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Assessment of VIE image quality using helical CT angiography: in vitro phantom study
Computerized Medical Imaging and Graphics 28 (2004) 3–12 www.elsevier.com/locate/compmedimag
Assessment of VIE image quality using helical CT angiography: in vitro phantom study Zhonghua Suna,*, John R. Windera, Barry E. Kellyb, Peter K. Ellisb, Peter T. Kennedyb, David G. Hirstc a
Room 15J 13, School of Applied Medical Sciences and Sports Studies, University of Ulster, Newtownabbey BT37 0QB, Northern Ireland, UK b Department of Radiology, Royal Victoria Hospital, Belfast BT12 6BA, Northern Ireland, UK c School of Biomedical Sciences, University of Ulster, Newtownabbey BT37 0QB, Northern Ireland, UK Received 15 April 2003; revised 28 July 2003; accepted 8 September 2003
Abstract The aim of this study is to quantify the effects of helical CT acquisitions parameters on the magnitude of three-dimensional stair-step artefacts, visualization of renal ostium and morphologies of suprarenal stents observed using virtual intravascular endoscopy. This was performed in a phantom of the human abdominal aorta with a stent graft in situ. Stair-step artefacts were quantified by measuring the standard deviation of signal intensity on surface shaded images and the influence of these artefacts on the visualization of arterial ostia and stent morphologies were assessed by three radiologists. The methodology may be used to optimise the CT system performance for helical CT angiography in aortic stent grafting. q 2003 Elsevier Ltd. All rights reserved. Keywords: Image quality; Artefacts; Stent; Computed tomography; Virtual endoscopy
1. Introduction Helical CT angiography (CTA) has been used widely in the evaluation of aortic stent grafting [1 – 3]. Helical CTA generated virtual intravascular endoscopy (VIE) is a rapidly evolving technology that permits interactive two-dimensional (2D) and three-dimensional (3D) visualization techniques to be used for the assessment of aortic stent grafts [4 –6]. This process involves the conversion of sets of 2D images into 3D representations that simulate views that may be obtained by conventional endoscopy. The final image quality is determined by several factors including slice thickness (beam collimation), pitch and image reconstruction interval. However, the acquisition step of CTA sets fundamental limits on the entire process [7]. It has been observed that stair-step artefacts caused by thick image slice acquisition can distort the depiction of polyp morphology in the colon [8]. Several studies have shown * Corresponding author. Tel.: þ44-289036-8121; fax: þ44-2890366028. E-mail address: [email protected] (Z. Sun). 0895-6111/$ - see front matter q 2003 Elsevier Ltd. All rights reserved. doi:10.1016/j.compmedimag.2003.09.001
that the image quality parameters of 3D helical CT obtained with a proper interpolation algorithm method either are better or are quite similar to the parameters of conventional 2D CT [9 –11]. However, stair-step artefacts associated with the interpolation may be generated [9]. A better understanding of this artefact may be important as future clinical applications of CTA extend beyond routine 2D image analysis with advances in 3D volumetric acquisition. Our purpose was to quantitatively assess the effect of specific single helical CT acquisition parameters on the stair-step artefact. We also aimed to qualitatively assess the artefact related to image quality in VIE of arterial ostia and aortic stents. We recognise that multislice CT is becoming more available and will provide increased image quality. However, we think that many imaging centres may only be able to perform single slice helical CT for some time. The measurement methodology for the artefact and results of this study may be used to optimise CT system performance for single slice helical CTA in aortic stent grafting. And the method of artefact quantification will equally apply to multislice systems.
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2. Materials and methods 2.1. Aorta phantom design A human aorta phantom was built with medical rapid prototyping [12]. A commercial aortic stent graft (Zenith, Cook, Europe, Denmark) was deployed inside the phantom to simulate a repaired aortic aneurysm (Fig. 1). The phantom was housed in a small container and filled with contrast medium. The CT attenuation of the phantom was measured 170 HU, therefore, the concentration of contrast medium was diluted to 450 HU which created a X-ray attenuation difference of 280 HU. This was similar to the difference in CT attenuation encountered between aorta vessel wall and enhanced blood in abdominal CT angiographic scanning. 2.2. CT unit and scanning parameters Helical CT scanning was performed on a Philips Medical Systems CT scanner (Tomoscan AV-E1). The rotation time of the CT scanner was 1 s and the tube voltage and current were 120 kVp and 125 mAs for the scanning protocol using 1.0 mm collimation and 120 kVp and 250 mAs for 3.0 and 5.0 mm collimation. The highest selectable X-ray tube current for this CT scanner was 125 mAs at 1.0 mm and 250 mAs for 3.0 and 5.0 mm. A field of view of 256 mm, matrix of 512 £ 512 and 1808 linear interpolation algorithm were used to reconstruct the images. This was similar to our routine abdominal CTA protocol. The phantom was placed parallel to the longitudinal axis of the table of CT scanner
[13]. Image data were acquired at collimations of 1.0, 3.0 and 5.0 mm, pitch 1.0, 1.5 and 2.0 and reconstruction interval of 0.5, 1.0, 2.0, 3.0 and 5.0 mm. A scout view was obtained to define the range of the phantom to be scanned. Scans were taken from the top of the model to the upper aspect of the aortic aneurysm. The total length of CT scanning was 70 mm. After scanning, the data were transferred to a SUN workstation for data processing using Analyze V 1.0 (Analyzedirect.com, Mayo Clinic, USA). All images were post-processed with and without data interpolation. Linear interpolation [14] was used to generate missing data between known surrounding points. A total of 54 data sets were generated. 2.3. Image analysis and assessment VIE images of the superior mesenteric artery (SMA), renal arteries and aortic aneurysm were generated from each volume data set. Two threshold ranges were chosen to create VIE images of the internal surface of the model and the stent wires. For images of the aortic ostia, the upper threshold was 320 HU whilst for the stent wire the lower threshold was 1200 HU. As many variables are involved in the quantification of stair-step artefacts based on the pixel value of the surfaced rendered VIE frames, consistent settings are required to keep the VIE images as same as possible in demonstrating each anatomical structure in all scanning protocols. Lighting and reflectivity are two important factors, which affect the results of measurements. The positioning of the virtual light source, which helps in the visualization of the 3D image and the level of the light that the surface reflects back to the viewing position were kept constant as any change of them produces different results (Fig. 2). Horizontal and vertical lights in our study were set 08 in all scanning parameters. The camera, ‘virtual eye position’ was kept at a constant angle of view (708) that was required to produce VIE images of the arterial ostia and the 3D relationship of stent struts to the arterial ostia in the same positions in all scanning protocols. 2.4. Quantitative image analysis
Fig. 1. (A) Stereolithography model of the human aorta. (B) Close-up view showing the stent deployed inside the model. (C) 3D CT image of the aorta phantom showing the aortic branches and aneurysm. It is noticed that the outer surface of the model was irregular due to manual editing when postprocessing the raw data. Arrow indicates the SMA, while short arrows indicate the renal arteries and arrowhead the fixing part inside the box. (D) Frontal view from the top of the model demonstrating suprarenal stent struts (red colour).
In an artefact free VIE image, the internal aortic wall should appear smooth. We required an objective numerical method for measuring the degree of stair-step artefact present. Line profile analysis produces a graphic display of pixel values from the VIE images. A plot of pixel intensity against pixel number enabled us to assess the degree of artefact present. Standard deviation (SD) of the line profile. To measure the degree of stair-step artefact, we used the SD of the pixel values, a higher SD indicating more variation in pixel intensity and therefore greater presence of artefact. The line profile was drawn across the images with the image intensity being recorded. The line was approximately
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Fig. 4. Measurements of the wire thickness on VIE image acquired at a collimation of 3.0 mm, pitch 1.0 (table speed of 3.0 mm per rotation) and reconstruction interval of 1.0 mm after interpolation. Five different locations were chosen to measure the transverse diameters of stent wires with thickness being 1.04, 1.23, 1.23, 1.29 and 1.29 mm, respectively, measured from the top to the bottom.
