Radiation Protection Dosimetry (2014), Vol. 158, No. 4, pp. 389 –398 Advance Access publication 25 October 2013
doi:10.1093/rpd/nct236
MEASUREMENTS OF COMPUTED TOMOGRAPHY DOSE INDEX FOR CLINICAL SCANS D. Trevisan1,*, D. Ravanelli2 and A. Valentini1 1 Department of Medical Physics, S. Chiara Hospital, APSS Trento, Italy 2 Medical Physics, University of Rome Tor Vergata, Rome, Italy *Corresponding author: [email protected]
Dose assessment in computed tomography is nowadays based on indicators such as the weighted computed tomography dose index (CTDIw) and the volume-weighted computed tomography dose index (CTDIvol ), both measured only on single-axial scans. The aim of this study is to extend the set of CT protocols suitable in CTDIvol,w evaluation and therefore to perform measurements directly on clinical CT scans. With this purpose, the present work follows the methodology proposed by the American Association of Physicists in Medicine in the Report No. 111 and focuses on the central cumulative dose DL(0), which predicts CTDIvol,w values for scan length equal to 100 mm. All measurements performed with ion chamber and gafchromic films suggest that it is possible to achieve accurate CTDIvol,w values without selecting single-axial scans tailored to dosimetry. Therefore, it is not always necessary to split dosimetric and clinical CT scanner set-up.
INTRODUCTION Dose assessment in computed tomography (CT) requires easily measurable indicators of dose delivered to organs–tissues. Weighted computed tomography dose index (CTDIw) and volume-weighted CT dose index(1, 2) (CTDIvol ) are nowadays valuable tools for this task, even though they seem to be obsolete because these are always measured only on single-axial scan tailored to CTDIw requirements(3 – 8). Moreover, the relation CTDIvol ¼ CTDIw/pitch used in this approach to extend dose assessment to helical scans is not always of easy implementation. This trouble occurs when some scan parameters with strong influence on CTDIw, such as the beam collimations, are not selectable on axial CT scans. Additionally, dosimetry based on axial scans can be also error prone whenever the presence or removal of different bow-tie filters is not displayed at CT console. In both the cases, CTDIw,vol assessments cannot be performed in an accurate manner. To overcome such limitations, the AAPM TG 111 proposes an alternative dose assessment methodology(9) focusing on the cumulative dose profile D(x)(10 – 12) for sequential as well as for helical CT scans. With reference to Figure 1, D(x) shows in the first case always a quasiperiodical behaviour of fundamental period k equal to the couch position displacement between two consecutive X-ray exposures. Conversely, oscillatory dose distribution is evident for helical scans only on the peripheral axes of the polymethylmethacrylate (PMMA) phantoms, where the periodicity k is the product of pitch and the nominal beam width. Dixon defines also the central cumulative dose DL(0)(10) as the value of D(x) at the centre (x ¼ 0) of the scan with length L. In particular, DL(0) of a helical CT scan with pitch p, if measured
in the centre of the PMMA phantom, comes from a periodicity free profile, and it is in relation to CTDIvol as follows: pDL ð0Þ ¼ CTDI100 for L ¼ 100 mm.
ð1Þ
Moreover, computing the averaging of peaks and valleys of D(x), Equation (1) can be generalised to every kind of CT protocols and locations in the PMMA phantoms. This averaging procedure can be calculated as convolution between D(x) and the rectangular function p(x/k)/k before taking its value at the central scan position (x ¼ 0)(9). As a consequence, Equation (1) is still valid for helical as well as for sequential scans (in this case, p is replaced by 1), and therefore: 1 c 2 p CTDIvol;w ¼ D100 ð0Þ þ D100 ð0Þ; 3 3
ð2Þ
where D100(0) is DL(0) for L ¼ 100 mm. Suffixes c and p indicate central and peripherals locations, respectively, in PMMA standard phantoms. The aim of this study is to overcome the limitations of the CTDIw-based approach, estimating CTDIw as well as CTDIvol directly on clinical CT scans by means of D100(0) measurements as indicated in Equation (2). In line with the AAPM TG 111 approach, all measurements were performed with a short ion chamber(9) (IC). The purpose of the present work is also to investigate the feasibility to use Gafchromic films (GF) for the same D100(0) measurements. The use of GF for CT dosimetry was already investigated by Rampado
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Received 8 May 2013; revised 3 September 2013; accepted 4 September 2013
D. TREVISAN ET AL.
et al.(13), obtaining a discrepancy from CTDIw measured with pencil chamber (PC) always of ,9 % for collimation of .10 mm. GF, unlike IC, have the advantage of choosing the region of interest (ROI) exactly of length k equal to the periodicity in D(x). Length k can change significantly among protocols and scanners, yielding different kind of ripples. Therefore, GF reduce, especially for wide collimation and pitch scans, all ripple-correlated uncertainties due to D(x) quasi-periodical behaviour. To validate the D100(0)-based methods described above, the present work compares all CTDIvol,w values to ones achieved by the standard approach(1, 2). In addition, as required in the study by International Electrotechnical Commission(14), it is also tested whether the relative disagreement between nominal and measured CTDIvol,w is always ,20 %. MATERIALS AND METHODS Dosimetry equipment At first, CTDIw measurements were conducted following exactly the standard dosimetric approach(1, 2) (Method 1), i.e. by means of head and body standard PMMA phantoms (diameter 160` 320 mm) and an 10` 5–3CT PC with charge-collection-length of 100 mm coupled to an Radcal 9010 electrometer (Radcal Corporation, CA, USA). DL(0) measurements were then performed on head and body extended (respectively to 200 and 300 mm length) PMMA phantoms as suggested by J.C. Martin(15). DL(0) was measured at first by a 10`5–0.6 IC (Method 2) with charge-collection-length of 20 mm
coupled to the same electrometer used in Method 1. The same procedure was then repeated using GF XR-QA2 (International Specialty Products, Wayne, NJ, USA, lot number A10121202, expiration date 31 October 2014). This last approach (Method 3) required also the preliminary calibration-set-up of the whole GF dosimetry system consisting of GF and a commercial flatbed scanner (Epson Perfection V750 Pro). Chromatic changes in GF were read by the flatbed scanner providing digital imaging of GF; all measurements on the images were performed by the use of ImageJ (Wayne Rusband, National Institute of Health, USA, http://rsb.info.nih.gov//ij version 1.44o) homemade plug-in. All experimental data were analysed by means of the Matlab 7.0 statistical tool. CTDIw obtained by Method 1 and nominal CTDIw,vol were used to benchmark Methods 2 and 3 proposed in this study. The current investigation involved performing measurements on a Phlips Brilliance 64S, (Tomograph A), a Siemens Biograph 64S (Tomograph B) and a Siemens Somatom Definition AS 128S (Tomograph C). The methods proposed in the present work were at first tested on Tomographs A and B. The advantages of Methods 2 and 3 were pointed out in particular on Tomograph C.
Dosimetry system set-up Ion chamber calibration Both PC and IC used in the present work, respectively, in Methods 1 and 2 had been previously calibrated in
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Figure 1. Typical cumulative dose profile DL(x) for axial and helical scans in different locations of the cylindrical PMMA phantoms. Its ripple is characterised by periodicity k.
DOSIMETRY ON CT CLINICAL SCANS
free-air for the RQA 9 beam according to IEC 61267 report (16). The calibration factor obtained in free-air was also used for both CTDIw and DL(0) in phantom measurements. This assumption is justified by published data of the energy dependence of IC(17) and by the limited effect of the beam spectral variations due to PMMA phantoms(13, 18). As a consequence in the present work IC provided a direct measurement of DL(0).
