- Email: [email protected]
Radiation-reduction strategies in cardiac computed tomographic angiography
Clinical Radiology 65 (2010) 859e867
Contents lists available at ScienceDirect
Clinical Radiology journal homepage: www.elsevierhealth.com/journals/crad
Review
Radiation-reduction strategies in cardiac computed tomographic angiography C.A. Roobottoma, b, *, G. Mitchellb, G. Morgan-Hughesb a b
Peninsula Medical School, University of Plymouth, Plymouth, UK Derriford Hospital, Plymouth, UK
art icl e i nformat ion Article history: Received 2 February 2010 Received in revised form 7 April 2010 Accepted 21 April 2010
Ionizing radiation has long been known to increase the risk of cancer. X-rays and g-rays are officially classified as a carcinogen by the World Health Organization’s International Agency for Research on Cancer.1 Of the 5 billion imaging investigations performed worldwide two-thirds employ ionizing radiation.2 Diagnostic x-rays are the largest man-made source of radiation exposure to the general population, and computed tomography (CT) represents the largest proportion of these.3 Diagnostic CT has seen a dramatic increase in applications in the last two decades, not least in the higher dose applications. Whilst the increased use of CT has undoubtedly been of patient benefit, it inevitably will be associated with an increase in malignancy due to medical exposure. In fact a recent study from the USA has estimated that the CT examinations performed in 2007 could result in 29,000 future cancers based on current risk estimations.4 Whilst the numbers in the UK will be less (only 4 million examinations are performed compared to 70 million), it is clear that it is the responsibility of all radiologists to carefully examine their CT techniques and protocols with the aim to reduce the dose of examinations without compromising their accuracy. Cardiac computed tomographic angiography (CTA) initially was a very high dose application. However, both clinicians and CT system manufacturers have done a large amount of work to reduce dose. Dramatic changes have been achieved and the aim of this review is to highlight these. However, such developments are not exclusively applicable to cardiac CTA and many can be utilized in CT in general. Ó 2010 The Royal College of Radiologists. Published by Elsevier Ltd. All rights reserved.
Biological effects of radiation Biological effects of ionizing radiation are classified as deterministic or stochastic. Deterministic effects, such as skin injuries and cataract formation, occur predictably when dose exceeds a certain threshold. Stochastic effects (for example, the induction of cancer and germ cell mutations) occur with a probability that increases with dose. Deterministic effects * Guarantor and correspondent: C.A. Roobottom, Derriford Hospital, Derriford Road, Plymouth, PL17 7DY, UK. Tel.: þ44 0 1752 792185. E-mail address: [email protected] (C.A. Roobottom).
are exceptionally rare at doses delivered by CT systems and will not be considered further. Stochastic effects are believed to have no radiation dose threshold and therefore are associated with the low radiation doses delivered during CT. These effects have a long latent period, occur randomly and are dependant on the level and type of irradiation delivered, the tissue irradiated and the age of the patient. Stochastic risks are believed to be cumulative, with an increasing risk occurring with successive exposures. However, the risks for an individual CT examination must be taken in the context of natural background irradiation, which varies widely but is approximately 2.5 mSv per year in the UK.
0009-9260/$ e see front matter Ó 2010 The Royal College of Radiologists. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.crad.2010.04.021
860
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
Data on the risks of radiation has largely been drawn from the follow-up of 25,000 individuals exposed to atomic bomb explosions in 1945 5 and to 407,000 radiation workers from 15 countries.6 Whilst controversy still remains about the interpretation of the data, it is generally accepted that a linear no threshold risk for radiation exposure exists and, as radiologists, our aim should always be to minimize dose wherever possible. The influence of age and sex has been examined in those individuals exposed to nuclear explosion.5,7 Decreasing age of exposure clearly increases risk. At all ages women have approximately twice the risk than men who receive the same exposure. There is a strong relationship between age and exposure of breast tissue, with higher risk for those below 40 years.8
Measurement of radiation dose Absorbed dose is the mean energy imparted to the matter in a volume by ionizing radiation, divided by the mass of the matter in the volume. The SI unit of absorbed dose is the gray (Gy). The biological effect of a given absorbed dose varies depending on the type and quality of radiation emitted by the radionuclide or external radiation field and the organs that it traverses. Current International Commission on Radiological Protection (ICRP) terminology uses a radiation-weighting factor (wR) to normalize for the radiation effect. In cardiac imaging, as emissions are photons (nuclear cardiology), and external radiation (CT and CTA), the wR is 1. However, it has been suggested that for x-rays the weighting may be an underestimation and could be double this.9 In addition to the absorbed dose and type of radiation, the probability of stochastic effects varies depending on the organ/tissue irradiated. A second weighting factor, the tissue-weighting factor (wT), is used to normalize for this effect. Equivalent dose multiplied by wT is termed weighted equivalent dose. The sum of weighted equivalent dose over all organs or tissues in an individual is termed the “effective dose.” The ICRP has calculated the weighting threshold for each organ. This weighting has been adjusted