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First published online May 25, 2006
British Journal of Radiology (2006) 79, 888-892
© 2006 British Institute of Radiology
doi: 10.1259/bjr/66519303
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Full paper

Conversion factor for CT dosimetry to assess patient dose using a 256-slice CT scanner

S Mori, PhD, MPH, RT K Nishizawa, PhD, MPH M Ohno, MSc, MPH and M Endo, PhD, MPH

Department of Medical Physics, National Institute of Radiological Sciences, Chiba 263-8555, Japan


   Abstract
Top
 Abstract
Introduction
 Materials and methods
 Results
 Discussion
 Conclusion
 References
 
Recent rapid progress in CT technology has yielded equipment with large numbers of detector rows and standard computed tomography dose index (CTDI) is therefore no longer an adequate integration range. An integration range of 300 mm is necessary to accurately measure dose under a nominal beam width of 128 mm due to scattered radiation. However, such a long phantom is inconvenient to use routinely in cone-beam CT patient dose checking. To assess patient dose accurately with standard dosimetry methods, we determined a conversion factor (CF) which was calculated from the weighted dose profile integral (DPIw) for the 300 mm integration range with a 300 mm long CTDI phantom using a 300 mm long ionization chamber divided by that for the 100 mm integration range with a standard CTDI phantom (140 mm long) with a 100 mm long chamber. CF values increase with increasing nominal beam width and effective energy in the range from 1.5 to 2.0. CF values can also be adapted for use with other CT systems as their dose profiles are thought to be analogous to those for the 300 mm phantom and are useful in any hospital situation to assess accurate patient doses using standard dosimetry methods.


   Introduction
Top
 Abstract
 Introduction
Materials and methods
 Results
 Discussion
 Conclusion
 References
 
State-of-the-art commercially based CT represents a marked improvement over conventional multislice CT (MSCT), especially in cardiac imaging. Its craniocaudal coverage without gantry movement, however, is typically only 20–40 mm and this has limited the width of coverage for cine imaging in the craniocaudal direction. To overcome this disadvantage, we developed a 256-slice CT scanner. Volumetric cine imaging is realised by scanning continuously at the same position (without table movement); this provides a large amount of diagnostic information and solves some of the limitations of present helical CT methods [1–5].

However, since the maximum nominal beam width is 128 mm, which is three or four times larger than the latest MSCT in common use, the dose is increased in proportion to the scan time. Therefore, it is very important to assess the dose for volumetric cine imaging.

For conventional CT dose measurement, a standard computed tomography dose index (CTDI) phantom (140 mm long) [6] and 100 mm long ionization chamber have generally been used. However, the conventional CT dose measurement is not sufficient in 256-slice CT because its beam width is larger than the 100 mm long ionization chamber. We therefore extended the FDA-recommended CTDI phantom [6] to a length of 300 mm, on the basis of our group's previous report that phantom length and integration range for dosimetry needed to be at least 300 mm to represent more than 90% of the line integral dose with a beam width between 20 mm and 138 mm [7]. Since a 300 mm CTDI phantom is quite inconvenient, determining a conversion factor (CF) is useful for all hospitals to assess patient dose accurately using standard dosimetry methods.

We propose CF values obtained by calculating ratios of the dose profile integrals (DPIs) with the standard CTDI and 300 mm long phantoms.


   Materials and methods
Top
 Abstract
 Introduction
 Materials and methods
Results
 Discussion
 Conclusion
 References
 
Second model of the 256-slice CT-scanner
The second model of our 256-slice CT scanner was based on the design of the first model [8, 9] which used a wide-area cylindrical 2D detector incorporating current CT technology. The second model was mounted on the gantry frame of a 16-slice CT [10] (Aquilion; Toshiba Medical Systems, Otawara, Tochigi). The 256-slice CT has 912 (transverse) x 256 (craniocaudal) elements, each approximately 0.5 mmx0.5 mm at the centre of rotation. The 128 mm total beam width allows the continuous use of several collimation sets (e.g. 256x0.5 mm, 128x1.0 mm, 64x2.0 mm). Large and small filters are shaped to compensate for the variable path length of each patient across the scan field of view (FOV). The small filter is used for an object under 240 mm field of view (FOV), and the large filter is used for over 240 mm-FOV (e.g. chest and abdomen). The second model of the 256-slice CT scanner incorporates several improvements over the first model, including the acceleration of rotation time from 1.0 s to 0.5 s per rotation, and elongation of the detector dynamic range from 16 to 18 bits [11]. The detector element consists of a Gd2O2S ceramic scintillator and single-crystal silicon photodiode, as used in conventional multislice CT. Maximum X-ray exposure time per scan is 60 s and exposure can be repeated in a series of scans according to a pre-determined program.