It is notable that Equation (4) fits saturation effects of GF and at the same time NetPV ¼ 0 in the absence of exposition. Scan length correction factor h 100(L) Clinical CT scans generally do not allow setting L ¼ 100 mm. The D100(0) values reported in Equation (2) were therefore estimated by the use of the correction functions h 100(L). These last describe the relative approach of DL(0) to D100(0) as function of L as:
Scanner set-up and GF calibration
PVunexp 1: NetPV ¼ PVexp
ð3Þ
NetPV values were then fitted with the following equation(13): Deq;air ¼ aNetPV2 þ bNetPV:
ð4Þ
D100 ð0Þ ¼ h100 ðLÞDL ð0Þ:
ð5Þ
h 100(L) were experimentally determined in line with AAPM Report no. 111(9). These correction factors were estimated only on Tomographs A and B, but not on Tomograph C. This choice is justified because the two scans on this last tomograph are always performed at L ¼ 100 mm (see Table 1), and therefore, the direct measurements of D100(0) were carried out. Operatively h 100(L) were obtained by means of central cumulative dose measurements repeated in central as well as in peripheral positions of both PMMA phantoms for scan length L ranging from 40 mm up to 160 mm. Experimental values, indicated in the present section as Dref(L) in order to emphasise their dependence of L, were measured by IC for the following scan parameters: 120 kVp, rotation time 1 s, pitch 0.68, nominal collimation 10 mm, 500 mA s/slice, FOV 300 mm on Tomograph A, and 120 kVp, rotation time 1 s, nominal collimation 19.2 mm, 500 mA s/slice, FOV 300 mm on Tomograph B. Dref(L) values were then fitted with the function: L : ð6Þ ¼ A 1 B exp Dref fit C with fit parameters A, B and C. In the present work, h100(L) functions were then obtained normalising ref Dref fit ðLÞ to Dfit ð100Þ as: h
100
1 Dref L fit ð100Þ ð0Þ ¼ ref : ¼ D 1 Bexp C Dfit ðLÞ
ð7Þ
with D ¼ 1 B expð100=CÞ: Dosimetry of CT clinical scans All CT protocols focused in the present study (always at 120 kVp X-ray tube voltage) are of wide use on patients and show very different scan parameters. All these are reported in Table 1, as well as the phantom type used in dosimetry. To obtain the CTDIvol,w by means of Method 1, all scan parameters reported in Table 1, with the exception of the pitch values and scan lengths L, were
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Whereas PC and IC provided directly CTDIw and DL(0) values, GF used in Method 3 required the preliminary characterisation of the whole dosimetry system. With this aim, the flatbed scanner was tested beforehand in line with the procedure indicated in Rampado et al.(19). Repeatability was between 0.1 and 0.9 %; non-uniformity was 1.3 % and ,0.5 % for 75 % of the ROI (8.0` 20 mm). The second step of the dosimetry system characterisation-set-up involved the calibration of the GF. To investigate the relationship air kerma (Kair) vs GF change in colour, 80 unexposed GF (10` 60 mm) were at first placed on the flatbed central area(20). All GF were then scanned as indicated by Devic et al.(21). On each GF image (TIFF RGB mode, 16 bit per colour, red channel only, resolution 72 dpi, no enhancement, filter or colour correction factors), the average pixel value PVunexp was recorded on a 8` 50 mm ROI. The IC was then placed along the central axis of the gantry solidal to the CT couch of Tomograph A and irradiated in free-air with following CT scan parameters: helical mode 120 kVp, nominal collimation 32`1.25 mm, pitch factor 0.971, rotation time 2 s and scan length 160 mm. Sixteen different mA s/slice values ranging from 65 up to 600 were selected. Particular care was also taken in positioning the centre of the CT scan (x ¼ 0) exactly on the centre of the IC. Under these conditions, the Kair value provided by the IC coincided with the air equilibrium dose Deq;air (9, 22). Exactly the same CT scan parameters used in IC irradiation were then selected for the exposure of the 80 beforehand scanned GF. The exposed GF were read only 24 h after its irradiation(19), following exactly the same procedure described above for the PVunexp evaluation, obtaining PVexp values. The change in the colour of each GF was then described by:
D. TREVISAN ET AL. Table 1. Nominal CTDIvol,w and scan parameters selected for central cumulative dose measurements (always at 120 kVp). Protocol
Rotation time (s)
Nominal collimation (mm)
A
1.5
B A B A B A B A B C C
Pitchb
Nominal CTDIdvol,w (mGy)
Field of view (mm)
Tube current (mA)
Scan lengthc (mm)
10.0
230
333
100.0
95.2
(H)
1.0
28.8
230
380
115.2
52.8
(H)
0.75 1.0 0.75 1.0 0.75 0.5 1.0 0.5 0.5 0.5
25.0 19.2 1.0 7.2 40.0 19.2 40.0 19.2 38.4 38.4
230 230 150 150 350 350 350 350 320 220
360 304 100 112 238 280 243 464 320 160
99.9 99.9 49.2 100.2 149.1 119.3 151.1 121.0 100.0 100.0
56.7 59.4 70.9 32.5 12.9 7.6 16.2 12.3 13.5 5.4
(H) (H) (H) (H) (B) (B) (B) (B) (B) (B)
0.675 0.8 0.375 0.8 0.891 1.4 0.97 1.45 0.8 0.8
a
A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S. Sequential CT scans are considered equivalent to helical ones with pitch ¼ 1. c Scan length is affected by +0.5 mm uncertainty. d CTDIvol for helical and CTDIw for sequential CT scans. In brackets is the phantom type (H head, B body). b
Figure 2. Left: description of dosimetric measurement set-up: a lateral view with phantom and dosemeters placed at scan centre x ¼ 0. Right: transversal view with X-ray source, phantom and GF oriented parallel to external phantom surface (inset figure).
selected on single-axial scans. Measured CTDIw values were then divided by the nominal pitch value of the corresponding clinical helical CT scan(1, 2)
obtaining CTDIvol. To estimate the uncertainty of the experimental data, each measurement was repeated five times.
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Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
Scannera
DOSIMETRY ON CT CLINICAL SCANS
NetPVfit ðxÞ ¼ ax2 þ b;
ð8Þ
where b was NetPVfit ðx ¼ 0Þ provided by Method 3. The corresponding DL(0) was eventually calculated from b, applying Equation (4) after replacing Deq,air with DL(0). This assumption, as explained above for IC and PC, was justified because of the limited effect of the beam spectral variations due to PMMA phantoms(13, 15 – 18). A last expedient was also adopted for low CTDIvol scans such as thorax and abdomen. To measure DL(0) values higher than GF minimum recordable dose, each film piece was irradiated three times. DL(0) per scan was then calculated as one-third of the DL(0) recorded value. Both DL(0) values provided by Method 2 and 3 were then extrapolated to D100(0) as indicated in the CT reference scan section. CTDIvol,w were eventually calculated as indicated in Equation (2). Uncertainty analysis Uncertainties were always estimated according to the standard error propagation theory(25), i.e. considering
Figure 3. Upper: ROI on GF used to read PVunexp(x) and PVexp(x) for each position x. Lower: calculated NetPV(x) profile with ripple (dots) and the smoothed one (line).
different contributions to total CTDIvol,w uncertainties, depending on the three methods focused in the study. For Method 1, these were the standard deviations of the experimental data and the uncertainty of the PC calibration factor equal to 1.3 % (Comecer, Corporation SpA, Castel Bolognese, 48014 Ravenna, Italy, certificate number: 11212/S/06/12). Exactly the same error sources were considered for DL(0) measurements for Methods 2, with IC calibration factor affected by the uncertainty of 1.3 %. Method 2 involved also the additional assessment due to h 100(L). This contribution was calculated as ref standard deviation of the Dref fit ð100Þ=Dfit ðLÞ values around the regression function(26) of h 100(L). The h 100(L)-related error was considered also for Method 3. In this last approach, particular care was taken in D L(0) error evaluation. This evaluation was carried out as indicated by Rampado 2 due to calibraet al. (19), adding the variance GFscal tion fit, to GFsb2 related to the uncertainty in
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Conversely, DL(0) assessments following Methods 2 and 3 were performed directly on clinical CT scans. All scan parameters (see Table 1) were directly selectable at console with the exception of the scan length L for helical scans due to the over-ranging effect. Therefore, over-ranging were beforehand quantified on Tomographs A and B as proposed by van der Molen et al. (23) for scanners without dynamic collimation. The same procedure was not implementable on Tomograph C because it was equipped with dynamic collimation. Therefore, over-ranging on Tomograph C scans was measured by computed radiography plates as indicated by Tien et al.(24). After that, to minimise the role of h100(L), L was always selected for each clinical scan (see Table 1) as closest as possible to L ¼ 100 mm. Both IC and then GF were located exactly in the x ¼ 0 position, i.e. in the centre of the scans (see Figure 1). To estimate the uncertainty of the experimental data, each measurement was repeated five times. In addition, as shown in Figure 2, all GF placed in the peripheral locations were also aligned parallel to the next external phantom surface. As a consequence, in line with Rampado et al.(13), the maximal GF geometric efficiency was achieved in correspondence to the maximal dose rate on GF. Methods 2 and 3 showed also significant differences in DL(0) assessment procedures, whereas Method 2 provided it directly and Method 3 required reading PVunexp and PVexp profiles as shown in Figure 3. In this case, Equation (3) was applied to each position, providing a NetPV(x) profile. NetPV(x) was in addition convoluted with the smoothing function p(x/k)/k described in Introduction section. The smoothed NetPVðxÞ profile was then fitted with the polynomial function:
D. TREVISAN ET AL.
b measurements as follows: 2 GF sDL ð0Þ
2 ¼ GF scal þ GF sb2 :
ð9Þ
2 and GF sb2 were estiIn the present work, GF scal mated by means of the same fit parameters of Equation (4) with:
2 GF scal ¼
@Deq;air @a
2
da2 þ
@Deq;air 2 2 db @b
¼ b4 da2 þ b2 db2
ð10Þ
2 GF sb ¼
@DLð0Þ 2 2 db ¼ ð2ab þ bÞ2 db2 ; @b
ð11Þ CTDIvol and CTDIw assessment
where da and db are the standard deviations of the fit parameters a and b of Equation (4) and db is the standard deviation in b measurements.