in 1991 and 2007 based on current evidence. The major difference in the recent publication compared to previous ones is a higher wT for the female breast and a lower factor for the gonads. Gender-specific dosimetry has been lacking, as is the effect of body habitus. In addition, radiosensitivity of many organs, such as the breast, has been observed to decrease with age.10 Moreover, a long lag time is typical from acute radiation exposure to the development of malignancy; e.g., a 12-year minimum latency period from radiation exposure to excess breast cancer risk has been described in Japanese atomic bomb survivors, therefore relative risk and weighting factors need to be age, sex and build specific and so dose estimations will continue to evolve. CT system manufacturers use dose data derived from measurements made in head and body phantoms to determine a weighted CT dose index (CTDI) for each CT system model at all available selections of tube voltage (kVp), tube current (mA), and rotation time. The selected pitch value is then incorporated to produce a CT dose index called the
CTDIvol. Once the scan length is determined from the topogram, the appropriate CTDIvol is combined with the actual length scanned in the patient to calculate the doseelength product (DLP). Manufacturer’s estimates are potentially prone to error and physicists onsite should check these measurements periodically to ensure reliability. This is done by the use of a pencil ionization chamber placed in a cylindrical polymethylmethacrylate phantom (16 cm for head simulations and 32 cm for body). As it is impossible to measure the absorbed dose to individuals’ organs, all effective dose calculations are estimates. Estimates are made either with the use of physical phantoms or by mathematical simulations (the commonest being the Monte Carlo simulation). Software is available to estimate dose derived from these models for current CT systems with the use of specific imaging protocols. Although such methods are the most accurate available, they are time-consuming and assume an “average” build for the patient. Because of the time-consuming nature of the above calculations, the most common method of dose estimation uses conversion factors. These are region-specific weighting factors that are determined by Monte Carlo methods and averaged for multiple CT systems. These conversion factors utilize the organ-specific weighting factors for a given anatomic area. Where effective dose (E) is obtained by the product of doseelength product (DLP) and the relevant conversion factor (EDLP). E ¼ EDLP DLP A committee determines tissue-weighting factors by a subjective balance of cancer incidence, cancer mortality, life shortening, and hereditary risk. Weighting factors are averaged over sex and age and effective dose risk assessment is only appropriate to a 30-year-old hermaphrodite. As a result, the estimated risk of cancer may be a factor of 3 higher or lower when applied to a reference patient and will be more variable when applied to an individual.11 Another shortcoming of this approach is the assumption that there is a uniform weighting across a scan volume. For example, in cardiac imaging the chest conversion factor is commonly used. Yet it is usually the lower chest and upper abdomen that is scanned. This scan volume includes a disproportionately large volume of breast tissue and often a significant amount of liver. As both of these organs are radiosensitive (and the weighting factor for breast has recently been increased) the current common method of calculation is likely to be an underestimate. Another important issue, which is particularly relevant to cardiac CT, is the increasing use of sequential axial imaging (the so-called “step and shoot” technique). This will be described in detail below but in brief involves the use of an axial rather than a spiral technique for just a small portion of the cardiac cycle in end diastole (when the heart is relatively stationary). Using this technique the scanner rotates in a fixed position and only turns the tube current on in end diastole. The time the tube is on is just over half a gantry rotation. The position relative to the patient is based on the
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
electrocardiogram (ECG) not the position of the tube relative to the patient and as such the half gantry rotation could begin and end in any position. This is important as if the half gantry rotation is anterior it will irradiate the breasts, therefore the effective dose will be high (due to the breasts relative radio sensitivity). In contrast if the gantry is posteriorly placed, the breasts will receive a lower dose. As yet the weighting factors take no count of this and it is very difficult to know at what gantry position the irradiation was on from the data provided by the CT systems. So effective dose, although easy to calculate, is an imperfect dose descriptor and is likely to be superseded.12 An alternative approach has been described that uses measured or calculated organ radiation doses, applies them to age and sex-specific organ risk estimates from the BEIR (biological effects of ionizing radiation) VII report,13 and calculates an effective risk from the examination. Effective risk attempts to estimate the risk of developing cancer from the partial-body irradiation of the examination. Although this is a new approach requiring further evaluation, it has the potential to improve the quantification of risk from CT radiation exposure to individuals. One such analysis suggested that the lifetime cancer risk estimate of a cardiac CT (performed retrospectively without dose modulation) could be as high as 1:143 for a 20-year-old woman, compared to 1:3621 for an 80-year-old man.14 Such figures, which are not unique to cardiac CT, highlight the ongoing need for appropriate patient selection and dose-reduction techniques. These will be discussed below.