The x–y coordinate plane is parallel to the transverse direction, and the z-coordinate axis is parallel to the craniocaudal direction.

Phantoms
The length of the FDA-recommended CTDI phantom [6] is at least 140 mm. This conventional phantom contains holes just large enough to accept a pencil-shaped ionization chamber. For cone-beam CT dose measurement, the phantom length should be longer, because the nominal beam width of 128 mm is longer than that of the conventional CTDI dosimetry method. Therefore, a CTDI phantom of 300 mm length designed for cone-beam CT (Kyoto-kagaku, Kyoto, Japan) [7] was used. The phantoms, 140 mm and 300 mm long, were made of PMMA (polymethylmethacrylate) with diameters of 160 mm for head examinations and 320 mm for body examinations. Holes of 10 mm diameter for the pencil-shaped ionization chamber were located parallel to the rotation axis, and the centres of the holes were located at the cylinder centre and also 10 mm below the cylinder surface at 90° intervals.

Dosemeter probe
A pencil-shaped ionization chamber with an active length of 100 mm (CT-10; Applied Engineering Inc., Tokyo, Japan) or 300 mm (CT-30; Applied Engineering Inc., Tokyo, Japan) was connected to a dosemeter (AE-132; Applied Engineering Inc., Tokyo, Japan). The 300 mm long chamber was an extended form of the 100 mm long pencil-shaped ionization chamber. The dosemeter was calibrated at NMIJ (National Metrology Institute of Japan) for the appropriate radiation qualities by comparison with the secondary radiation standards.

Effective energy
Effective energy was calculated from the attenuation curve of X-ray intensity. The 300 mm-length ionization chamber was positioned at the horizontal plane that passed through the rotational axis centre plane. The X-ray tube was fixed under the ionization chamber, and X-ray irradiation was initiated. The X-ray intensity was measured by setting aluminium attenuators of various thicknesses between the X-ray tube and the ionization chamber [12]. The effective energy was derived from an attenuation length that gave half of the X-ray intensity produced without aluminium attenuators (half-value layer (HVL)), calculated from the attenuation curve of the beam intensity [13]. Effective energy was measured at 0 mm, 50 mm, 100 mm and 150 mm along the transverse direction for 80 kV, 100 kV, 120 kV and 135 kV with the nominal beam width of 128 mm.

The effective energy at the periphery was averaged from x = 0 mm to x = 100 mm for the small filter and from x = 0 mm to x = 150 mm for the large filters. Those measurement ranges almost completely covered the head and body phantoms, respectively. Average effective energy (Eave) was estimated using the following equation:


Formula 001

where Ec and Ep were defined as the effective energy at the centre and periphery, respectively.

Dose profile integral (DPI)
The DPIs were measured with the 300 mm long chamber for the 300 mm long phantoms (extended dosimetry) and with the 100 mm long chamber for the 140 mm long phantoms. The DPI was given as the output of the ionization chamber. The weighted DPI (DPIw) for x, y coordinates was given by:


Formula 002

where DPIc and DPIp denote DPIs in the measurement range L at the centre and periphery of the phantom. {omega} denotes the beam width.

Then we calculated the DPIw ratio as the CF value:


Formula 003

Again, we omitted the variable {omega} from the notation of DPI. Using the above equations we calculated the CF for various beam sizes.