RESULTS Central cumulative dose DL(0) The present work is focused on DL(0) measurements summarised in Table 2. The fit parameters of Equation (4) required in the Method 3 are valid for all tomographs as showed in Figure 4. DL(0) experimental data from Methods 2 and 3 are good overlapped each other: their relative discrepancies are ,5 % for 65 % of the data. A maximal difference of 12 %
Tomographs A and C always allow the selection of all scan parameters required by Method 1 (see Table 1), whereas for Tomograph B, the total collimation of 7.2 and 19.2 mm is not available among the axial scans saved in the CT protocols database. The six CTDIvol,w values obtained by Method 1 on Tomographs A and B reproduce the nominal ones very well (see Table 3), with relative disagreement of always ,5 %. On the contrary, there is an unexpected disagreement of 22 % on Tomograph C for the paediatric abdomen protocol. Methods 2 and 3 proposed in the present study allow CTDIvol,w assessments for each clinical CT scan without any exception. CTDIvol,w evaluations are performed multiplying DL(0) with corresponding h 100(L) as indicated in Equation (5). Figures 5 and 6 show all
Table 2. Central cumulative dose DL(0) in centre and in periphery of the PMMA phantoms. Protocol
Scannera
DL(0) (mGy) Centre
Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
A B A B A B A B A B C C
Periphery
IC
GF
IC
GF
88.0 (1.3) 51.0 (1.3) 50.3 (1.3) 56.5 (1.3) 45.6 (1.3) 30.4 (1.3) 9.7 (1.3) 5.2 (1.4) 12.2 (1.3) 8.5 (1.4) 8.7 (1.4) 3.8 (1.4)
87.9 (4.2) 54.8 (7.8) 55.9 (4.5) 57.9 (6.7) 45.1 (5.5 33.4 (5.2) 10.0 (4.6) 5.2 (6.5) 12.9 (5.7) 9.2 (7.2) 8.8 (4.5) 3.9 (6.5)
100.5 (1.3) 53.9 (1.4) 57.2 (1.4) 60.6 (1.9) 59.4 (1.4) 33.0 (1.3) 17.0 (1.7) 9.0 (9.6) 21.9 (4.3) 15.3 (8.7) 15.8 (4.3) 6.2 (4.4)
94.5 (5.5) 53.9 (5.5) 64.6 (4.8) 59.6 (8.1) 59.5 (4.9) 33.1 (4.7) 17.7 (4.1) 8.7 (5.5) 20.9 (5.6) 15.1 (5.7) 15.7 (5.6) 6.4 (7.0)
Brackets, the relative uncertainty (%). A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S. GF: GF XR-QA2; IC: ion chamber 10` 520.6. a
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and
is recorded for the head spiral protocol of Tomograph A. Experimental data indicate also that Method 2 is generally more precise than Method 3. This evidence is emphasised in the centre of the PMMA phantoms, where DL(0) relative uncertainty for Method 1 is always 1.3 %, whereas the one of Method 2 is in the range of 4.2–7.8 %. Nevertheless, Method 2 increases systematically in uncertainty passing to the peripheral of the phantoms, whereas the same behaviour is absent for GF (Method 3). Instead, Method 3 shows an increment in precision in the peripheral locations of the body phantom. In particular, for thorax and abdomen protocols on Tomograph B, relative uncertainty of GF measurements is 5.5 and 5.7 %, whereas that of Method 2 is 9.6 and 8.7 %, respectively.
DOSIMETRY ON CT CLINICAL SCANS
h 100(L) functions. Their values used in D100(0) evaluation are reported in Table 4, with their relative uncertainty. It can be noted that this last parameter is not taken into account for h 100(L) ¼ 1. This assumption is justified because h 100(L) correction is not required for L ¼ 100 mm. CTDIvol,w obtained with Method 2 are also reported in Table 3. All measurements reproduce nominal CTDIvol,w values very well with relative disagreement ranging from 22.5 up to 4.1 %. Method 3
DISCUSSION Comparison of IC vs GF in DL(0) measurements The general good accordance of Methods 2 and 3 in DL(0) measurements is in agreement with the results of Rampado et al.(13) about CTDIw measurement with GF. This evidence is remarkable because of the significant procedural differences of these two methods proposed in this study. Experimental data indicate that Method 2 is generally more precise than Method 3. This evidence is justified because Method 3 always requires a GF calibration introducing the add2 . Nevertheless, itional uncertainty contribute GF scal these GF drawbacks are not so evident for peripheral locations of body phantoms. This experimental evidence has two reasons: at first, DL(0) values are here approximately twice than those in the centre of the phantom. As a consequence, as shown in Figure 4, in accordance with Devic et al. (27), increased DL(0) values involve a decrement of GF scal =DL ð0Þ: Additionally, Method 3 evaluates the NetPVðxÞ profile obtained by averaging NetPV(x) exactly on the periodicity k found on NetPV(x) itself, as explained in Introduction section. In particular, this smoothing procedure is very effective for large nominal collimation and pitch such as for thorax and abdomen CT protocols of Tomograph B (see Figure 7). Figure 7 shows also the length L of the IC in comparison with the periodicity of the abdomen protocol DL(x) profile. In particular,
Table 3. Comparison between nominal and measured CTDIvol,w (with its relative uncertainty). Protocol
Scannera
Measured CTDIbvol,w (mGy)
Nominal CTDIvol,w (mGy) Method 1
Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
A B A B A B A B A B C C
95.2 52.8 56.7 59.4 70.9 32.5 12.9 7.6 16.2 12.3 13.5 5.4
95.8 (1.5 %) 50.2 (1.6 %) 56.2 (1.6 %) n.a. 69.5 (1.5 %) n.a. 12.8 (1.8 %) n.a. 16.0 (1.6 %) n.a. 13.3 (2.0 %) 6.6 (1.6 %)
Disagreement from nominal CTDIvol,w (%)
Method 2
Method 3
1
2
3
96.3 (1.0 %) 51.0 (1.3 %) 54.9 (1.0 %) 59.2 (1.4 %) 68.0 (1.1 %) 32.1 (1.0 %) 13.1 (1.8 %) 7.3 (7.8 %) 16.6 (3.8 %) 12.3 (7.8 %) 13.4 (3.4 %) 5.4 (3.4 %)
92.3 (4.0 %) 52.2 (4.7 %) 61.7 (3.6 %) 59.0 (5.9 %) 67.9 (3.1 %) 33.2 (3.6 %) 13.6 (3.9 %) 7.1 (4.8 %) 16.2 (5.2 %) 12.4 (5.0 %) 13.5 (4.5 %) 5.6 (5.6 %)
20.6 4.9 0.9 n.a. 2.0 n.a. 0.8 n.a. 1.2 n.a. 1.5 222.2
21.2 3.4 3.2 0.3 4.1 1.2 21.6 3.9 22.5 0.0 0.7 0.0
3.0 1.1 28.8 0.7 4.2 22.2 25.4 6.6 0.0 20.8 0.0 23.7
Method 1: axial scan mode and 100-mm PC; Method 2: clinical mode with 20-mm IC; Method 3: clinical mode with GF. a A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S. b CTDIvol for helical and CTDIw for sequential CT scans.
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Figure 4. Calibration of GF in free-air: air equilibrium dose (Deq,air) measured values were fitted as function of NetPV(#). Inset figure provides the uncertainties of the fit parameters da, db and the determined relative calibration uncertainty GFscal (%).
provides a wider disagreement in the interval between 28.8 and 6.6 % (see Table 3).
D. TREVISAN ET AL. Table 4. h 100(L) values for each CT scans with scan length L in different locations of the PMMA phantoms. Protocol
Scannera
h 100(L) (#)
Scan length (mm)
Centre
Figure 5. Dimensionless h 100(L) functions of Tomograph A in the centre (BC) and in the peripheral axes (BP) of the body phantom and in the centre (HC) and in the peripheral axes (HP) of the head phantom. Fit parameters B, C, D and their uncertainties dB, dC, dD are reported.