Fig. 2. Effect of the angle of virtual lightning on the measurements of SD through a line profile was demonstrated in non-interpolated VIE images of the right renal ostium acquired at a collimation of 5.0 mm, pitch 1.0 (table speed of 5.0 mm per rotation) and reconstruction interval of 3.0 mm. The SD was measured 35.5, 51.4 and 22.9, respectively, when the horizontal and vertical light was set at 08 (A), horizontal light 308 and vertical light 08 (B), horizontal light 08 and vertical light 308 (C).
107 –114 pixels long, with the first 100 pixels being used for analysis (Fig. 3). The measurements were performed at three different positions with two being near to the aortic ostia (SMA and renal ostia), the remaining one in the aortic lumen. Width of the artefacts. The purpose of this measurement was to determine whether there is any measurable difference in the width of the stair-step artefact generated by each scanning protocol. A line profile was recorded in the VIE images including three high and three low intensity areas. This included three stair-step artefacts (Fig. 3). Similar regions were chosen as to those in the measurement of SD. 2.5. Measurement of aortic stent strut thickness It was noted that the stent wire (strut) diameter in the CT image was not representative of the real wire thickness [12]; we wanted to determine the relationship of apparent strut diameter to the CT scanning protocol. Small objects of high contrast in CT imaging may become distorted due to
blurring by the system point spread function. Measurements of the wire diameter were made to determine its appearance versus actual thickness on VIE. The measurements were performed at two locations: renal artery and SMA levels. At each level, five measurements of the transverse diameter of the stent wire were made with an average thickness being calculated (Fig. 4). 2.6. Subjective image analysis The purpose of subjective image analysis was to determine the optimum CT scanning parameters from a clinical perspective. The 54 data sets were 3D rendered to create VIE images of the SMA, renal arteries and aortic aneurysm together with the 3D display of the relationship of aortic stents to the arterial ostia. All examinations were evaluated by three independent experienced reviewers (consultant radiologists). The reviewers were blinded to the scanning protocols and recorded their assessment of the following four aspects based on a 4-point scale scoring method † Clarity of the renal ostium. Score 1 indicated that the renal ostium was not visualized, 2 visualized but distorted, 3 visualised and normal and 4 visualized and perfect. † Definition of suprarenal component. Score 1 indicated the discontinued stent struts, 2 irregular, 3 smooth and 4 smooth with hooks visible. † Presence and degree of stair-step artefacts. Score 1 indicated marked artefacts, 2 moderate, 3 minimal and 4 no artefacts. 2.7. Statistical analysis
Fig. 3. Method to measure the SD through a line profile. A line is drawn on a VIE image (A) with the distance of around 100 pixels. The corresponding line profile shows the degree of the artefacts (B).
The results were analysed using SPSS 11.0 for Windows (SPSS, Inc) to determine the degree of agreement among reviewers. The agreement among the reviewers for the assessment of image quality of VIE was evaluated using the Kendall’s W test [15]. Kendall’s W ranges between 0 (no agreement) and 1 (complete agreement). A p value , 0.05 was considered to be a statistically significant difference.
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3. Results 3.1. Stair-step artefacts quantification The SD of line profile measurements from each scanning protocol was as follows. The mean SD for non-interpolated and interpolated data were 18.6– 32.3 and 5.89 – 21.12, respectively (Table 1). It is clear from these measurements that the magnitude of the stair-step artefacts was much smaller for interpolated data compared to non-interpolated data. However, the SD measured in scanning protocols of 1.0 mm collimation was slightly higher than that measured in 3.0 and 5.0 mm collimations. Table 2 shows that the average scores from the three reviewers for each scanning protocol ranged from 1.0 to 2.7 in non-interpolated and 1.3 to 3.3 in interpolated image data. We considered the score of 2 or 3 (moderate or minimal artefacts) to be acceptable for clinical diagnosis. This meant that all scans of 1.0 mm collimation and 3.0, 5.0 mm collimation with reconstruction interval of 1.0 mm were acceptable for diagnosis in non-interpolated data. While in interpolated data, the range of acceptable scanning protocols was expanded to include scans of 3.0 and 5.0 mm collimation with reconstruction interval of 2.0 mm. Figs. 5– 7 show the relationship of stair-step artefacts to collimation, pitch and reconstruction interval in
non-interpolated data, while Figs. 8 –10 show the same relationship in interpolated data. When measuring the width of the artefacts in all of the scanning protocols in both non-interpolated and interpolated data sets, it was noted that there was no significant difference in the width of artefacts in scans obtained with collimation of 1.0 and 3.0 mm, whether the data were interpolated or not ðp . 0:05Þ: However, the width of the artefacts was significantly smaller in interpolated data than that in noninterpolated data when the collimation was 5.0 mm with reconstruction intervals of 3.0 and 5.0 mm ðp , 0:05Þ: 3.2. Visualization of the renal ostium Table 2 shows that the mean scores from the subjective visual assessment of the renal ostium by three reviewers ranged from 1.0 to 4.0 and 1.7 to 4.0 in non-interpolated and interpolated data, respectively. A renal ostium was required to be smooth or of perfect configuration to be acceptable for clinical diagnosis. Therefore, it was noted that only a small range of scanning protocols were acceptable, including all scans of 1.0 mm collimation and 3.0 mm collimation with reconstruction interval of 1.0 mm in non-interpolated data. In contrast, following interpolation, the acceptable scanning ranges were all scans of 3.0 and 1.0 mm collimation. It was also noted that the visualization of the renal ostium was
Table 1 The mean SD measured in all scanning protocols in both non-interpolated and interpolated image data Scanning parameters
Mean standard deviation (SD)
Collimation (mm)
Pitch
Reconstruction interval (mm)
Non-interpolated data
Interpolated data
1.0 1.0 1.0 1.0 1.0 1.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0
1 1 1.5 1.5 2 2 1 1 1 1.5 1.5 1.5 2 2 2 1 1 1 1 1.5 1.5 1.5 1.5 2 2 2 2
0.5 1.0 0.5 1.0 0.5 1.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0
19.8 26 22.1 28 19.9 28 18.6 27.9 27.5 19.6 29.7 27.4 19.1 28 26.7 24.7 32.3 31 24 21.2 31.8 30.9 26.1 24.6 29.8 30.1 25.1
8.34 8.99 11.26 10.92 11.68 11.04 6.54 5.89 7.69 7.87 5.98 8.75 8.26 6.94 8.56 8.05 9.84 13.1 17.63 9.56 9.51 10.96 17.28 15.34 17.82 17.22 21.12
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Table 2 Mean ranks of subjective image quality for VIE images in the assessment of stair-step artefacts, renal ostium and aortic stent struts in all scanning protocols in both non-interpolated and interpolated image data Scanning parameters Collimation (mm)
1.0 1.0 1.0 1.0 1.0 1.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0
Pitch
1 1 1.5 1.5 2 2 1 1 1 1.5 1.5 1.5 2 2 2 1 1 1 1 1.5 1.5 1.5 1.5 2 2 2 2
Mean ranks of image quality Reconstruction interval (mm)
0.5 1.0 0.5 1.0 0.5 1.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0
Non-interpolated data
Interpolated data
Stair-step artefacts
Renal ostium
Aortic stents
Stair-step artefacts
Renal ostium
Aortic stents
2.3 2 2.7 2.7 2.3 2 2.3 1 1.3 2 1 1 2 1 1 2 1 1 1 2 1 1 1 1.7 1 1 1
4 3.7 3 3.3 4 3 3.3 2 2 3 2 2 3 2 2 3 2 2 2 2 1.7 3 2 1.7 1.3 1.3 1
3.3 3 2.7 3.3 3.3 2.3 3 2 1.7 2.7 2 1.7 2.7 2 1.7 2.7 1.7 1 1 1.7 1.3 1 1 2 1 1 1
3.3 3 1.7 2.7 2 2.7 2.3 2 1.7 3 2 2 2 2 1.7 2 2 1.7 1.3 2.3 2.3 2 1.3 3 1.7 1.7 1.7
4 3.7 3.3 3.7 3.3 3.7 3 3.3 3 3 3 3.3 3 3 3.3 1.7 1.7 1.7 1.7 1.7 2 1.7 1.7 1.7 1.7 1.7 1.7
3.7 3.7 3.7 3.3 3.7 3.7 3 3 1.7 3 2.3 2 1.3 3 1.7 1.3 1.3 1 1 1.3 1.3 1 1.7 1.7 1.7 1.7 1.7
Fig. 5. Non-interpolated VIE images of the right renal ostium obtained with a collimation of 1.0, 3.0 and 5.0 mm (A –C) with pitch of 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) and reconstruction interval of 1.0 mm. The artefacts in image C slightly distorted the renal ostium (arrows).