Technical considerations Cardiac CT is the most demanding of all CT applications. The coronary arteries are small (1e5 mm) and are almost continuously moving. Therefore, it follows that their accurate visualization requires both high temporal and spatial resolution. Spatial resolution is dependant on focal spot and detector size and the number of reconstructed views performed [increased by the use of “wobbling” the focal spot in either the z-axis (Siemens) or the xey axis (GE with their HD scanner)]. This resolution, as such, is fixed for any individual scanner (although the scanner may have a highresolution mode). Temporal resolution is fundamentally linked to gantry rotation. Currently 180 of gantry rotation are required to obtain sufficient information to reconstruct an image. This is obtained either rotating one-detector/tube combination through 180 , or a two-tube/detector combination (with dual headed systems) through 90 . It follows that increasing gantry rotation will increase temporal resolution. However, detector tube combinations are extremely heavy (and have got even heavier with the introduction of 256 and 320 detector array systems) and the ability to increase gantry rotation above a third of a second is limited with current technology. Temporal resolutions of 175 ms are thus typically achievable with a single source system (with a 350 ms gantry rotation) when all the data are obtained from one cardiac cycle. This is usually adequate for patients with heart rates below 65 beats per minute (b.p.m.). Obtaining the
861
relevant data over two or more cardiac cycles can increase temporal resolution further. These partial segmental reconstruction techniques (in two segment mode) collect half the data from one cardiac cycle and the rest from the second and, in so doing, halve temporal resolution compared to single segment mode (where all the data are from one cardiac cycle). However, such techniques rely on a very stable RR interval. If this is variable, data from different parts of the cardiac cycle are mixed and result in image degradation.
Effects of the cardiac cycle In order to minimize movement blur it is fundamentally important to reconstruct images of the coronary arteries when they are most stationary. This occurs when the ventricles are undergoing iso-volumetric contraction and relaxation, which occurs at end systole and end diastole. If two successive R waves in the ECG are defined as 0 and 100%, these windows occur around 40% for end systole and 75% for end diastole. Iso-volumetric contraction is relatively constant but the time the heart is stationary in diastole is indirectly proportional to the heart rate. The slower the heart rate, the longer this time window is and so the less chance of movement blur. This explains why good heart rate control improves image quality and the key reconstruction windows for coronary CT are centred around 75% and 40% of the RR interval.
Retrospectively gated cardiac CT Retrospectively gated cardiac CT utilizes a spiral acquisition to obtain the relevant data for reconstruction. The data required have to be sufficient to not only reconstruct an image, but to reconstruct an image at any phase of the cardiac cycle. For this reason the pitch required is low (0.2e0.5) and is varied according to the heart rate. The data are stored in conjunction with the ECG data and thus any image from any phase of the cardiac cycle can be reconstructed if required. This has a number of advantages. Images can be reconstructed through out the cardiac cycle to obtain functional data. The different parts of the cardiac cycle can be analysed (either manually or by automatic software) to visualize the coronaries (and other cardiac structures) when there is least movement blur. There is also usually enough data present to allow the removal of data (which is usually when an ectopic occurs) and still be able to reconstruct an image. For these reasons the majority of cardiac CT up until recently has been performed using this technique. However, the downside to such flexibility is the increased radiation dose. As dose is inversely proportional to pitch, the dose of cardiac CT (whose pitch may be as low as 0.2) can be five-times greater than a standard helical scan with a pitch of one. This has been reflected in the reported doses for cardiac CT, where doses up to 15.2e21.4 mSv have been reported.15 Such high radiation doses, particularly when applied to patients of low to intermediate risk (where the incidence of disease is low), have made workers question the riskebenefit balance of its use in such patients.16 For this reason radiation reduction strategies have been a priority for workers using cardiac CT and major advances
862
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
in technique have been made. Such advances have not necessarily been taken up by all clinicians using cardiac CT. This was highlighted by the recent PROTECTION 1 trial; a multicentre observational study of 50 sites performing cardiac CT.17 Despite a high proportion of university sites (42%) and possible selection bias from the centres willingness to take part in such a research study, there were dramatic differences in dose. Median doses ranged from 5e30 mSv with wide variation in the use of dose reduction strategies. For this reason it is appropriate to go through such techniques in detail.