A
100.0
1.0
1.0
B
115.2
0.95 (1.0) 0.97 (0.9)
A
99.9
1.0
1.0
B
99.9
1.0
1.0
A B A B A B C C
49.2 100.2 149.1 119.3 151.1 121.0 100.0 100.0
1.40 (0.4) 1.0 0.82 (1.7) 0.90 (1.2) 0.81 (1.7) 0.90 (1.3) 1.0 1.0
1.18 (0.5) 1.0 0.92 (1.3) 0.96 (1.0) 0.91 (1.3) 0.96 (1.0) 1.0 1.0
Brackets, the relative standard deviation (only for h100(L) = 1) (%). a A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S.
Figure 7. NetPV(x) profiles, compared with chamber length (L ¼ 20 mm), in the peripheral axes of the phantom for very different scan parameters: inner ear CT protocol for Tomograph A (triangles), CT abdomen protocol (line) for Tomograph B and its smoothed profile (dashed line).
Figure 6. Dimensionless h 100(L) functions of Tomograph B in the centre (BC) and in the peripheral axes (BP) of the body phantom and in the centre (HC) and in the peripheral axes (HP) of the head phantom. Fit parameters B, C, D and their uncertainties dB, dC, dD are reported.
this suggests that the ripple effect in NetPV(x), and therefore in DL(x), has significant consequences on uncertainties of Method 2: by repeating the measurements, the IC (with sensitive length of 20 mm) can be located with equal probability in a peak or valley of DL(x). As a consequence, the relative uncertainty in peripheral location of the body phantom is drastically increased.
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Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
Periphery
DOSIMETRY ON CT CLINICAL SCANS
CTDIvol,w assessment
CONCLUSIONS In computed tomography practice, CTDIvol,w values are usually recorded and stored for each clinical scan. These data must be reliable because they play an important role in each dosimetric evaluation on population. The traditional CTDIw-based approach is generally satisfactory in CT dosimetry. However, it can be used only for those helical scans that can be adapted to axial ones. As a consequence, CT protocols of wide use cannot be easily checked. An additional limitation of the traditional CTDIwbased method can take place whenever some scan parameters, such as the presence of different bow-tie filters, are automatically selected. As outlined in this study, a significant discrepancy between nominal and measured CTDIvol up to 22 % are achieved under these conditions. The present work indicates a way to overcome these troubles because all measurements reported in the present work are always carried out following the same procedure, even on different clinical CT protocols. Moreover, the introduced correction factor h 100(L) is a dimensionless quantity independent of pitch as well as of X-ray beam collimation. Therefore, it can be measured only at the acceptance test or after significant maintenance operations. The method proposed in the present study is also of easier implementation if a short IC is used. In this case, particular care must be taken for helical CT scans with large collimation and pitch increasing the number of measurements in the peripheral location of the dosimetry phantom. FUNDING This work was supported by the Azienda Provinciale per i Servizi Sanitari, Trento, Italy. REFERENCES
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1. International Electrotechnical Commission (IEC). Medical electrical equipment part 2– 44: particular requirements for the safety of X-ray equipment for computed tomography, third edn. IEC 60601-2-44. IEC Geneva Switzerland (2009). 2. The American Association of Physicists in Medicine (AAPM). The measurement, reporting and management
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The traditional CTDIw-based approach (Method 1) always requires adapting each CT clinical scan to the corresponding axial one. This procedure can be used only if the scan parameters summarised in Table 1, with the exception of pitch and scan length, are available even for axial protocols. In particular, this is not possible for Tomograph B, for which some nominal collimations are selectable only for helical scans, and not for axial one. As a consequence, it is not possible to perform CTDIw measurements with collimations of 7.2 and 19.2 mm. It is also remarkable that 19.2 mm is the most frequently selected nominal collimation for helical CT protocols of Tomograph B. For this reason, both DL(0)-based methods proposed in the present study suggest a way to overcome this limitation. In particular, the possibility of performing CTDIvol,w assessment for each CT clinical scan is crucial. Even when all scan parameters of the helical scan seem to be selectable to axial one, the traditional Method 1 can be still error prone. This is the case of Tomograph C, for which two different bow-tie filters are working for adult and paediatric abdomen protocols(28), even if all scan parameters summarised in Table 1 are the same for both kind of patients. In particular, Method 1 provides for the paediatric protocol a CTDIvol of 6.6 mGy, which is a lot higher than the nominal one of 5.4 mGy. This gap of 22 % is not acceptable in the dosimetric praxis(14). Otherwise, both DL(0)-based approaches (Method 2 and 3) eliminate this error source, pointing out the reliability of the nominal CTDIvol of the paediatric abdomen protocol. In addition, performing dosimetry directly on clinical scans as proposed in the present work does not compromise accuracy of CTDIvol,w evaluations, even when the tomograph does not allow selecting L ¼ 100 mm. This is in particular the case of Tomograph A, where the abdomen and inner ear protocols always require L of .151.1 mm and ,49.2 mm, respectively. Although h 100(L) values are significantly different from one (see Table 4), measured CTDIvol values are overlapped very good to those obtained by Method 1, with discrepancies from the nominal data of ,5 % as shown in Table 3. CTDIw obtained for sequential CT protocols confirm also the robustness of the dosimetric approach proposed in the present work, and in line with Descamps et al.(29), they suggest to use h 100(L) for each kind of CT protocols. The DL(0)-based approach in CTDIvol,w assessment seems to be easier in feasibility following Method 2. In fact, GF require a calibration for each film lot. In addition, Method 2 is generally more precise and accurate than Method 3. This aspect is evident for head spiral protocol of Tomograph A with nominal CTDIvol of 56.7 mGy, for which Methods 1 and 2 provide high accurate CTDIvol estimations.
Conversely, CTDIvol assessment by means of Method 3 is less than satisfying because 61.7 mGy is higher than all others CTDIvol evaluations. This disagreement reproduces the maximal one obtained by Rampado et al.(13) In particular, it seems to be related to GF non-uniformity, indicated in previous works as an important error source in GF dosimetry(27, 30). Nevertheless, even in this case, Method 3 allows performing the dosimetric check required in the study by International Electrotechnical Commission(14).
D. TREVISAN ET AL.
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16. International Electrotechnical Commission (IEC). Medical Diagnostic X-ray Equipment-Radiation Conditions for Use in the Determination of Characteristics, second edn. IEC 61267. IEC Geneva Switzerland (2005). 17. Radcal Corporation. Reference class-chamber—technical specifications. Available on http://www.radcal.com/10x50-6ct (accessed 18 October 2012). 18. Butson, M. J., Cheung, T. and Yu, P. K. Measurement of energy dependence for XRCT radiochromic film. Med. Phys. 33, 2923– 2925 (2006). 19. Rampado, O., Garelli, E., Deagostini, S. and Ropolo, R. Dose and energy dependence of response of Gafchromicw XR-QA film for kilovoltage X-ray beams. Phys. Med. Biol. 51, 2871–2881 (2006). 20. Menegotti, L., Delana, A. and Martignano, A. Radiochromic film dosimetry with flatbed scanners: a fast and accurate method for dose calibration and uniformity correction with single film exposure. Med. Phys. 35, 3078–2508 (2008). 21. Devic, S., Seuntjens, J., Sham, E., Podgorsak, E. B., Schmidtlein, C. R., Kirov, A. S. and Soares, C. G. Precise radiochromic film dosimetry using a flat-bed document scanner. Med. Phys. 32, 2245– 2253 (2005). 22. Merimaa, K., Tapiovaara, M., Kosunen, A. and Toroi, P. Calibration and features of air-kerma length product meters. Radiat. Prot. Dosim. 152, 384–392 (2012). 23. Van der Molen, A. J. and Geleijns, J. Overranging in multisection CT: quantification and relative contribution to dose-comparison of four 16-section CT scanners. Radiology 242, 208–216 (2007). 24. Tien, C. J., Winslow, J. F. and Hintenlang, D. E. A methodology for direct quantification of over-ranging length in helical computed tomography with real-time dosimetry. J. Appl. Clin. Med. Phys. 12, 350– 361 (2010). 25. Taylor, J. R. An introduction to error analysis, second edn. University Science Books. ISBN 0935702423 (1997). 26. Glantz, S. A. Primer of biostatistics, fifth edn. McGrawHill. ISBN 0071379460 (2002). 27. Devic, S., Seuntjens, J., Hegyi, G., Podgorsak, E. B., Soares, C. G., Kirov, A. S., Ali, I., Williamson, J. F. and Elizondo, A. Dosimetric properties of improved Gafchromic films for seven different digitizers. Med. Phys. 31, 2392–2401 (2004). 28. Siemens Corporation. Computed Tomograph Somatom Definition AS – User Manual, Siemens Corporation (2012). 29. Descamps, C., Gonzalez, M., Garrigo, E., Germanier, A. and Venencia, D. Measurements of the dose delivered during CT exams using AAPM Task Group No. 111. J. Appl. Med. Phys. 13, 293– 302 (2012). 30. Bouchard, H., Lacroix, F., Beaudoin, G., Carrier, J.-F. and Kawrakow, I. On the characterization and uncertainty analysis of radiochromic film dosimetry. Med. Phys. 36, 1931–1946 (2009).