Fig. 6. Non-interpolated VIE images of the right renal ostium obtained with scanning protocols of 3.0 mm collimation, pitch 1.0, 1.5 and 2.0 (table speed of 3.0, 4.5 and 6.0 mm per rotation, respectively) with reconstruction interval of 1.0 mm (A –C). No apparent change of the artefacts was noticed with the increase of pitch values.
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Fig. 7. Non-interpolated VIE images obtained with a collimation of 5.0 mm, pitch 1.0 (table speed of 5.0 mm per rotation) with reconstruction interval of 1.0, 2.0, 3.0 and 5.0 mm (A –D). Artefacts slightly distorted the renal ostium in A and seriously compromised the renal ostia in B –D due to wide reconstruction intervals.
Fig. 9. Interpolated VIE images of the right renal ostium obtained with scanning protocols of collimation 3.0 mm, pitch 1.0, 1.5 and 2.0 (table speed of 3.0, 4.5 and 6.0 mm per rotation, respectively) and reconstruction interval of 1.0 mm. No apparent difference in the degree of stair-step artefacts was noticed.
determined by the collimation and reconstruction interval, however, it was independent of pitch values. Figs. 5, 7, 8 and 10 show examples of the effect of collimation and reconstruction interval on the visualization of the renal ostia in both non-interpolated and interpolated data. Figs. 6 and 9 show that the visualization of the renal ostium on VIE images was independent of pitch.
scanning protocols ranged from 1.0 to 3.3 and 1.0 to 3.7 in non-interpolated and interpolated data, respectively. A score of 3 indicated a smooth appearance of aortic stent struts visualized on VIE and was considered acceptable for determining the stent position. Most of the scans with 1.0 and 3.0 mm collimation, pitch 1.0 and reconstruction interval of 1.0 mm were acceptable in non-interpolated data. After interpolation, all scans with 1.0 mm collimation and most of the scans with 3.0 mm collimation were acceptable. All scans with 5.0 mm collimation were below the acceptable criteria as their scores ranged from 1.0 to 2.7. Figs. 11 and 12 show the effect of scanning protocols on the visualization of aortic stent struts in both non-interpolated
3.3. Visualization of the stent strut Table 2 shows that the average scores from the subjective assessment of the visualization of the aortic stent struts in all
Fig. 8. Interpolated VIE images of the right renal ostium obtained with scanning protocols of collimation 1.0, 3.0 and 5.0 mm, pitch 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) with reconstruction interval of 1.0 mm. The renal ostium (C) is slightly distorted when the collimation reached 5.0 mm.
Fig. 10. Interpolated VIE images obtained with scanning protocol of 5.0 mm collimation, pitch 1.0 (table speed of 5.0 mm per rotation) and reconstruction of 1.0, 2.0, 3.0 and 5.0 mm (A –D). Note that artefacts existing in these images make renal ostia distorted in all of these images.
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were interpolated or not. When two wires were very close together, they merged, producing an apparent thickness of 3.14 – 3.87 mm (Fig. 11C). Subjective assessment showed that there was significant evidence of concordance among the reviewers in the visualization of the renal ostium, stent struts (Kendall’s W ranged from 0.849 to 0.94, p , 0:05). Significant evidence of concordance was found among the reviewers in the assessment of stair-step artefacts in non-interpolated data (Kendall’s W 0.869, p , 0:05), while no significant evidence of concordance in interpolated data (Kendall’s W 0.401, p . 0:05). This could be because the degree of artefact was not as apparent in interpolated data as that observed in non-interpolated data, and reviewers scored most of the interpolated images as having moderate or minimal artefacts. Fig. 11. VIE images of aortic stents observed in the scanning protocols of 1.0, 3.0 and 5.0 mm collimation, pitch 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) with reconstruction interval of 1.0 mm (A –C) without interpolation. Details of aortic stent struts such as strut junctions (arrow in A) were only visualized in the scanning of 1.0 mm collimation.
and interpolated data. It was noted that details of the stent struts such as strut junctions can only be visualized in images of 1.0 mm collimation. 3.4. Thickness of the aortic stent wire The actual diameter of the stent wire was 0.4 mm. The thickness of the wire ranged from 1.64 to 3.14 mm on axial CT and from 0.82 to 1.29 mm on VIE. The measured thickness of the stent struts did not depend on the collimation used (Fig. 11), regardless of whether the data
Fig. 12. VIE images of aortic stents observed in the scanning protocols of 1.0, 3.0 and 5.0 mm collimation, pitch 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) and reconstruction interval of 1.0 mm (A –C) after interpolation. It is noted that the stents become slightly irregular when the collimation reached 5.0 mm (arrows).
4. Discussion Helical CTA is accepted as the imaging modality of choice for follow up of endovascular stent grafting. Most research has focused on the study of optimal scanning parameters for CT colonography in the detection of colonic polyps [16,17], or CT bronchoscopy in the detection of endobronchial lesions [18,19]. Evaluation of VIE in aortic stent grafting is limited [4 –6], and assessment of image quality of VIE in aortic stent grafting and optimal scanning parameters of CTA together with VIE images has not been reported. Image quality of the original CTA images affects the evaluation of aortic stent grafts as well as the relationship of aortic stents to the arterial branches. From a clinical point of view, 3.0 and 5.0 mm collimation are the most commonly used parameters in helical CTA of aortic stent grafting, therefore, we included these slice thicknesses in our phantom study. The reason we included 1.0 mm collimation in our study was that we considered it may improve spatial resolution, resulting in better visualization of small anatomical structures such as the renal ostium and details of aortic struts. One millimetre collimation is not routinely used in abdominal scanning because of increased image noise, decreased image quality, lack of anatomical coverage and X-ray dose [20]. Increased image noise was found in our study as it was shown that the SD of the stair-step artefacts measured in 1.0 mm scanning protocols was relatively higher than that measured in 3.0 and 5.0 mm collimation scans. However, image quality of VIE images in scanning protocols of 1.0 mm was scored better than that obtained in scans with wider collimation. Stair-step artefacts are a well-recognised problem in clinical CT imaging. However, explicit information on the degree of artefact in relation to acquisition parameters is scarce [21]. Our results showed that for helical CT generated VIE the magnitude of stair-step artefacts was generally related to the collimation and reconstruction interval used but independent of pitch values. This was
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found for both subjective and objective assessment of the VIE images. Therefore, a thinner collimation with a higher pitch, e.g. 2.0, may be suggested to produce good quality VIE images. For evaluation of aortic stent grafts, thinner collimation produced smoother aortic ostia and more detailed images of the stents wires. According to our study, scanning protocols with 1.0 and 3.0 mm collimation after interpolation produced the best VIE images for visualizing the ostium and stent wires. Our results concurred with those obtained with multislice CT regarding the effect of collimation and pitch values on the degree of artefacts [22]. They reported that a large pitch and narrower collimation is much more favourable for suppressing the artefacts. The quality of VIE images in aortic stent grafting was determined by the CT scanning parameters. An acceptable VIE image was required to demonstrate clearly the arterial ostia, stent wires and wire to ostia relationship. Previous studies have shown that in order to eliminate stair-step artefacts, both the collimation and the pitch should be less than the longitudinal dimension of the important feature on inclined surfaces, and the reconstruction interval should be less than the table speed [9,19]. Our study showed that a hybrid CT scanning protocol (1.0 mm collimation around the renal artery and 3.0 mm collimation with pitch 2.0 for the rest of the abdominal aorta) resulted in acceptable VIE image quality and better delineation of the arterial ostia, stent wires as well as wire to ostia relationship. One of the disadvantages of helical CTA in the imaging of stent grafts is the blurring effect that results in overestimation of the thickness of stent wires as displayed in both axial CT and VIE images. It was noted that nearly 50% of the renal ostium appeared covered by the stent wires on VIE images (Figs. 11 and 12). We already know only a small proportion of the renal ostium was covered in reality. It is very important to appreciate this when evaluating the interference of stent struts to the arterial ostia in clinical studies as the interference may cause effects on renal blood flow. Our results showed that the stent morphology depends on the collimation and is independent of pitch values. This did not concur in part with a recent study investigating the effect of artefacts on various stent grafts by multislice CT [23]. The collimation mainly influenced the sharpness of the image and delineation of the stent struts, which is consistent with our results. Their results also showed that higher pitch values caused more pronounced artefacts and influenced the size of the artefacts. The reason for this difference is that their study investigated the effects of scanning parameters on the axial and multiplanar reformatted images, while ours concentrated on virtual endoscopic images. There are several limitations in our study. Firstly, the phantom wall is much denser than the patient’s aortic wall with a CT number of around 170 HU, compared with about
40 HU for aortic tissue. Although, the contrast difference between the aortic wall and the inside medium is similar to the contrast enhanced blood, which was about 200 HU, images obtained from the phantom are slightly different from those of patient data. Secondly, we used static contrast material for imaging. In vivo, blood is flowing and contrast density varies along the vessel and these factors affect image quality. Thirdly, other factors such as the angulation of the aneurysm neck and the direction of the renal arteries in vivo could lead to individual variation of artery and aortic stent visualization even if the angle of view is kept constant. Therefore, in the processing of patient data, the virtual eye position should be customised to ensure the optimal visualization of specific anatomic structures. A potential limitation of phantom studies is that quantitatively obtained results of image quality cannot be translated directly into the in vivo applications. This is because of the difference between the environment in vitro and in vivo. However, in this study, we built the vessel phantom based on real patient data with an infrarenal (abdominal aortic aneurysm) AAA. The size of the aorta and its branches were the same size as in a real patient. The stent graft placed in the phantom was the same product used to treat patients with AAA in our clinical cohort. Therefore, results obtained from the phantom were applicable to clinical studies, as the VIE images of the arterial ostia and the 3D relationship of stent struts to the arterial ostia reflect a similar environment as in real patients.