The phantom was placed on the patient table and its centre was aligned at the isocentre. An ionization chamber was inserted into either the central or one of the peripheral cavities of the phantom (other cavities were filled with PMMA rods). The exposure (expressed as C kg–1) was obtained with the ionization chamber dosemeter and converted to the values of the absorbed dose for PMMA. The exposure length-integral (expressed as C kg–1 cm) was obtained with the ionization chamber dosemeter and converted to the values of DPI for PMMA with the f-factors (0.898 cGy (C kg–1)–1) [14]. Absorbed doses at the centre and periphery of the phantom were calculated using the effective energy at the isocentre (x = 0 mm) and the periphery, respectively.

All scans for the dose measurement were made in the axial scan (non-helical) mode using the 256-slice CT. Scan conditions were tube voltage of 80 kV, 100 kV, 120 kV, or 135 kV; 100 mAs; and nominal beam width of 8 mm, 32 mm, 64 mm, 96 mm, or 128 mm. The results were averaged for 8 repeated DPI measurements.


   Results
Top
 Abstract
 Introduction
 Materials and methods
 Results
Discussion
 Conclusion
 References
 
The effective energies for 80 kV, 100 kV, 120 kV and 135 kV, which were used in converting the exposure to the values of absorbed dose, are summarized in Table 1Go. The effective energy for the large filter is slightly higher than that for the small filter due to the different filter shapes.


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  		Table 1. Effective energies for CT dose index(CTDI) phantoms

 
Figure 1Go shows DPI300,w and DPI100,w which are normalized against 100 mAs for the body phantom. The relationship between DPI300,w and beam width shows good linearity and the linear correlation coefficient is 0.9996 for the 135 kV tube voltage. Other tube voltages have similar values of linear correlation coefficient. DPI100,w shows reasonably good linearity between the beam widths of 8 mm and 128 mm (e.g. the linear correlation coefficient is 0.9910 for 135 kV). Since the 140 mm long phantom could not cover primary dose and scattered dose in the wider beam width such as 128 mm, second term polynomial fitting was better for DPI100,w (e.g. the fitting correlation is 0.9991 for 135 kV). DPI100,w and DPI300,w were represented as follows:


Formula 004


Figure 1
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  		Figure 1. The relationships between weighted dose profile integral(DPIw) at the centre of the body phantom and the beam width. (a) DPI300,w. (b) DPI100,w.

 
These coefficients are summarized in Table 2Go.


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  		Table 2. Coefficients for the fitting equations for DPI100,w, and DPI300,w

 
The DPIws also increase with increasing tube voltage; in particular, DPI300,w for 135 kV for the body phantom is 4.5 times higher than at 80 kV. DPIs for the head phantom (results not shown) are similar to those for the body phantom.

Figure 2Go shows the relationships between CF value and the nominal beam width for each tube voltage. CF value increases gradually up to the nominal beam width of 128 mm with increasing nominal beam width as well as tube voltage. CF values range from 1.5 to 2.0 for the beam widths and tube voltages.


Figure 2
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  		Figure 2. The relationships between the conversion factor and the beam width for(a) the body and (b) the head phantoms.

 
The relationships between CF value and the effective energy are shown in Figure 3Go. The CF increases with increasing effective energy, and it has the same tendency as in Figure 2Go.


Figure 3
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  		Figure 3. The relationships between the conversion factor and the effective energy for(a) the body and (b) the head phantoms.

 

   Discussion
Top
 Abstract
 Introduction
 Materials and methods
 Results
 Discussion
Conclusion
 References
 
DPIws were measured using the standard CTDI and the 300 mm long phantoms for nominal beam widths of 9–128 mm. We believe that the standard 100 mm ionization chamber and 140 mm long phantom are not reliable, even for the 8 mm beam width [7, 15].

Although the standard phantoms measured DPIws for the nominal beam width less than 100 mm, CF was not 1.0. In particular, CF value ranged from 1.5 to 1.8 for the nominal beam widths of 8 mm and 32 mm, which can be selected by the MSCT. These results can be explained by scattered radiation which depends on the phantom length and which affects DPI strongly [7]. Therefore, CF value is important to both MSCT and 256-slice CT in order to assess doses accurately.

CF should be adaptable to MSCT systems because the scatter tails of their dose profiles are thought to be analogous to those for the 256-slice CT, if the effective energies are similar [16].


