatch deformation algorithm (c). Differences in the DVFmeasured from both ends (front : figure 6a, back: figure 6b of the elastomer by the eformable Scintillation Dosimeter II : Dose and DVF measurements ± standard deviations over the 19points. Fixed Deformed Y p o s i t i o n [ c m ] FixedDeformed Z p o s i t i o n [ c m ] x [cm] 0.03 ± ± ± ± ± ± Différence mesure vs DIR (c)(a) (b) (c)
Figure 6: Deformation vector fields measured from scintillation on the first (a) and last(b) dosimeter surfaces, and computed with the Plastimatch DIR algorithm (c). Thegrey circle represents the dosimeter contour in its fixed state.Globally, the DVF computed by the DIR algorithm presents the same shape, andmagnitude as the one obtained optically. Overall, the applied compression resulted ina downward shift in the vertical axis and a shift towards the edges in the horizontalaxis. Moreover, the compressed dosimeter develop a convex shape towards the cameras(CCD1 and sCMOS1) as a result of the applied deformation. The curve was opticallydetected by the depth (Z) variation in the 3D tracking. The largest vertical deformationwas obtained at the top of the dosimeter with measured and computed displacement of6.7 ± ± eformable Scintillation Dosimeter II : Dose and DVF measurements F r e q u e n c y [ - ] X 1 0 1Difference [mm]Y 1 0 1Z
Figure 7: Distribution of differences between the measured and computed DVF in thex, y and z axis, respectively.
Dose distributions were acquired in fixed and deformed conditions. The compressionof the deformable dosimeter led to movement, i.e. translation and rotations, of thescintillators. Signal was accordingly corrected to account for variations in the systemcollection efficiency. Figure 8 presents the angular, distal and vignetting corrections thatwere applied to each scintillator. Angulation and distance from CCD1’s sensor center
Facteurs de correction sur la dose
123 4567 89101112 13141516 171819
Figure 8: Angular, distal and vignetting correction factors applied to the deformedmeasurements for each detecting scintillator. The scintillator labels are defined on theright figure.were measured. The angulation correction coefficient results from the combined tilt ofthe scintillators in the elastomer and their position relative to the camera. Deformingthe dosimeter led to tilts of the fibers as presented on figure 9. Measured θ presents asymmetry along the x-axis, as expected.Dose distribution from the 19 scintillators are presented on figure 10, for differentfield sizes. For each field, crossline profiles and depth dose were extracted and comparedwith the Hyperscint measurements and computation from the treatment planningsystem (figure 11). An uncertainty of 1% was estimated on scintillators measurementswhich mainly takes into account the correction factors uncertainty. Uncertainties onTPS calculations corresponds to dose variations resulting from 1 mm translations toaccount for setup variations, whereas the uncertainties on the Hyperscint correspond eformable Scintillation Dosimeter II : Dose and DVF measurements -14.836.29-15.172.31-14.86-0.03 -8.386.39-10.162.6-10.15-0.91-9.17-1.74 -0.375.36-0.033.3-0.5-0.78-0.39-4.0-0.04-1.98 7.642.639.55-0.439.06-4.017.94-4.47 13.13-1.3714.27-4.3812.59-6.57 Measured [°]Measured [°]
Figure 9: Angulation of the fibers in the φ and θ direction.to the standard deviation over 10 consecutive measurements. For depth dose and FixedDeformed 4x3 cm Figure 10: Dose distributions measured from 19 scintillating points for various fieldsizes in the fixed (top) and deformed (bottom) dosimeter’s states. Each colored circlerepresents the dose measured with a scintillator. Size of the circles is larger than theactual size of scintillators for better visualisation.profiles, most differences between scintillators measurements and TPS calculationsremained within the uncertainty margins of 1%. In the beam direction, deformationof the dosimeter results in dose shifts along the depth dose line as scintillators werebrought closer to the surface. Scintillators towards the sides of the dosimeter exhibitlarger variations between the fixed and deformed conditions. Differences between thefixed and deformed conditions up to 37 cGy (60%) were obtained, which refers to ascintillator moving through the beam’s edge following deformation. It was calculated eformable Scintillation Dosimeter II : Dose and DVF measurements × profile, were likely caused by a 0.9 mm re-positioning shift, when the probe was insertedin the dosimeter. Figure 11: Scintillation dose measurement comparison with calculation from thetreatment planning system (TPS) and Hyperscint measurements. Left images presentdepth dose as right images present dose profiles.