5. Summary We have tested a series of scanning protocols of CTA in a human vessel phantom with the purpose of finding the optimal one in aortic stent grafting. The phantom was built from a typical aortic aneurysm and a commercial stent graft was deployed inside to simulate the environment to that in real patient’s treatment. Our study demonstrated that image quality of VIE images would benefit from linear interpolation. This dramatically decreased the effect of stair-step artefacts in 3D post-processing images. The beam collimation determines image quality of CTA in aortic stent grafting more than any other factors and image quality was independent of pitch values, according to both subjective and objective assessment. A thin collimation and a high pitch produced better image quality than a wide collimation and a small pitch in the assessment of stair-step artefacts, renal ostium, aortic stent struts as well as the 3D relationship of stent wire to renal ostium. For the purpose of visualizing the encroachment of stent struts to the arterial ostia, a hybrid protocol of slice thickness 1.0 mm around the renal artery and slice thickness of 3.0 mm with pitch 2.0 and reconstruction interval of 1.0 mm for the rest of the abdominal aorta is recommended to provide the greatest coverage with acceptable image quality for visualization of the aortic ostia and suprarenal stent struts on VIE.
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Acknowledgements We thank Mr Gil Stevenson for his assistance in the statistical analysis of the results.
References [1] Moritz JD, Rotermund S, Keating DD, Oestmann JW. Infrarenal abdominal aortic aneurysms: implication of CT evaluation of size and configuration for placement of endovascular aortic grafts. Radiology 1996;198:463 –6. [2] Broeders IAMJ, Blankensteijn JD, Olree M, Mali W, Eikelboom BC. Preoperative sizing of grafts for transfemoral endovascular aneurysm management: a prospective comparative study of spiral CT angiography and conventional CT imaging. J Endovasc Surg 1997;4: 252–61. [3] Rozenblitz A, Marin ML, Veith FJ, Cynamon J, Wahl SI, Bakal CW. Endovascular repair of abdominal aortic aneurysm: value of postoperative follow-up with helical CT. AJR 1995;165(1473):1479. [4] Beier J, Diebold T, Vehse H. Virtual endoscopy in the assessment of implanted aortic stents. CARS 1997;183–8. [5] Rodenwaldt J, Kopka L, Lofti S. CT-supported intra-arterial endoscopy after stent implantation: phantom studies and the initial clinical experience. ROFO 1997;166:180–4. [6] Bartolozzi C, Neri E, Caramella D. CT in vascular pathologies. Eur Radiol 1998;8:679–84. [7] Whiting BR, McFarland EG, Brink JA. Influence of image acquisition parameters on CT artifacts and polyp depiction in spiral CT colonography: in vitro evaluation. Radiology 2000;217:165–72. [8] Whiting BR, McFarland EG, Dellabarca C, Brunsden BS. Study of CT artifacts in virtual colonoscopy. Proc SPIE 1999;3660:110 –6. [9] Wang G, Vannier MW. Stair-step artifacts in three-dimensional helical CT: an experimental study. Radiology 1994;191:79– 83. [10] Brink JA, Heiken JP, Balfe DM, Sagel SS, DiCroce J, Vannier MW. Spiral CT: decreased spatial resolution in vivo due to broadening of section-sensitivity profile. Radiology 1992;185:469 –74. [11] Polacin A, Kalender WA, Marchal G. Evaluation of sectionsensitivity profiles and image noise in spiral CT. Radiology 1992; 185:29–35. [12] Winder J, Sun Z, Kelly B, Ellis PK, Hirst DG. Abdominal aortic aneurysm and stent graft phantom manufactured by medical rapid prototyping. J Med Engng Technol 2002;26(2):75– 8. [13] Ogata I, Yamashita S, Sumi S, Nishiharu T, Mitsukazi K, Takahashi M. Pitfalls in image reconstruction of helical CT angiography: an experimental study. Comput Med Imaging Graph 1999;23:143–54. [14] Udupa JK. Three-dimensional visualisation and analysis methodologies: a current perspective. Radiographics 1999;19:783–806. [15] Siegel S. Nonparametric methods for the behavioural sciences. New York: McGraw-Hill; 1956. p. 229. [16] Springer P, Stohr B, Giacomuzzi SM, Bodner G, Klingler A, Jaschke W, zur Nedden D. Virtual computed tomography colonoscopy: artifacts, image quality and radiation dose load in a cadaver study. Eur Radiol 2000;10:183–7. [17] Dachman AH, Lieberman J, Osnis RB, Chen SY, Hoffmann KR, Chen CT, Newmark GM, McGill J. Small simulated polyps in pig colon: sensitivity of CT virtual colography. Radiology 1997;203: 427–30. [18] Summers RM, Shaw DJ, Shelhamer JH. CT virtual bronchoscopy of simulated endobronchial lesions: effect of scanning reconstruction, and display settings and potential pitfalls. AJR 1998;170:947–50. [19] Rodentwaldt J, Kopka L, Roedel R, Margos A, Grabbe E. 3D virtual endoscopy of the upper airway: optimization of the scan parameters in
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a cadaver phantom and clinical assessment. J Comput Assist Tomogr 1997;21:405 –11. Urban BA, Fishman EK. Tailored helical CT evaluation of acute abdomen. Radiographics 2000;20:725– 49. Fleischmann D, Rubin GD, Paik DS, Yen SY, Hilfiker PR, Beaulieu CF, Napel S. Stair-step artifacts with single versus multiple detectorrow helical CT. Radiology 2000;216:185–96. Klingenbeck-Regn KSS, Flohr T, Ohnesorge B, Ohnesorge B, Kopp AF, Baum U. Subsecond multi-slice computed tomography: basics and applications. Eur J Radiol 1999;31:110–24. Maintz D, Fischbach R, Juergens KU, Allkemper T, Wessling J, Heindel W. Multislice CT angiography of the iliac arteries in the presence of various stents: in vitro evaluation of artefacts and lumen visibility. Invest Radiol 2001;36(12):699–704.
Zhonghua Sun received his degree in medicine (MB) at the Harbin Medical University, China in 1989. He had worked in the Department of Radiology, Peking Union Medical College Hospital as a radiologist for 9 years after his graduation. He obtained his PhD from the University of Ulster, United Kingdom in November 2002. He is currently a lecturer in medical imaging at the University of Ulster. His research interests include 3D medical image processing, CT virtual endoscopy.