Radiation reduction strategies ECG-linked tube current modulation
visualized at 75% of the RR interval and as such this is the only phase that will be required. Maximum tube current therefore should be centred at this point (i.e. 75%) and a narrow window should be used (say 10%). If this is done the image quality will be good for phases 70, 75, and 80% but be noisy elsewhere. However, it is important to stress that there is still phase data throughout the cardiac cycle, and it is possible to obtain ejection fractions and view wall-motion abnormalities by creating multiphase images throughout the cardiac cycle even on the noisy datasets, as the structures being evaluated are large (Fig. 2). By minimizing the maximum tube current window you are effectively reducing the dose to the patient to its maximum with ECG-dose modulation. Such tactics are effective in patients with slow regular heart rates. If a patient has a tachycardia (and in cardiac CT terms this means above 65 b.p.m.) or variation in the RR interval (as indicated by a changing heart rate on the ECG monitor) the first thing undertaken should be to reduce the heart rate if possible. This is most commonly done with the use of either oral or intravenous metoprolol. If the patient cannot receive beta blockade, or it is ineffectual, then the tube modulation protocol will need modification. In patients with faster heart rates not all coronary segments are likely to be seen in one phase. A combination of both end diastolic and end systolic images are likely to be required for full evaluation. For this reason maximum tube current will be required for 35% of the RR interval as well as 75%. Centring at 55% of the RR interval and widening the maximum tube current window to 40% of the cardiac cycle can achieve this. Such tube modulation is thus used for patients with heart rates above 65 b.p.m. or with variation of heart rate during breath-hold.
Retrospective cardiac CT was initially performed with a constant tube current. However, it is possible to vary the current that is applied relative to the ECG (Fig. 1). The rationale is that in patients with a slow steady heart rate it is possible to predict that all coronary segments will be free of motion blur in end diastole (that is around 75% of the RR interval). Therefore, maximum tube current is only required around this part of the cardiac cycle and can be reduced during the remainder. Typically maximum tube current has been applied during mid diastole (60e80% of the RR interval) and is reduced by between 50e80% for the remainder. Using such a strategy, radiation exposures can be reduced by 37% on 64-section CT systems18 in patients with stable heart rates while maintaining diagnostic accuracy. The degree of ECG-linked tube modulation can be varied. The amount of modulation required will be dependant on the patient’s absolute heart rate and the stability of the RR interval. In patients with a slow stable heart rate, with a stable RR interval, it is probable that all the coronaries will be
Prospective axial gating
Figure 1 Illustration of ECG dose modulation. The tube current in this case is maximum throughout diastole but reduced to the minimum possible for the rest of the cardiac cycle.
The main disadvantages of retrospective gating are that the tube current is always on, even with the most aggressive forms of dose modulation and the requirement for a low pitch. In order to overcome this, a more radiation-efficient method of acquisition was developed.19 This technique is performed axially rather than spirally and on a 64-section system involving obtaining a 4 cm block of axial images during a predetermined part of the cardiac cycle. During the next cardiac cycle the table moves on an increment of 3.5 cm and during the cycle after that obtains a further 4 cm block (Fig. 3) and so on till the appropriate volume has been obtained (usually three to four acquisitions for a coronary study). This so-called “step and shoot” technique has a number of advantages. Being an axial scan, the z-axis resolution is improved; the pitch is dramatically reduced (effectively to 0.875 due to the overlapping datasets); and the tube current is zero outside the predetermined window. The latter results in dramatic reductions in dose compared to axial scanning techniques and doses well below 5 mSv are regularly achieved.20e23 These results were also borne out in clinical practice in PROTECTION 1.17 The only disadvantages of the technique are the fact multi-sector reconstruction techniques (which use data from more than one cardiac cycle to create the image) have not been developed. This means the patient’s heart rate
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
863
Figure 2 These two images demonstrate the effect of dose modulation. The image on the left demonstrates image quality with full tube current (in end diastole), whilst the image on the right is the same section but with minimum tube current (at end systole). The image has much more quantum mottle but the cardiac chambers can easily be assessed to perform an ejection fraction.