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of radiation dose in CT. AAPM, Report of AAPM Task Group 23. AAPM One Physics Ellipse, College Park, MD 20740-3846 (2008). Brenner, D. J. Is it time to retire the CTDI for CT quality assurance and dose optimization?. Med. Phys. 32, 3225–3226 (2005). Brenner, D. J., McCollough, C. H. and Orton, C. G. Is it time to retire the computed tomography dose index (CTDI) for CT quality assurance and dose optimization. Med. Phys. 33, 1189–1191 (2006). Boone, J. M. The trouble with CTDI100. Med. Phys. 34, 1364–1371 (2007). Kim, K. P., Lee, J. and Bolch, W. E. CT dosimetry computer codes: their influence on radiation dose estimates and the necessity for their revision under new ICRP radiation protection standards. Radiat. Prot. Dosim. 146, 252– 255 (2011). Geleijns, J., Artells, M. S., De Bruin, P. W., Mather, R., Muramatsu, Y. and McNitt-Gray, M. F. Computed tomography dose assessment for 160 mm wide, 320 detector row, cone beam CT scanner. Phys. Med. Biol. 54, 3141–3159 (2009). De Denaro, M. and Bregant, P. Dosimetric evaluation of a 320 detector row CT scanner unit. Radiol. Oncol. 45, 64–67 (2011). The American Association of Physicists in Medicine (AAPM). Comprehensive methodology for the evaluation of radiation dose in X-ray computed tomography. AAPM, Report of AAPM Task Group 111. AAPM One Physics Ellipse, College Park, MD 20740-3846 (2010). Dixon, R. L. A new look at CT dose measurement: beyond CTDI. Med. Phys. 30, 1272– 1280 (2003). Dixon, R. L. and Ballard, A. C. Experimental validation of a versatile system of CT dosimetry using a conventional ion chamber: beyond CTDI100. Med. Phys. 34, 3399–3413 (2007). Dixon, R. L. and Boone, J. M. Cone beam CT dosimetry: a unified and self-consistent approach including all scan modalities—with or without phantom motion. Med. Phys. 37, 2703– 2718 (2010). Rampado, O., Garelli, E. and Ropolo, R. Computed tomography dose measurement with radiochromic films and a flatbed scanner. Med. Phys. 37, 189 – 196 (2010). International Electrotechnical Commission (IEC). Evaluation and routine testing in medical imaging departments – Part 3–5: acceptance tests—imaging performance of computed tomography X-ray equipment. IEC 61223. IEC Geneva Switzerland (2006). Martin, C. J., Gentle, D. J., Sookpeng, S. and Loveland, J. Application of Gafchromic film in the study of dosimetry methods phantoms. J. Radiol. Prot. 31, 389–409 (2011).
doi:10.1093/rpd/nct236
MEASUREMENTS OF COMPUTED TOMOGRAPHY DOSE INDEX FOR CLINICAL SCANS D. Trevisan1,*, D. Ravanelli2 and A. Valentini1 1 Department of Medical Physics, S. Chiara Hospital, APSS Trento, Italy 2 Medical Physics, University of Rome Tor Vergata, Rome, Italy *Corresponding author: [email protected]
Dose assessment in computed tomography is nowadays based on indicators such as the weighted computed tomography dose index (CTDIw) and the volume-weighted computed tomography dose index (CTDIvol ), both measured only on single-axial scans. The aim of this study is to extend the set of CT protocols suitable in CTDIvol,w evaluation and therefore to perform measurements directly on clinical CT scans. With this purpose, the present work follows the methodology proposed by the American Association of Physicists in Medicine in the Report No. 111 and focuses on the central cumulative dose DL(0), which predicts CTDIvol,w values for scan length equal to 100 mm. All measurements performed with ion chamber and gafchromic films suggest that it is possible to achieve accurate CTDIvol,w values without selecting single-axial scans tailored to dosimetry. Therefore, it is not always necessary to split dosimetric and clinical CT scanner set-up.
INTRODUCTION Dose assessment in computed tomography (CT) requires easily measurable indicators of dose delivered to organs–tissues. Weighted computed tomography dose index (CTDIw) and volume-weighted CT dose index(1, 2) (CTDIvol ) are nowadays valuable tools for this task, even though they seem to be obsolete because these are always measured only on single-axial scan tailored to CTDIw requirements(3 – 8). Moreover, the relation CTDIvol ¼ CTDIw/pitch used in this approach to extend dose assessment to helical scans is not always of easy implementation. This trouble occurs when some scan parameters with strong influence on CTDIw, such as the beam collimations, are not selectable on axial CT scans. Additionally, dosimetry based on axial scans can be also error prone whenever the presence or removal of different bow-tie filters is not displayed at CT console. In both the cases, CTDIw,vol assessments cannot be performed in an accurate manner. To overcome such limitations, the AAPM TG 111 proposes an alternative dose assessment methodology(9) focusing on the cumulative dose profile D(x)(10 – 12) for sequential as well as for helical CT scans. With reference to Figure 1, D(x) shows in the first case always a quasiperiodical behaviour of fundamental period k equal to the couch position displacement between two consecutive X-ray exposures. Conversely, oscillatory dose distribution is evident for helical scans only on the peripheral axes of the polymethylmethacrylate (PMMA) phantoms, where the periodicity k is the product of pitch and the nominal beam width. Dixon defines also the central cumulative dose DL(0)(10) as the value of D(x) at the centre (x ¼ 0) of the scan with length L. In particular, DL(0) of a helical CT scan with pitch p, if measured
in the centre of the PMMA phantom, comes from a periodicity free profile, and it is in relation to CTDIvol as follows: pDL ð0Þ ¼ CTDI100 for L ¼ 100 mm.
ð1Þ
Moreover, computing the averaging of peaks and valleys of D(x), Equation (1) can be generalised to every kind of CT protocols and locations in the PMMA phantoms. This averaging procedure can be calculated as convolution between D(x) and the rectangular function p(x/k)/k before taking its value at the central scan position (x ¼ 0)(9). As a consequence, Equation (1) is still valid for helical as well as for sequential scans (in this case, p is replaced by 1), and therefore: 1 c 2 p CTDIvol;w ¼ D100 ð0Þ þ D100 ð0Þ; 3 3
ð2Þ
where D100(0) is DL(0) for L ¼ 100 mm. Suffixes c and p indicate central and peripherals locations, respectively, in PMMA standard phantoms. The aim of this study is to overcome the limitations of the CTDIw-based approach, estimating CTDIw as well as CTDIvol directly on clinical CT scans by means of D100(0) measurements as indicated in Equation (2). In line with the AAPM TG 111 approach, all measurements were performed with a short ion chamber(9) (IC). The purpose of the present work is also to investigate the feasibility to use Gafchromic films (GF) for the same D100(0) measurements. The use of GF for CT dosimetry was already investigated by Rampado
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Received 8 May 2013; revised 3 September 2013; accepted 4 September 2013
D. TREVISAN ET AL.
et al.(13), obtaining a discrepancy from CTDIw measured with pencil chamber (PC) always of ,9 % for collimation of .10 mm. GF, unlike IC, have the advantage of choosing the region of interest (ROI) exactly of length k equal to the periodicity in D(x). Length k can change significantly among protocols and scanners, yielding different kind of ripples. Therefore, GF reduce, especially for wide collimation and pitch scans, all ripple-correlated uncertainties due to D(x) quasi-periodical behaviour. To validate the D100(0)-based methods described above, the present work compares all CTDIvol,w values to ones achieved by the standard approach(1, 2). In addition, as required in the study by International Electrotechnical Commission(14), it is also tested whether the relative disagreement between nominal and measured CTDIvol,w is always ,20 %. MATERIALS AND METHODS Dosimetry equipment At first, CTDIw measurements were conducted following exactly the standard dosimetric approach(1, 2) (Method 1), i.e. by means of head and body standard PMMA phantoms (diameter 160` 320 mm) and an 10` 5–3CT PC with charge-collection-length of 100 mm coupled to an Radcal 9010 electrometer (Radcal Corporation, CA, USA). DL(0) measurements were then performed on head and body extended (respectively to 200 and 300 mm length) PMMA phantoms as suggested by J.C. Martin(15). DL(0) was measured at first by a 10`5–0.6 IC (Method 2) with charge-collection-length of 20 mm
coupled to the same electrometer used in Method 1. The same procedure was then repeated using GF XR-QA2 (International Specialty Products, Wayne, NJ, USA, lot number A10121202, expiration date 31 October 2014). This last approach (Method 3) required also the preliminary calibration-set-up of the whole GF dosimetry system consisting of GF and a commercial flatbed scanner (Epson Perfection V750 Pro). Chromatic changes in GF were read by the flatbed scanner providing digital imaging of GF; all measurements on the images were performed by the use of ImageJ (Wayne Rusband, National Institute of Health, USA, http://rsb.info.nih.gov//ij version 1.44o) homemade plug-in. All experimental data were analysed by means of the Matlab 7.0 statistical tool. CTDIw obtained by Method 1 and nominal CTDIw,vol were used to benchmark Methods 2 and 3 proposed in this study. The current investigation involved performing measurements on a Phlips Brilliance 64S, (Tomograph A), a Siemens Biograph 64S (Tomograph B) and a Siemens Somatom Definition AS 128S (Tomograph C). The methods proposed in the present work were at first tested on Tomographs A and B. The advantages of Methods 2 and 3 were pointed out in particular on Tomograph C.