John R. Winder graduated from the University of Ulster in Physics and Chemistry in 1983. He attained an MSc in Physics in 1985 from University College of Wales. He worked as a Medical Physicist for 13 years and is now a lecturer in Medical Imaging at the University of Ulster. His research interests include medical rapid prototyping, 3D medical visualisation and virtual reality in medicine.
Barry E. Kelly qualified MB BCh BAO in 1984. Following basic surgical training and FRCS, he trained in the NI radiology scheme, attaining the FRCR and FFRRCSI in 1993. Subsequently, he was SHERT Fellow at the University of Glasgow from 1993 to 1995 and was subsequently appointed consultant at the Royal Victoria Hospital, Belfast. Currently his interests are in cross-sectional imaging. He is also programme director of the NI radiology training scheme, and lecturer in radiology at the Queen’s University, Belfast.
Peter K. Ellis received his degree in medicine (MB BCh) at The Queen’s University, Belfast in 1989. He specialised in Radiology with subsequent subspecialty training as a Fellow in Interventional Radiology at The University of California, Irvine 1996 – 1997. Currently, he is a consultant radiologist at The Royal Victoria Hospital, Belfast with particular interests in interventional and trauma radiology. He was involved in the creation of the aortic stenting service in Belfast and current research interests are very much based on endovascular topics.
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Peter T. Kennedy is a consultant radiologist in the Royal Victoria Hospital, Belfast. He obtained his medical degree from Oxford University in 1989 and initially trained in General Surgery until 1994. His Radiological training was then carried out in the teaching hospital programme in Northern Ireland. This was followed by a year as the Vascular and Interventional Radiology Fellow in the University of Texas Medical Branch, Galveston, Texas. Following that, he obtained his current post in 1999 and has since developed an interest in vascular imaging and recently developed interventional techniques such as aortic stent graft and radiofrequency ablation of liver malignancy.
David G. Hirst graduated in 1972 in Physiology from the University of St Andrews, and completed a PhD from the same institution in 1975. After a period of 5 years as a radiation biologist at the Cancer Research Campaign Gray Laboratory, he worked in the Department of Radiation Oncology at Stanford University Medical Centre in California as a Research Associate and then as Associate Professor until 1989, when he returned to the Gray Laboratory as a Group Head. In 1994, he took up the position and currently holds at the University of Ulster as Professor of Radiation Science. His research interests are in vascular physiology and cancer gene therapy.
Assessment of VIE image quality using helical CT angiography: in vitro phantom study Zhonghua Suna,*, John R. Windera, Barry E. Kellyb, Peter K. Ellisb, Peter T. Kennedyb, David G. Hirstc a
Room 15J 13, School of Applied Medical Sciences and Sports Studies, University of Ulster, Newtownabbey BT37 0QB, Northern Ireland, UK b Department of Radiology, Royal Victoria Hospital, Belfast BT12 6BA, Northern Ireland, UK c School of Biomedical Sciences, University of Ulster, Newtownabbey BT37 0QB, Northern Ireland, UK Received 15 April 2003; revised 28 July 2003; accepted 8 September 2003
Abstract The aim of this study is to quantify the effects of helical CT acquisitions parameters on the magnitude of three-dimensional stair-step artefacts, visualization of renal ostium and morphologies of suprarenal stents observed using virtual intravascular endoscopy. This was performed in a phantom of the human abdominal aorta with a stent graft in situ. Stair-step artefacts were quantified by measuring the standard deviation of signal intensity on surface shaded images and the influence of these artefacts on the visualization of arterial ostia and stent morphologies were assessed by three radiologists. The methodology may be used to optimise the CT system performance for helical CT angiography in aortic stent grafting. q 2003 Elsevier Ltd. All rights reserved. Keywords: Image quality; Artefacts; Stent; Computed tomography; Virtual endoscopy
1. Introduction Helical CT angiography (CTA) has been used widely in the evaluation of aortic stent grafting [1 – 3]. Helical CTA generated virtual intravascular endoscopy (VIE) is a rapidly evolving technology that permits interactive two-dimensional (2D) and three-dimensional (3D) visualization techniques to be used for the assessment of aortic stent grafts [4 –6]. This process involves the conversion of sets of 2D images into 3D representations that simulate views that may be obtained by conventional endoscopy. The final image quality is determined by several factors including slice thickness (beam collimation), pitch and image reconstruction interval. However, the acquisition step of CTA sets fundamental limits on the entire process [7]. It has been observed that stair-step artefacts caused by thick image slice acquisition can distort the depiction of polyp morphology in the colon [8]. Several studies have shown * Corresponding author. Tel.: þ44-289036-8121; fax: þ44-2890366028. E-mail address: [email protected] (Z. Sun). 0895-6111/$ - see front matter q 2003 Elsevier Ltd. All rights reserved. doi:10.1016/j.compmedimag.2003.09.001
that the image quality parameters of 3D helical CT obtained with a proper interpolation algorithm method either are better or are quite similar to the parameters of conventional 2D CT [9 –11]. However, stair-step artefacts associated with the interpolation may be generated [9]. A better understanding of this artefact may be important as future clinical applications of CTA extend beyond routine 2D image analysis with advances in 3D volumetric acquisition. Our purpose was to quantitatively assess the effect of specific single helical CT acquisition parameters on the stair-step artefact. We also aimed to qualitatively assess the artefact related to image quality in VIE of arterial ostia and aortic stents. We recognise that multislice CT is becoming more available and will provide increased image quality. However, we think that many imaging centres may only be able to perform single slice helical CT for some time. The measurement methodology for the artefact and results of this study may be used to optimise CT system performance for single slice helical CTA in aortic stent grafting. And the method of artefact quantification will equally apply to multislice systems.
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2. Materials and methods 2.1. Aorta phantom design A human aorta phantom was built with medical rapid prototyping [12]. A commercial aortic stent graft (Zenith, Cook, Europe, Denmark) was deployed inside the phantom to simulate a repaired aortic aneurysm (Fig. 1). The phantom was housed in a small container and filled with contrast medium. The CT attenuation of the phantom was measured 170 HU, therefore, the concentration of contrast medium was diluted to 450 HU which created a X-ray attenuation difference of 280 HU. This was similar to the difference in CT attenuation encountered between aorta vessel wall and enhanced blood in abdominal CT angiographic scanning. 2.2. CT unit and scanning parameters Helical CT scanning was performed on a Philips Medical Systems CT scanner (Tomoscan AV-E1). The rotation time of the CT scanner was 1 s and the tube voltage and current were 120 kVp and 125 mAs for the scanning protocol using 1.0 mm collimation and 120 kVp and 250 mAs for 3.0 and 5.0 mm collimation. The highest selectable X-ray tube current for this CT scanner was 125 mAs at 1.0 mm and 250 mAs for 3.0 and 5.0 mm. A field of view of 256 mm, matrix of 512 £ 512 and 1808 linear interpolation algorithm were used to reconstruct the images. This was similar to our routine abdominal CTA protocol. The phantom was placed parallel to the longitudinal axis of the table of CT scanner
[13]. Image data were acquired at collimations of 1.0, 3.0 and 5.0 mm, pitch 1.0, 1.5 and 2.0 and reconstruction interval of 0.5, 1.0, 2.0, 3.0 and 5.0 mm. A scout view was obtained to define the range of the phantom to be scanned. Scans were taken from the top of the model to the upper aspect of the aortic aneurysm. The total length of CT scanning was 70 mm. After scanning, the data were transferred to a SUN workstation for data processing using Analyze V 1.0 (Analyzedirect.com, Mayo Clinic, USA). All images were post-processed with and without data interpolation. Linear interpolation [14] was used to generate missing data between known surrounding points. A total of 54 data sets were generated. 2.3. Image analysis and assessment VIE images of the superior mesenteric artery (SMA), renal arteries and aortic aneurysm were generated from each volume data set. Two threshold ranges were chosen to create VIE images of the internal surface of the model and the stent wires. For images of the aortic ostia, the upper threshold was 320 HU whilst for the stent wire the lower threshold was 1200 HU. As many variables are involved in the quantification of stair-step artefacts based on the pixel value of the surfaced rendered VIE frames, consistent settings are required to keep the VIE images as same as possible in demonstrating each anatomical structure in all scanning protocols. Lighting and reflectivity are two important factors, which affect the results of measurements. The positioning of the virtual light source, which helps in the visualization of the 3D image and the level of the light that the surface reflects back to the viewing position were kept constant as any change of them produces different results (Fig. 2). Horizontal and vertical lights in our study were set 08 in all scanning parameters. The camera, ‘virtual eye position’ was kept at a constant angle of view (708) that was required to produce VIE images of the arterial ostia and the 3D relationship of stent struts to the arterial ostia in the same positions in all scanning protocols. 2.4. Quantitative image analysis
Fig. 1. (A) Stereolithography model of the human aorta. (B) Close-up view showing the stent deployed inside the model. (C) 3D CT image of the aorta phantom showing the aortic branches and aneurysm. It is noticed that the outer surface of the model was irregular due to manual editing when postprocessing the raw data. Arrow indicates the SMA, while short arrows indicate the renal arteries and arrowhead the fixing part inside the box. (D) Frontal view from the top of the model demonstrating suprarenal stent struts (red colour).