   Conclusion
Top
 Abstract
 Introduction
 Materials and methods
 Results
 Discussion
 Conclusion
References
 
Use of the standard 100 mm ionization chamber and 140 mm phantoms instead of the inconvenient 300 mm long alternatives is made possible by use of the CF values determined in our paper. It was found that the standard system was unreliable even at 8 mm beam width. The relationship demonstrated between the CF and effective energy will allow this method to be adapted to other CT systems.

Received for publication November 14, 2005. Revision received March 14, 2006. Accepted for publication March 14, 2006.


   References
Top
 Abstract
 Introduction
 Materials and methods
 Results
 Discussion
 Conclusion
 References
 

  1. Mori S, Endo M, Obata T, Murase K, Fujiwara H, Susumu K, et al. Clinical potentials of the prototype 256-detector row CT-scanner. Acad Radiol 2005;12:148–54.[CrossRef][Medline]
  2. Mori S, Kondo C, Suzuki N, Yamashita H, Hattori A, Kusakabe M, et al. Volumetric cine imaging for cardiovascular circulation using prototype 256–detector row computed tomography scanner (4-dimensional computed tomography) a preliminary study with a porcine model. J Compt Assist Tomogr 2005;29:26–30.[CrossRef][Medline]
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  4. Mori S, Obata T, Nakajima N, Ichihara N, Endo M. Volumetric perfusion CT using prototype 256-detector row computed tomography scanner: preliminary study with healthy porcine model. Am J Neuroradiol 2005;26:2536–41.[Abstract/Free Full Text]
  5. Kondo C, Mori S, Endo M, Kusakabe K, Suzuki N, Hattori A, et al. Real-time volumetric imaging of human heart without ECG gating by 256-detector row computed tomography: initial experience. J Compt Assist Tomogr 2005;29:694–8.[CrossRef][Medline]
  6. US FDA Code of Federal Regulations. Diagnostic x-ray systems and their major components. Govt. Printing Office, August 1984;21CFR §1020.23
  7. Mori S, Endo M, Nishizawa K, Tsunoo T, Aoyama T, Fujiwara H, et al. Enlarged longitudinal dose profiles in cone-beam CT and the need for modified dosimetry. Med Phys 2005;32:1061–9.[CrossRef][Medline]
  8. Endo M, Mori S, Tsunoo T, Kandatsu S, Tanada S, Aradate H, et al. Development and performance evaluation of the first model of 4D CT-scanner. IEEE Trans Nucl Sci 2003;50:1667–71.[CrossRef]
  9. Mori S, Endo M, Tsunoo T, Kandatsu S, Tanada S, Aradatte H, et al. Physical performance evaluation of a 256-slice CT-scanner for four-dimensional imaging. Med Phys 2004;31:1348–56.[CrossRef][Medline]
  10. Saito Y, Aradate H, Igarashi K, Ide H. Large area 2-dimensional detector for real-time 3–dimensional CT (4D CT). Proc SPIE 2001;4320:775–82.[CrossRef]
  11. Endo M, Mori S, Kandatsu S, Tanada S, Sugihara N, Saito Y, et al. Development of real 4D CT with real-time reconstruction and display. IEEE NSS/MIC Conf Record 2006;M11–252.
  12. Edyvean S, Lewis M, Britten A. CT scanner dose survey: measurement protocol Ver. 5. http://www.impactscan.org/1997 [Accessed 5 May 2006]
  13. Hubbell J, Selzer S. Tables of x-ray mass attenuation coefficients and mass energy-absorption coefficients. National Institute of Standard and Technology (NIST) physical reference data 1996
  14. International Commission on Radiation Units and Measurements. Radiation dosimetry: X rays generated at potentials of 5 to 150 kV: ICRU. Publication. 17, Washington D.C.:ICRU Publications, 1970
  15. Nakaonechny K, Fallone B, Rathee S. Novel methods of measuring single scan dose profiles and cumulative dose in CT. Med Phys 2005;32:98–109.[CrossRef][Medline]
  16. Felmlee JP, Gray JE, Leetzow ML, Price JC. Estimated fetal radiation dose from multislice CT studies. AJR Am J Roentgenol 1990;154:185–90.[Abstract/Free Full Text]



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