Dosimetry system set-up Ion chamber calibration Both PC and IC used in the present work, respectively, in Methods 1 and 2 had been previously calibrated in
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Figure 1. Typical cumulative dose profile DL(x) for axial and helical scans in different locations of the cylindrical PMMA phantoms. Its ripple is characterised by periodicity k.
DOSIMETRY ON CT CLINICAL SCANS
free-air for the RQA 9 beam according to IEC 61267 report (16). The calibration factor obtained in free-air was also used for both CTDIw and DL(0) in phantom measurements. This assumption is justified by published data of the energy dependence of IC(17) and by the limited effect of the beam spectral variations due to PMMA phantoms(13, 18). As a consequence in the present work IC provided a direct measurement of DL(0).
It is notable that Equation (4) fits saturation effects of GF and at the same time NetPV ¼ 0 in the absence of exposition. Scan length correction factor h 100(L) Clinical CT scans generally do not allow setting L ¼ 100 mm. The D100(0) values reported in Equation (2) were therefore estimated by the use of the correction functions h 100(L). These last describe the relative approach of DL(0) to D100(0) as function of L as:
Scanner set-up and GF calibration
PVunexp 1: NetPV ¼ PVexp
ð3Þ
NetPV values were then fitted with the following equation(13): Deq;air ¼ aNetPV2 þ bNetPV:
ð4Þ
D100 ð0Þ ¼ h100 ðLÞDL ð0Þ:
ð5Þ
h 100(L) were experimentally determined in line with AAPM Report no. 111(9). These correction factors were estimated only on Tomographs A and B, but not on Tomograph C. This choice is justified because the two scans on this last tomograph are always performed at L ¼ 100 mm (see Table 1), and therefore, the direct measurements of D100(0) were carried out. Operatively h 100(L) were obtained by means of central cumulative dose measurements repeated in central as well as in peripheral positions of both PMMA phantoms for scan length L ranging from 40 mm up to 160 mm. Experimental values, indicated in the present section as Dref(L) in order to emphasise their dependence of L, were measured by IC for the following scan parameters: 120 kVp, rotation time 1 s, pitch 0.68, nominal collimation 10 mm, 500 mA s/slice, FOV 300 mm on Tomograph A, and 120 kVp, rotation time 1 s, nominal collimation 19.2 mm, 500 mA s/slice, FOV 300 mm on Tomograph B. Dref(L) values were then fitted with the function: L : ð6Þ ¼ A 1 B exp Dref fit C with fit parameters A, B and C. In the present work, h100(L) functions were then obtained normalising ref Dref fit ðLÞ to Dfit ð100Þ as: h
100
1 Dref L fit ð100Þ ð0Þ ¼ ref : ¼ D 1 Bexp C Dfit ðLÞ
ð7Þ
with D ¼ 1 B expð100=CÞ: Dosimetry of CT clinical scans All CT protocols focused in the present study (always at 120 kVp X-ray tube voltage) are of wide use on patients and show very different scan parameters. All these are reported in Table 1, as well as the phantom type used in dosimetry. To obtain the CTDIvol,w by means of Method 1, all scan parameters reported in Table 1, with the exception of the pitch values and scan lengths L, were
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Whereas PC and IC provided directly CTDIw and DL(0) values, GF used in Method 3 required the preliminary characterisation of the whole dosimetry system. With this aim, the flatbed scanner was tested beforehand in line with the procedure indicated in Rampado et al.(19). Repeatability was between 0.1 and 0.9 %; non-uniformity was 1.3 % and ,0.5 % for 75 % of the ROI (8.0` 20 mm). The second step of the dosimetry system characterisation-set-up involved the calibration of the GF. To investigate the relationship air kerma (Kair) vs GF change in colour, 80 unexposed GF (10` 60 mm) were at first placed on the flatbed central area(20). All GF were then scanned as indicated by Devic et al.(21). On each GF image (TIFF RGB mode, 16 bit per colour, red channel only, resolution 72 dpi, no enhancement, filter or colour correction factors), the average pixel value PVunexp was recorded on a 8` 50 mm ROI. The IC was then placed along the central axis of the gantry solidal to the CT couch of Tomograph A and irradiated in free-air with following CT scan parameters: helical mode 120 kVp, nominal collimation 32`1.25 mm, pitch factor 0.971, rotation time 2 s and scan length 160 mm. Sixteen different mA s/slice values ranging from 65 up to 600 were selected. Particular care was also taken in positioning the centre of the CT scan (x ¼ 0) exactly on the centre of the IC. Under these conditions, the Kair value provided by the IC coincided with the air equilibrium dose Deq;air (9, 22). Exactly the same CT scan parameters used in IC irradiation were then selected for the exposure of the 80 beforehand scanned GF. The exposed GF were read only 24 h after its irradiation(19), following exactly the same procedure described above for the PVunexp evaluation, obtaining PVexp values. The change in the colour of each GF was then described by:
D. TREVISAN ET AL. Table 1. Nominal CTDIvol,w and scan parameters selected for central cumulative dose measurements (always at 120 kVp). Protocol
Rotation time (s)
Nominal collimation (mm)
A
1.5
B A B A B A B A B C C
Pitchb
Nominal CTDIdvol,w (mGy)
Field of view (mm)
Tube current (mA)
Scan lengthc (mm)
10.0
230
333
100.0
95.2
(H)
1.0
28.8
230
380
115.2
52.8
(H)
0.75 1.0 0.75 1.0 0.75 0.5 1.0 0.5 0.5 0.5
25.0 19.2 1.0 7.2 40.0 19.2 40.0 19.2 38.4 38.4
230 230 150 150 350 350 350 350 320 220
360 304 100 112 238 280 243 464 320 160
99.9 99.9 49.2 100.2 149.1 119.3 151.1 121.0 100.0 100.0
56.7 59.4 70.9 32.5 12.9 7.6 16.2 12.3 13.5 5.4
(H) (H) (H) (H) (B) (B) (B) (B) (B) (B)
0.675 0.8 0.375 0.8 0.891 1.4 0.97 1.45 0.8 0.8
a
A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S. Sequential CT scans are considered equivalent to helical ones with pitch ¼ 1. c Scan length is affected by +0.5 mm uncertainty. d CTDIvol for helical and CTDIw for sequential CT scans. In brackets is the phantom type (H head, B body). b
Figure 2. Left: description of dosimetric measurement set-up: a lateral view with phantom and dosemeters placed at scan centre x ¼ 0. Right: transversal view with X-ray source, phantom and GF oriented parallel to external phantom surface (inset figure).
selected on single-axial scans. Measured CTDIw values were then divided by the nominal pitch value of the corresponding clinical helical CT scan(1, 2)
obtaining CTDIvol. To estimate the uncertainty of the experimental data, each measurement was repeated five times.