In an artefact free VIE image, the internal aortic wall should appear smooth. We required an objective numerical method for measuring the degree of stair-step artefact present. Line profile analysis produces a graphic display of pixel values from the VIE images. A plot of pixel intensity against pixel number enabled us to assess the degree of artefact present. Standard deviation (SD) of the line profile. To measure the degree of stair-step artefact, we used the SD of the pixel values, a higher SD indicating more variation in pixel intensity and therefore greater presence of artefact. The line profile was drawn across the images with the image intensity being recorded. The line was approximately
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Fig. 4. Measurements of the wire thickness on VIE image acquired at a collimation of 3.0 mm, pitch 1.0 (table speed of 3.0 mm per rotation) and reconstruction interval of 1.0 mm after interpolation. Five different locations were chosen to measure the transverse diameters of stent wires with thickness being 1.04, 1.23, 1.23, 1.29 and 1.29 mm, respectively, measured from the top to the bottom.
Fig. 2. Effect of the angle of virtual lightning on the measurements of SD through a line profile was demonstrated in non-interpolated VIE images of the right renal ostium acquired at a collimation of 5.0 mm, pitch 1.0 (table speed of 5.0 mm per rotation) and reconstruction interval of 3.0 mm. The SD was measured 35.5, 51.4 and 22.9, respectively, when the horizontal and vertical light was set at 08 (A), horizontal light 308 and vertical light 08 (B), horizontal light 08 and vertical light 308 (C).
107 –114 pixels long, with the first 100 pixels being used for analysis (Fig. 3). The measurements were performed at three different positions with two being near to the aortic ostia (SMA and renal ostia), the remaining one in the aortic lumen. Width of the artefacts. The purpose of this measurement was to determine whether there is any measurable difference in the width of the stair-step artefact generated by each scanning protocol. A line profile was recorded in the VIE images including three high and three low intensity areas. This included three stair-step artefacts (Fig. 3). Similar regions were chosen as to those in the measurement of SD. 2.5. Measurement of aortic stent strut thickness It was noted that the stent wire (strut) diameter in the CT image was not representative of the real wire thickness [12]; we wanted to determine the relationship of apparent strut diameter to the CT scanning protocol. Small objects of high contrast in CT imaging may become distorted due to
blurring by the system point spread function. Measurements of the wire diameter were made to determine its appearance versus actual thickness on VIE. The measurements were performed at two locations: renal artery and SMA levels. At each level, five measurements of the transverse diameter of the stent wire were made with an average thickness being calculated (Fig. 4). 2.6. Subjective image analysis The purpose of subjective image analysis was to determine the optimum CT scanning parameters from a clinical perspective. The 54 data sets were 3D rendered to create VIE images of the SMA, renal arteries and aortic aneurysm together with the 3D display of the relationship of aortic stents to the arterial ostia. All examinations were evaluated by three independent experienced reviewers (consultant radiologists). The reviewers were blinded to the scanning protocols and recorded their assessment of the following four aspects based on a 4-point scale scoring method † Clarity of the renal ostium. Score 1 indicated that the renal ostium was not visualized, 2 visualized but distorted, 3 visualised and normal and 4 visualized and perfect. † Definition of suprarenal component. Score 1 indicated the discontinued stent struts, 2 irregular, 3 smooth and 4 smooth with hooks visible. † Presence and degree of stair-step artefacts. Score 1 indicated marked artefacts, 2 moderate, 3 minimal and 4 no artefacts. 2.7. Statistical analysis
Fig. 3. Method to measure the SD through a line profile. A line is drawn on a VIE image (A) with the distance of around 100 pixels. The corresponding line profile shows the degree of the artefacts (B).
The results were analysed using SPSS 11.0 for Windows (SPSS, Inc) to determine the degree of agreement among reviewers. The agreement among the reviewers for the assessment of image quality of VIE was evaluated using the Kendall’s W test [15]. Kendall’s W ranges between 0 (no agreement) and 1 (complete agreement). A p value , 0.05 was considered to be a statistically significant difference.
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3. Results 3.1. Stair-step artefacts quantification The SD of line profile measurements from each scanning protocol was as follows. The mean SD for non-interpolated and interpolated data were 18.6– 32.3 and 5.89 – 21.12, respectively (Table 1). It is clear from these measurements that the magnitude of the stair-step artefacts was much smaller for interpolated data compared to non-interpolated data. However, the SD measured in scanning protocols of 1.0 mm collimation was slightly higher than that measured in 3.0 and 5.0 mm collimations. Table 2 shows that the average scores from the three reviewers for each scanning protocol ranged from 1.0 to 2.7 in non-interpolated and 1.3 to 3.3 in interpolated image data. We considered the score of 2 or 3 (moderate or minimal artefacts) to be acceptable for clinical diagnosis. This meant that all scans of 1.0 mm collimation and 3.0, 5.0 mm collimation with reconstruction interval of 1.0 mm were acceptable for diagnosis in non-interpolated data. While in interpolated data, the range of acceptable scanning protocols was expanded to include scans of 3.0 and 5.0 mm collimation with reconstruction interval of 2.0 mm. Figs. 5– 7 show the relationship of stair-step artefacts to collimation, pitch and reconstruction interval in
non-interpolated data, while Figs. 8 –10 show the same relationship in interpolated data. When measuring the width of the artefacts in all of the scanning protocols in both non-interpolated and interpolated data sets, it was noted that there was no significant difference in the width of artefacts in scans obtained with collimation of 1.0 and 3.0 mm, whether the data were interpolated or not ðp . 0:05Þ: However, the width of the artefacts was significantly smaller in interpolated data than that in noninterpolated data when the collimation was 5.0 mm with reconstruction intervals of 3.0 and 5.0 mm ðp , 0:05Þ: 3.2. Visualization of the renal ostium Table 2 shows that the mean scores from the subjective visual assessment of the renal ostium by three reviewers ranged from 1.0 to 4.0 and 1.7 to 4.0 in non-interpolated and interpolated data, respectively. A renal ostium was required to be smooth or of perfect configuration to be acceptable for clinical diagnosis. Therefore, it was noted that only a small range of scanning protocols were acceptable, including all scans of 1.0 mm collimation and 3.0 mm collimation with reconstruction interval of 1.0 mm in non-interpolated data. In contrast, following interpolation, the acceptable scanning ranges were all scans of 3.0 and 1.0 mm collimation. It was also noted that the visualization of the renal ostium was