4. Discussion
We developed a novel real-time deformable dosimeter that can simultaneously measuredose and deformation vector fields with a system of cameras. Using plastic scintillators,we were able to develop a water-equivalent phantom compatible with most imagingmodalities. In addition, given the dosimeter’s density homogeneity, the scintillatorsdo not act as fiducial markers and allow the evaluation of deformable registrationalgorithms without influencing their outcomes. However, measuring the light outputfrom displaced scintillators with fixed cameras created new challenges. Hence, it wasdemonstrated that such system requires precise position and orientation tracking of thescintillators to account for signal variations arising from changes in their optical couplingwith the cameras [26]. In this case, compressing the dosimeter by 1 cm necessitatedcorrection factors of up to 5.6%. As displacements of the fibers were lower than 0.71 cm, eformable Scintillation Dosimeter II : Dose and DVF measurements × field. The differences between the dose measurements and TPS reached2.2%. Other differences between Hyperscint and scintillator measurements remainedunder the positioning uncertainties. Overall, agreement with the TPS was expected asscintillators were calibrated against calculation from the TPS itself, but with differentirradiation conditions than the ones used for analysis. Ideally, the system should becalibrated independently from the TPS. However, dose calibration of the system remainstricky because each detecting scintillator needs to be individually calibrated, to accountfor variations in the polishing for example, and the phantom doesn’t provide sufficientscattering conditions for AAPM TG-51 reference dose calibration [35]. As such, usingan external dosimetry tool, like a standard PSD dosimeter, to calibrate rather thanvalidate the system would be an interesting avenue.Deforming the dosimeter with an antero-posterior compression resulted in two maindosimetric effects : 1) along the depth dose, dose to scintillators increased as they werebrought closer to the surface and 2) the deformation increased the off-axis distance ofscintillators which resulted in dose decrease for scintillators moving from in-field towardsthe beam penumbra. 4 × field profile measurements especially stressed the needfor accurate understanding of the deformation as small shift near dose gradients canresults in significant dose differences. In that case, 1 mm lateral shifts could resultin dose differences up to 40 cGy as the scintillator is close to the beam’s edge. Theincreased complexity of modern radiotherapy techniques, such has IMRT and VMATtype deliveries, further enforces the need for efficient and quantitative dose distributionmeasurements [36]. Similarly, previous work have demonstrated that a small discrepancyin the computed DVF can significantly impact the warped dose, especially in highgradient regions, highlighting the need for validation [1].AAPM Task Group 132 stated that an ideal DIR validation tools should enablean error detection smaller than the DIR pixel size [6]. In our case, the tomographicimages were acquired with an in-plane pixel size of 0.35 mm. The set of stereoscopicpairs of cameras provided an optical measurement of the deformation vector field witha previously demonstrated precision 0.3 mm [26]. Hence, the system has the potentialto accurately portray deformation vector field for quality assurance applications. Thedeformation vector field computed with the DIR algorithm presented differences upto 1.5 mm with the one optically measured. AAPM TG-132 stated that an overallregistration accuracy within 2 mm is desired for clinical applications [6]. Scan quality,image artifacts and image distortions, amongst others, can affect the resulting qualityof a registration. In this work, scan quality was optimized with a tube current-timeof 1000 mAs. Contrast was further enhanced by choosing a head scanning protocol.Nevertheless, a computed DVF with Plastimatch present differences from the onepredicted by the optical measurements. Those differences are attributed to the known eformable Scintillation Dosimeter II : Dose and DVF measurements
5. Conclusion
Anatomical motion and deformation challenge the calculation of the dose delivered,raising the need for adapted quality assurance tools. We developed a dosimeterthat enables measurements in fixed and deformable conditions, while tracking thedeformation itself. The water-equivalent composition of the dosimeter further endows itwith the quality to act both as a phantom and detector. Moreover, the detector allowsa wide variety of 2D and 3D geometric or anthropomorphous designs since its shapeand size is solely determined by the mold used to cast the elastomer. Such a detectorcould be used for the quality assurance of DIR algorithms and to explore the dosimetricimpact of organ deformations.
6. Acknowledgement
We thank Medscint, especially Benjamin Cˆot´e and Simon Lambert-Girard, for theirsupport and for kindly providing a customized probe as well as the Hyperscint researchplatform for the measurements. We also thank Jonathan Boivin and `Eve Chamberlandfor their assistance in CT image acquisition and dose calculations, respectively. Thiswork was financed by the Natural Sciences and Engineering Research Council ofCanada (NSERC) Discovery grants
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