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Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
Scannera
DOSIMETRY ON CT CLINICAL SCANS
NetPVfit ðxÞ ¼ ax2 þ b;
ð8Þ
where b was NetPVfit ðx ¼ 0Þ provided by Method 3. The corresponding DL(0) was eventually calculated from b, applying Equation (4) after replacing Deq,air with DL(0). This assumption, as explained above for IC and PC, was justified because of the limited effect of the beam spectral variations due to PMMA phantoms(13, 15 – 18). A last expedient was also adopted for low CTDIvol scans such as thorax and abdomen. To measure DL(0) values higher than GF minimum recordable dose, each film piece was irradiated three times. DL(0) per scan was then calculated as one-third of the DL(0) recorded value. Both DL(0) values provided by Method 2 and 3 were then extrapolated to D100(0) as indicated in the CT reference scan section. CTDIvol,w were eventually calculated as indicated in Equation (2). Uncertainty analysis Uncertainties were always estimated according to the standard error propagation theory(25), i.e. considering
Figure 3. Upper: ROI on GF used to read PVunexp(x) and PVexp(x) for each position x. Lower: calculated NetPV(x) profile with ripple (dots) and the smoothed one (line).
different contributions to total CTDIvol,w uncertainties, depending on the three methods focused in the study. For Method 1, these were the standard deviations of the experimental data and the uncertainty of the PC calibration factor equal to 1.3 % (Comecer, Corporation SpA, Castel Bolognese, 48014 Ravenna, Italy, certificate number: 11212/S/06/12). Exactly the same error sources were considered for DL(0) measurements for Methods 2, with IC calibration factor affected by the uncertainty of 1.3 %. Method 2 involved also the additional assessment due to h 100(L). This contribution was calculated as ref standard deviation of the Dref fit ð100Þ=Dfit ðLÞ values around the regression function(26) of h 100(L). The h 100(L)-related error was considered also for Method 3. In this last approach, particular care was taken in D L(0) error evaluation. This evaluation was carried out as indicated by Rampado 2 due to calibraet al. (19), adding the variance GFscal tion fit, to GFsb2 related to the uncertainty in
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Conversely, DL(0) assessments following Methods 2 and 3 were performed directly on clinical CT scans. All scan parameters (see Table 1) were directly selectable at console with the exception of the scan length L for helical scans due to the over-ranging effect. Therefore, over-ranging were beforehand quantified on Tomographs A and B as proposed by van der Molen et al. (23) for scanners without dynamic collimation. The same procedure was not implementable on Tomograph C because it was equipped with dynamic collimation. Therefore, over-ranging on Tomograph C scans was measured by computed radiography plates as indicated by Tien et al.(24). After that, to minimise the role of h100(L), L was always selected for each clinical scan (see Table 1) as closest as possible to L ¼ 100 mm. Both IC and then GF were located exactly in the x ¼ 0 position, i.e. in the centre of the scans (see Figure 1). To estimate the uncertainty of the experimental data, each measurement was repeated five times. In addition, as shown in Figure 2, all GF placed in the peripheral locations were also aligned parallel to the next external phantom surface. As a consequence, in line with Rampado et al.(13), the maximal GF geometric efficiency was achieved in correspondence to the maximal dose rate on GF. Methods 2 and 3 showed also significant differences in DL(0) assessment procedures, whereas Method 2 provided it directly and Method 3 required reading PVunexp and PVexp profiles as shown in Figure 3. In this case, Equation (3) was applied to each position, providing a NetPV(x) profile. NetPV(x) was in addition convoluted with the smoothing function p(x/k)/k described in Introduction section. The smoothed NetPVðxÞ profile was then fitted with the polynomial function:
D. TREVISAN ET AL.
b measurements as follows: 2 GF sDL ð0Þ
2 ¼ GF scal þ GF sb2 :
ð9Þ
2 and GF sb2 were estiIn the present work, GF scal mated by means of the same fit parameters of Equation (4) with:
2 GF scal ¼
@Deq;air @a
2
da2 þ
@Deq;air 2 2 db @b
¼ b4 da2 þ b2 db2
ð10Þ
2 GF sb ¼
@DLð0Þ 2 2 db ¼ ð2ab þ bÞ2 db2 ; @b
ð11Þ CTDIvol and CTDIw assessment
where da and db are the standard deviations of the fit parameters a and b of Equation (4) and db is the standard deviation in b measurements.
RESULTS Central cumulative dose DL(0) The present work is focused on DL(0) measurements summarised in Table 2. The fit parameters of Equation (4) required in the Method 3 are valid for all tomographs as showed in Figure 4. DL(0) experimental data from Methods 2 and 3 are good overlapped each other: their relative discrepancies are ,5 % for 65 % of the data. A maximal difference of 12 %
Tomographs A and C always allow the selection of all scan parameters required by Method 1 (see Table 1), whereas for Tomograph B, the total collimation of 7.2 and 19.2 mm is not available among the axial scans saved in the CT protocols database. The six CTDIvol,w values obtained by Method 1 on Tomographs A and B reproduce the nominal ones very well (see Table 3), with relative disagreement of always ,5 %. On the contrary, there is an unexpected disagreement of 22 % on Tomograph C for the paediatric abdomen protocol. Methods 2 and 3 proposed in the present study allow CTDIvol,w assessments for each clinical CT scan without any exception. CTDIvol,w evaluations are performed multiplying DL(0) with corresponding h 100(L) as indicated in Equation (5). Figures 5 and 6 show all
Table 2. Central cumulative dose DL(0) in centre and in periphery of the PMMA phantoms. Protocol
Scannera
DL(0) (mGy) Centre
Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
A B A B A B A B A B C C
Periphery
IC
GF
IC
GF
88.0 (1.3) 51.0 (1.3) 50.3 (1.3) 56.5 (1.3) 45.6 (1.3) 30.4 (1.3) 9.7 (1.3) 5.2 (1.4) 12.2 (1.3) 8.5 (1.4) 8.7 (1.4) 3.8 (1.4)
87.9 (4.2) 54.8 (7.8) 55.9 (4.5) 57.9 (6.7) 45.1 (5.5 33.4 (5.2) 10.0 (4.6) 5.2 (6.5) 12.9 (5.7) 9.2 (7.2) 8.8 (4.5) 3.9 (6.5)
100.5 (1.3) 53.9 (1.4) 57.2 (1.4) 60.6 (1.9) 59.4 (1.4) 33.0 (1.3) 17.0 (1.7) 9.0 (9.6) 21.9 (4.3) 15.3 (8.7) 15.8 (4.3) 6.2 (4.4)
94.5 (5.5) 53.9 (5.5) 64.6 (4.8) 59.6 (8.1) 59.5 (4.9) 33.1 (4.7) 17.7 (4.1) 8.7 (5.5) 20.9 (5.6) 15.1 (5.7) 15.7 (5.6) 6.4 (7.0)
Brackets, the relative uncertainty (%). A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S. GF: GF XR-QA2; IC: ion chamber 10` 520.6. a
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and
is recorded for the head spiral protocol of Tomograph A. Experimental data indicate also that Method 2 is generally more precise than Method 3. This evidence is emphasised in the centre of the PMMA phantoms, where DL(0) relative uncertainty for Method 1 is always 1.3 %, whereas the one of Method 2 is in the range of 4.2–7.8 %. Nevertheless, Method 2 increases systematically in uncertainty passing to the peripheral of the phantoms, whereas the same behaviour is absent for GF (Method 3). Instead, Method 3 shows an increment in precision in the peripheral locations of the body phantom. In particular, for thorax and abdomen protocols on Tomograph B, relative uncertainty of GF measurements is 5.5 and 5.7 %, whereas that of Method 2 is 9.6 and 8.7 %, respectively.
DOSIMETRY ON CT CLINICAL SCANS
h 100(L) functions. Their values used in D100(0) evaluation are reported in Table 4, with their relative uncertainty. It can be noted that this last parameter is not taken into account for h 100(L) ¼ 1. This assumption is justified because h 100(L) correction is not required for L ¼ 100 mm. CTDIvol,w obtained with Method 2 are also reported in Table 3. All measurements reproduce nominal CTDIvol,w values very well with relative disagreement ranging from 22.5 up to 4.1 %. Method 3
DISCUSSION Comparison of IC vs GF in DL(0) measurements The general good accordance of Methods 2 and 3 in DL(0) measurements is in agreement with the results of Rampado et al.(13) about CTDIw measurement with GF. This evidence is remarkable because of the significant procedural differences of these two methods proposed in this study. Experimental data indicate that Method 2 is generally more precise than Method 3. This evidence is justified because Method 3 always requires a GF calibration introducing the add2 . Nevertheless, itional uncertainty contribute GF scal these GF drawbacks are not so evident for peripheral locations of body phantoms. This experimental evidence has two reasons: at first, DL(0) values are here approximately twice than those in the centre of the phantom. As a consequence, as shown in Figure 4, in accordance with Devic et al. (27), increased DL(0) values involve a decrement of GF scal =DL ð0Þ: Additionally, Method 3 evaluates the NetPVðxÞ profile obtained by averaging NetPV(x) exactly on the periodicity k found on NetPV(x) itself, as explained in Introduction section. In particular, this smoothing procedure is very effective for large nominal collimation and pitch such as for thorax and abdomen CT protocols of Tomograph B (see Figure 7). Figure 7 shows also the length L of the IC in comparison with the periodicity of the abdomen protocol DL(x) profile. In particular,
Table 3. Comparison between nominal and measured CTDIvol,w (with its relative uncertainty). Protocol
Scannera
Measured CTDIbvol,w (mGy)
Nominal CTDIvol,w (mGy) Method 1
Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
A B A B A B A B A B C C
95.2 52.8 56.7 59.4 70.9 32.5 12.9 7.6 16.2 12.3 13.5 5.4
95.8 (1.5 %) 50.2 (1.6 %) 56.2 (1.6 %) n.a. 69.5 (1.5 %) n.a. 12.8 (1.8 %) n.a. 16.0 (1.6 %) n.a. 13.3 (2.0 %) 6.6 (1.6 %)
Disagreement from nominal CTDIvol,w (%)
Method 2
Method 3
1
2
3
96.3 (1.0 %) 51.0 (1.3 %) 54.9 (1.0 %) 59.2 (1.4 %) 68.0 (1.1 %) 32.1 (1.0 %) 13.1 (1.8 %) 7.3 (7.8 %) 16.6 (3.8 %) 12.3 (7.8 %) 13.4 (3.4 %) 5.4 (3.4 %)
92.3 (4.0 %) 52.2 (4.7 %) 61.7 (3.6 %) 59.0 (5.9 %) 67.9 (3.1 %) 33.2 (3.6 %) 13.6 (3.9 %) 7.1 (4.8 %) 16.2 (5.2 %) 12.4 (5.0 %) 13.5 (4.5 %) 5.6 (5.6 %)
20.6 4.9 0.9 n.a. 2.0 n.a. 0.8 n.a. 1.2 n.a. 1.5 222.2
21.2 3.4 3.2 0.3 4.1 1.2 21.6 3.9 22.5 0.0 0.7 0.0
3.0 1.1 28.8 0.7 4.2 22.2 25.4 6.6 0.0 20.8 0.0 23.7
Method 1: axial scan mode and 100-mm PC; Method 2: clinical mode with 20-mm IC; Method 3: clinical mode with GF. a A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S. b CTDIvol for helical and CTDIw for sequential CT scans.