Table 1 The mean SD measured in all scanning protocols in both non-interpolated and interpolated image data Scanning parameters
Mean standard deviation (SD)
Collimation (mm)
Pitch
Reconstruction interval (mm)
Non-interpolated data
Interpolated data
1.0 1.0 1.0 1.0 1.0 1.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0
1 1 1.5 1.5 2 2 1 1 1 1.5 1.5 1.5 2 2 2 1 1 1 1 1.5 1.5 1.5 1.5 2 2 2 2
0.5 1.0 0.5 1.0 0.5 1.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0
19.8 26 22.1 28 19.9 28 18.6 27.9 27.5 19.6 29.7 27.4 19.1 28 26.7 24.7 32.3 31 24 21.2 31.8 30.9 26.1 24.6 29.8 30.1 25.1
8.34 8.99 11.26 10.92 11.68 11.04 6.54 5.89 7.69 7.87 5.98 8.75 8.26 6.94 8.56 8.05 9.84 13.1 17.63 9.56 9.51 10.96 17.28 15.34 17.82 17.22 21.12
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Table 2 Mean ranks of subjective image quality for VIE images in the assessment of stair-step artefacts, renal ostium and aortic stent struts in all scanning protocols in both non-interpolated and interpolated image data Scanning parameters Collimation (mm)
1.0 1.0 1.0 1.0 1.0 1.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 3.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0 5.0
Pitch
1 1 1.5 1.5 2 2 1 1 1 1.5 1.5 1.5 2 2 2 1 1 1 1 1.5 1.5 1.5 1.5 2 2 2 2
Mean ranks of image quality Reconstruction interval (mm)
0.5 1.0 0.5 1.0 0.5 1.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0 1.0 2.0 3.0 5.0
Non-interpolated data
Interpolated data
Stair-step artefacts
Renal ostium
Aortic stents
Stair-step artefacts
Renal ostium
Aortic stents
2.3 2 2.7 2.7 2.3 2 2.3 1 1.3 2 1 1 2 1 1 2 1 1 1 2 1 1 1 1.7 1 1 1
4 3.7 3 3.3 4 3 3.3 2 2 3 2 2 3 2 2 3 2 2 2 2 1.7 3 2 1.7 1.3 1.3 1
3.3 3 2.7 3.3 3.3 2.3 3 2 1.7 2.7 2 1.7 2.7 2 1.7 2.7 1.7 1 1 1.7 1.3 1 1 2 1 1 1
3.3 3 1.7 2.7 2 2.7 2.3 2 1.7 3 2 2 2 2 1.7 2 2 1.7 1.3 2.3 2.3 2 1.3 3 1.7 1.7 1.7
4 3.7 3.3 3.7 3.3 3.7 3 3.3 3 3 3 3.3 3 3 3.3 1.7 1.7 1.7 1.7 1.7 2 1.7 1.7 1.7 1.7 1.7 1.7
3.7 3.7 3.7 3.3 3.7 3.7 3 3 1.7 3 2.3 2 1.3 3 1.7 1.3 1.3 1 1 1.3 1.3 1 1.7 1.7 1.7 1.7 1.7
Fig. 5. Non-interpolated VIE images of the right renal ostium obtained with a collimation of 1.0, 3.0 and 5.0 mm (A –C) with pitch of 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) and reconstruction interval of 1.0 mm. The artefacts in image C slightly distorted the renal ostium (arrows).
Fig. 6. Non-interpolated VIE images of the right renal ostium obtained with scanning protocols of 3.0 mm collimation, pitch 1.0, 1.5 and 2.0 (table speed of 3.0, 4.5 and 6.0 mm per rotation, respectively) with reconstruction interval of 1.0 mm (A –C). No apparent change of the artefacts was noticed with the increase of pitch values.
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Fig. 7. Non-interpolated VIE images obtained with a collimation of 5.0 mm, pitch 1.0 (table speed of 5.0 mm per rotation) with reconstruction interval of 1.0, 2.0, 3.0 and 5.0 mm (A –D). Artefacts slightly distorted the renal ostium in A and seriously compromised the renal ostia in B –D due to wide reconstruction intervals.
Fig. 9. Interpolated VIE images of the right renal ostium obtained with scanning protocols of collimation 3.0 mm, pitch 1.0, 1.5 and 2.0 (table speed of 3.0, 4.5 and 6.0 mm per rotation, respectively) and reconstruction interval of 1.0 mm. No apparent difference in the degree of stair-step artefacts was noticed.
determined by the collimation and reconstruction interval, however, it was independent of pitch values. Figs. 5, 7, 8 and 10 show examples of the effect of collimation and reconstruction interval on the visualization of the renal ostia in both non-interpolated and interpolated data. Figs. 6 and 9 show that the visualization of the renal ostium on VIE images was independent of pitch.
scanning protocols ranged from 1.0 to 3.3 and 1.0 to 3.7 in non-interpolated and interpolated data, respectively. A score of 3 indicated a smooth appearance of aortic stent struts visualized on VIE and was considered acceptable for determining the stent position. Most of the scans with 1.0 and 3.0 mm collimation, pitch 1.0 and reconstruction interval of 1.0 mm were acceptable in non-interpolated data. After interpolation, all scans with 1.0 mm collimation and most of the scans with 3.0 mm collimation were acceptable. All scans with 5.0 mm collimation were below the acceptable criteria as their scores ranged from 1.0 to 2.7. Figs. 11 and 12 show the effect of scanning protocols on the visualization of aortic stent struts in both non-interpolated
3.3. Visualization of the stent strut Table 2 shows that the average scores from the subjective assessment of the visualization of the aortic stent struts in all
Fig. 8. Interpolated VIE images of the right renal ostium obtained with scanning protocols of collimation 1.0, 3.0 and 5.0 mm, pitch 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) with reconstruction interval of 1.0 mm. The renal ostium (C) is slightly distorted when the collimation reached 5.0 mm.
Fig. 10. Interpolated VIE images obtained with scanning protocol of 5.0 mm collimation, pitch 1.0 (table speed of 5.0 mm per rotation) and reconstruction of 1.0, 2.0, 3.0 and 5.0 mm (A –D). Note that artefacts existing in these images make renal ostia distorted in all of these images.
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were interpolated or not. When two wires were very close together, they merged, producing an apparent thickness of 3.14 – 3.87 mm (Fig. 11C). Subjective assessment showed that there was significant evidence of concordance among the reviewers in the visualization of the renal ostium, stent struts (Kendall’s W ranged from 0.849 to 0.94, p , 0:05). Significant evidence of concordance was found among the reviewers in the assessment of stair-step artefacts in non-interpolated data (Kendall’s W 0.869, p , 0:05), while no significant evidence of concordance in interpolated data (Kendall’s W 0.401, p . 0:05). This could be because the degree of artefact was not as apparent in interpolated data as that observed in non-interpolated data, and reviewers scored most of the interpolated images as having moderate or minimal artefacts. Fig. 11. VIE images of aortic stents observed in the scanning protocols of 1.0, 3.0 and 5.0 mm collimation, pitch 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) with reconstruction interval of 1.0 mm (A –C) without interpolation. Details of aortic stent struts such as strut junctions (arrow in A) were only visualized in the scanning of 1.0 mm collimation.
and interpolated data. It was noted that details of the stent struts such as strut junctions can only be visualized in images of 1.0 mm collimation. 3.4. Thickness of the aortic stent wire The actual diameter of the stent wire was 0.4 mm. The thickness of the wire ranged from 1.64 to 3.14 mm on axial CT and from 0.82 to 1.29 mm on VIE. The measured thickness of the stent struts did not depend on the collimation used (Fig. 11), regardless of whether the data
Fig. 12. VIE images of aortic stents observed in the scanning protocols of 1.0, 3.0 and 5.0 mm collimation, pitch 1.0 (table speed of 1.0, 3.0 and 5.0 mm per rotation, respectively) and reconstruction interval of 1.0 mm (A –C) after interpolation. It is noted that the stents become slightly irregular when the collimation reached 5.0 mm (arrows).
4. Discussion Helical CTA is accepted as the imaging modality of choice for follow up of endovascular stent grafting. Most research has focused on the study of optimal scanning parameters for CT colonography in the detection of colonic polyps [16,17], or CT bronchoscopy in the detection of endobronchial lesions [18,19]. Evaluation of VIE in aortic stent grafting is limited [4 –6], and assessment of image quality of VIE in aortic stent grafting and optimal scanning parameters of CTA together with VIE images has not been reported. Image quality of the original CTA images affects the evaluation of aortic stent grafts as well as the relationship of aortic stents to the arterial branches. From a clinical point of view, 3.0 and 5.0 mm collimation are the most commonly used parameters in helical CTA of aortic stent grafting, therefore, we included these slice thicknesses in our phantom study. The reason we included 1.0 mm collimation in our study was that we considered it may improve spatial resolution, resulting in better visualization of small anatomical structures such as the renal ostium and details of aortic struts. One millimetre collimation is not routinely used in abdominal scanning because of increased image noise, decreased image quality, lack of anatomical coverage and X-ray dose [20]. Increased image noise was found in our study as it was shown that the SD of the stair-step artefacts measured in 1.0 mm scanning protocols was relatively higher than that measured in 3.0 and 5.0 mm collimation scans. However, image quality of VIE images in scanning protocols of 1.0 mm was scored better than that obtained in scans with wider collimation. Stair-step artefacts are a well-recognised problem in clinical CT imaging. However, explicit information on the degree of artefact in relation to acquisition parameters is scarce [21]. Our results showed that for helical CT generated VIE the magnitude of stair-step artefacts was generally related to the collimation and reconstruction interval used but independent of pitch values. This was
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found for both subjective and objective assessment of the VIE images. Therefore, a thinner collimation with a higher pitch, e.g. 2.0, may be suggested to produce good quality VIE images. For evaluation of aortic stent grafts, thinner collimation produced smoother aortic ostia and more detailed images of the stents wires. According to our study, scanning protocols with 1.0 and 3.0 mm collimation after interpolation produced the best VIE images for visualizing the ostium and stent wires. Our results concurred with those obtained with multislice CT regarding the effect of collimation and pitch values on the degree of artefacts [22]. They reported that a large pitch and narrower collimation is much more favourable for suppressing the artefacts. The quality of VIE images in aortic stent grafting was determined by the CT scanning parameters. An acceptable VIE image was required to demonstrate clearly the arterial ostia, stent wires and wire to ostia relationship. Previous studies have shown that in order to eliminate stair-step artefacts, both the collimation and the pitch should be less than the longitudinal dimension of the important feature on inclined surfaces, and the reconstruction interval should be less than the table speed [9,19]. Our study showed that a hybrid CT scanning protocol (1.0 mm collimation around the renal artery and 3.0 mm collimation with pitch 2.0 for the rest of the abdominal aorta) resulted in acceptable VIE image quality and better delineation of the arterial ostia, stent wires as well as wire to ostia relationship. One of the disadvantages of helical CTA in the imaging of stent grafts is the blurring effect that results in overestimation of the thickness of stent wires as displayed in both axial CT and VIE images. It was noted that nearly 50% of the renal ostium appeared covered by the stent wires on VIE images (Figs. 11 and 12). We already know only a small proportion of the renal ostium was covered in reality. It is very important to appreciate this when evaluating the interference of stent struts to the arterial ostia in clinical studies as the interference may cause effects on renal blood flow. Our results showed that the stent morphology depends on the collimation and is independent of pitch values. This did not concur in part with a recent study investigating the effect of artefacts on various stent grafts by multislice CT [23]. The collimation mainly influenced the sharpness of the image and delineation of the stent struts, which is consistent with our results. Their results also showed that higher pitch values caused more pronounced artefacts and influenced the size of the artefacts. The reason for this difference is that their study investigated the effects of scanning parameters on the axial and multiplanar reformatted images, while ours concentrated on virtual endoscopic images. There are several limitations in our study. Firstly, the phantom wall is much denser than the patient’s aortic wall with a CT number of around 170 HU, compared with about
40 HU for aortic tissue. Although, the contrast difference between the aortic wall and the inside medium is similar to the contrast enhanced blood, which was about 200 HU, images obtained from the phantom are slightly different from those of patient data. Secondly, we used static contrast material for imaging. In vivo, blood is flowing and contrast density varies along the vessel and these factors affect image quality. Thirdly, other factors such as the angulation of the aneurysm neck and the direction of the renal arteries in vivo could lead to individual variation of artery and aortic stent visualization even if the angle of view is kept constant. Therefore, in the processing of patient data, the virtual eye position should be customised to ensure the optimal visualization of specific anatomic structures. A potential limitation of phantom studies is that quantitatively obtained results of image quality cannot be translated directly into the in vivo applications. This is because of the difference between the environment in vitro and in vivo. However, in this study, we built the vessel phantom based on real patient data with an infrarenal (abdominal aortic aneurysm) AAA. The size of the aorta and its branches were the same size as in a real patient. The stent graft placed in the phantom was the same product used to treat patients with AAA in our clinical cohort. Therefore, results obtained from the phantom were applicable to clinical studies, as the VIE images of the arterial ostia and the 3D relationship of stent struts to the arterial ostia reflect a similar environment as in real patients.
5. Summary We have tested a series of scanning protocols of CTA in a human vessel phantom with the purpose of finding the optimal one in aortic stent grafting. The phantom was built from a typical aortic aneurysm and a commercial stent graft was deployed inside to simulate the environment to that in real patient’s treatment. Our study demonstrated that image quality of VIE images would benefit from linear interpolation. This dramatically decreased the effect of stair-step artefacts in 3D post-processing images. The beam collimation determines image quality of CTA in aortic stent grafting more than any other factors and image quality was independent of pitch values, according to both subjective and objective assessment. A thin collimation and a high pitch produced better image quality than a wide collimation and a small pitch in the assessment of stair-step artefacts, renal ostium, aortic stent struts as well as the 3D relationship of stent wire to renal ostium. For the purpose of visualizing the encroachment of stent struts to the arterial ostia, a hybrid protocol of slice thickness 1.0 mm around the renal artery and slice thickness of 3.0 mm with pitch 2.0 and reconstruction interval of 1.0 mm for the rest of the abdominal aorta is recommended to provide the greatest coverage with acceptable image quality for visualization of the aortic ostia and suprarenal stent struts on VIE.
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Acknowledgements We thank Mr Gil Stevenson for his assistance in the statistical analysis of the results.
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Zhonghua Sun received his degree in medicine (MB) at the Harbin Medical University, China in 1989. He had worked in the Department of Radiology, Peking Union Medical College Hospital as a radiologist for 9 years after his graduation. He obtained his PhD from the University of Ulster, United Kingdom in November 2002. He is currently a lecturer in medical imaging at the University of Ulster. His research interests include 3D medical image processing, CT virtual endoscopy.
John R. Winder graduated from the University of Ulster in Physics and Chemistry in 1983. He attained an MSc in Physics in 1985 from University College of Wales. He worked as a Medical Physicist for 13 years and is now a lecturer in Medical Imaging at the University of Ulster. His research interests include medical rapid prototyping, 3D medical visualisation and virtual reality in medicine.
Barry E. Kelly qualified MB BCh BAO in 1984. Following basic surgical training and FRCS, he trained in the NI radiology scheme, attaining the FRCR and FFRRCSI in 1993. Subsequently, he was SHERT Fellow at the University of Glasgow from 1993 to 1995 and was subsequently appointed consultant at the Royal Victoria Hospital, Belfast. Currently his interests are in cross-sectional imaging. He is also programme director of the NI radiology training scheme, and lecturer in radiology at the Queen’s University, Belfast.
Peter K. Ellis received his degree in medicine (MB BCh) at The Queen’s University, Belfast in 1989. He specialised in Radiology with subsequent subspecialty training as a Fellow in Interventional Radiology at The University of California, Irvine 1996 – 1997. Currently, he is a consultant radiologist at The Royal Victoria Hospital, Belfast with particular interests in interventional and trauma radiology. He was involved in the creation of the aortic stenting service in Belfast and current research interests are very much based on endovascular topics.
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Peter T. Kennedy is a consultant radiologist in the Royal Victoria Hospital, Belfast. He obtained his medical degree from Oxford University in 1989 and initially trained in General Surgery until 1994. His Radiological training was then carried out in the teaching hospital programme in Northern Ireland. This was followed by a year as the Vascular and Interventional Radiology Fellow in the University of Texas Medical Branch, Galveston, Texas. Following that, he obtained his current post in 1999 and has since developed an interest in vascular imaging and recently developed interventional techniques such as aortic stent graft and radiofrequency ablation of liver malignancy.
David G. Hirst graduated in 1972 in Physiology from the University of St Andrews, and completed a PhD from the same institution in 1975. After a period of 5 years as a radiation biologist at the Cancer Research Campaign Gray Laboratory, he worked in the Department of Radiation Oncology at Stanford University Medical Centre in California as a Research Associate and then as Associate Professor until 1989, when he returned to the Gray Laboratory as a Group Head. In 1994, he took up the position and currently holds at the University of Ulster as Professor of Radiation Science. His research interests are in vascular physiology and cancer gene therapy.