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Figure 4. Calibration of GF in free-air: air equilibrium dose (Deq,air) measured values were fitted as function of NetPV(#). Inset figure provides the uncertainties of the fit parameters da, db and the determined relative calibration uncertainty GFscal (%).
provides a wider disagreement in the interval between 28.8 and 6.6 % (see Table 3).
D. TREVISAN ET AL. Table 4. h 100(L) values for each CT scans with scan length L in different locations of the PMMA phantoms. Protocol
Scannera
h 100(L) (#)
Scan length (mm)
Centre
Figure 5. Dimensionless h 100(L) functions of Tomograph A in the centre (BC) and in the peripheral axes (BP) of the body phantom and in the centre (HC) and in the peripheral axes (HP) of the head phantom. Fit parameters B, C, D and their uncertainties dB, dC, dD are reported.
A
100.0
1.0
1.0
B
115.2
0.95 (1.0) 0.97 (0.9)
A
99.9
1.0
1.0
B
99.9
1.0
1.0
A B A B A B C C
49.2 100.2 149.1 119.3 151.1 121.0 100.0 100.0
1.40 (0.4) 1.0 0.82 (1.7) 0.90 (1.2) 0.81 (1.7) 0.90 (1.3) 1.0 1.0
1.18 (0.5) 1.0 0.92 (1.3) 0.96 (1.0) 0.91 (1.3) 0.96 (1.0) 1.0 1.0
Brackets, the relative standard deviation (only for h100(L) = 1) (%). a A: Philips Brilliance 64S; B: Siemens Biograph 64S; C: Siemens Somatom Definition AS 128S.
Figure 7. NetPV(x) profiles, compared with chamber length (L ¼ 20 mm), in the peripheral axes of the phantom for very different scan parameters: inner ear CT protocol for Tomograph A (triangles), CT abdomen protocol (line) for Tomograph B and its smoothed profile (dashed line).
Figure 6. Dimensionless h 100(L) functions of Tomograph B in the centre (BC) and in the peripheral axes (BP) of the body phantom and in the centre (HC) and in the peripheral axes (HP) of the head phantom. Fit parameters B, C, D and their uncertainties dB, dC, dD are reported.
this suggests that the ripple effect in NetPV(x), and therefore in DL(x), has significant consequences on uncertainties of Method 2: by repeating the measurements, the IC (with sensitive length of 20 mm) can be located with equal probability in a peak or valley of DL(x). As a consequence, the relative uncertainty in peripheral location of the body phantom is drastically increased.
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Head sequential Head sequential Head spiral Head spiral Inner ear Inner ear Thorax Thorax Abdomen Abdomen Abdomen Paediatric abdomen
Periphery
DOSIMETRY ON CT CLINICAL SCANS
CTDIvol,w assessment
CONCLUSIONS In computed tomography practice, CTDIvol,w values are usually recorded and stored for each clinical scan. These data must be reliable because they play an important role in each dosimetric evaluation on population. The traditional CTDIw-based approach is generally satisfactory in CT dosimetry. However, it can be used only for those helical scans that can be adapted to axial ones. As a consequence, CT protocols of wide use cannot be easily checked. An additional limitation of the traditional CTDIwbased method can take place whenever some scan parameters, such as the presence of different bow-tie filters, are automatically selected. As outlined in this study, a significant discrepancy between nominal and measured CTDIvol up to 22 % are achieved under these conditions. The present work indicates a way to overcome these troubles because all measurements reported in the present work are always carried out following the same procedure, even on different clinical CT protocols. Moreover, the introduced correction factor h 100(L) is a dimensionless quantity independent of pitch as well as of X-ray beam collimation. Therefore, it can be measured only at the acceptance test or after significant maintenance operations. The method proposed in the present study is also of easier implementation if a short IC is used. In this case, particular care must be taken for helical CT scans with large collimation and pitch increasing the number of measurements in the peripheral location of the dosimetry phantom. FUNDING This work was supported by the Azienda Provinciale per i Servizi Sanitari, Trento, Italy. REFERENCES
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The traditional CTDIw-based approach (Method 1) always requires adapting each CT clinical scan to the corresponding axial one. This procedure can be used only if the scan parameters summarised in Table 1, with the exception of pitch and scan length, are available even for axial protocols. In particular, this is not possible for Tomograph B, for which some nominal collimations are selectable only for helical scans, and not for axial one. As a consequence, it is not possible to perform CTDIw measurements with collimations of 7.2 and 19.2 mm. It is also remarkable that 19.2 mm is the most frequently selected nominal collimation for helical CT protocols of Tomograph B. For this reason, both DL(0)-based methods proposed in the present study suggest a way to overcome this limitation. In particular, the possibility of performing CTDIvol,w assessment for each CT clinical scan is crucial. Even when all scan parameters of the helical scan seem to be selectable to axial one, the traditional Method 1 can be still error prone. This is the case of Tomograph C, for which two different bow-tie filters are working for adult and paediatric abdomen protocols(28), even if all scan parameters summarised in Table 1 are the same for both kind of patients. In particular, Method 1 provides for the paediatric protocol a CTDIvol of 6.6 mGy, which is a lot higher than the nominal one of 5.4 mGy. This gap of 22 % is not acceptable in the dosimetric praxis(14). Otherwise, both DL(0)-based approaches (Method 2 and 3) eliminate this error source, pointing out the reliability of the nominal CTDIvol of the paediatric abdomen protocol. In addition, performing dosimetry directly on clinical scans as proposed in the present work does not compromise accuracy of CTDIvol,w evaluations, even when the tomograph does not allow selecting L ¼ 100 mm. This is in particular the case of Tomograph A, where the abdomen and inner ear protocols always require L of .151.1 mm and ,49.2 mm, respectively. Although h 100(L) values are significantly different from one (see Table 4), measured CTDIvol values are overlapped very good to those obtained by Method 1, with discrepancies from the nominal data of ,5 % as shown in Table 3. CTDIw obtained for sequential CT protocols confirm also the robustness of the dosimetric approach proposed in the present work, and in line with Descamps et al.(29), they suggest to use h 100(L) for each kind of CT protocols. The DL(0)-based approach in CTDIvol,w assessment seems to be easier in feasibility following Method 2. In fact, GF require a calibration for each film lot. In addition, Method 2 is generally more precise and accurate than Method 3. This aspect is evident for head spiral protocol of Tomograph A with nominal CTDIvol of 56.7 mGy, for which Methods 1 and 2 provide high accurate CTDIvol estimations.
Conversely, CTDIvol assessment by means of Method 3 is less than satisfying because 61.7 mGy is higher than all others CTDIvol evaluations. This disagreement reproduces the maximal one obtained by Rampado et al.(13) In particular, it seems to be related to GF non-uniformity, indicated in previous works as an important error source in GF dosimetry(27, 30). Nevertheless, even in this case, Method 3 allows performing the dosimetric check required in the study by International Electrotechnical Commission(14).
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