needs to be well controlled (below 65 b.p.m.). As the data are reconstructed from several separate blocks of data (albeit with 5 mm overlap) step artefacts can also occur and so has been one of the drivers for very large detector CT systems. In reality they are rarely a problem and look very different to stenotic disease (Fig. 4). Although only applicable in patients with steady, slow heart rates, aggressive use of beta blockade opens this technique up to the majority of patients. A recent audit in our department showed 86% of patients were scanned in prospective mode. It is a common misconception that prospective gating does not allow multi-phase data analysis. This is incorrect. In order to reconstruct a single-phase image, 180 of gantry rotation is required. However, if gantry rotation is extended more phase data can be obtained. The larger the degree of “padding” the more phase data are obtainable. Just as with dose modulation, the centre of reconstruction and the amount of padding can be varied (up to 200 ms). Thus, one can use a small amount of padding centred around end diastole in patients with slow heart rates (to obtain small phase increments, say to reduce problems of step artefact) or in patients with higher or variable heart rates, move the centre to 55% and have maximum padding to obtain end systolic (35%) and end diastolic data (75%; Fig. 5). The penalty of such flexibility, of course, is increased radiation dose. The dose of maximum padding is roughly double that of using zero padding. This is because the tube is turned on in each location for approximately double the time. However, the use of such a protocol still offers significant dose savings over the use of retrospective spiral scanning, even with aggressive dose modulation.
Patient specific protocols Tube voltage and tube current Spiral imaging in the majority of CT applications benefits from the use of attenuation dependant, tube current modulation. This is the tube current, and therefore the photon flux is
adjusted relative to the attenuation of the different parts of the patient. For example, the attenuation of the patient is higher when the tube/detector combination is scanning in the coronal plane and is lower in the anteroposterior (AP) plane. To address this manufacturers have introduced programs that adjust the tube current depending on the attenuation of the object in both the transverse (x, y) and the longitudinal (z) directions to minimize either photon-starved or photon-rich projections. Such techniques have been shown to significantly reduce dose without degradation of image quality 24e26 and are used routinely in general imaging protocols. Unfortunately such techniques are not available for cardiac imaging at present. Therefore, if one protocol is used for cardiac patients it will be set up with a high enough tube current to cover large patients. Therefore, it follows that smaller patients will receive unnecessary doses. To avoid this, patient-specific protocols must be used in routine clinical practice. At present the most reliable method of doing this is the use of the body mass index (BMI), which is derived from the patient's height and weight, which should always be measured prior to imaging (patients are notoriously inaccurate, particularly when it comes to weight!). Although the BMI should generally be used to work out the appropriate tube voltage and tube current, the clinician always has the option to vary this according to the patient’s specific build. The size of the patient around mid-thorax is the most important, so ideally fat distribution, muscle mass, and breast size should be taken into account. Once the BMI has been calculated a specific tube voltage/ tube current combination should be prescribed. This will be dependant on the system used [for example, we have found that we can use a lower tube voltage technique in larger patients with our GE HD750 system than we can with our GE VCT-XT (because of the garnet detector technology in the HD750-see below)]. Therefore, these will need evaluation by each centre. However Table 1 serves to provide typical figures as used in our department.
864
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
Figure 3 The “step and shoot” technique. In contrast to retrospective imaging the step and shoot technique is performed using sequential axial scans with a small degree of overlap to construct the cardiac image.
The operator has the option of varying both the tube current and the tube voltage. Dose is proportional to tube current in a linear fashion and to the square of tube voltage. This means that tube voltage should be reduced in preference to tube current when possible. Typically, CT protocols use 120 kVp as the standard tube voltage. However, diagnostic image quality using 100 or 80 kVp has been used in small adults and paediatric patients18 and has been shown to be effective in cardiac CTA.27 The limitations of decreased tube voltage scanning are decreased penetration of the beam with a resultant need to increase the tube current. This means the technique is not applicable in larger patients. The use of low tube voltage techniques also results in an increase in subject contrast (as the beam has more photons approximating to the k edge of iodine). This allows a reduction of the iodine load to the patient and requires
that image visualization be performed with a wider window width.
Patient specific scan length
As dose a CTDI length scanned, it follows that the larger the volume of the patient that is scanned, the larger the dose received by the patient. In the PROTECTION 1 trial the average dose for a CTA was 12 mSv and the average scan volume was 12 cm.17 Therefore, for every centimetre of volume added (at least with retrospective scanning) a further millisievert of dose is given to the patient. The volume of acquisition must therefore be individually “prescribed” for each individual patient. When only the native coronaries are being examined the scan volume can either be “guessed” from the topogram or more accurate delineation can be made from the use of a low-dose calcium
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
865
Figure 4 Two examples of step artefact. This artefact is not easily confused with stenotic disease because of its typical appearance and the fact that it affects not only the coronary arteries but also all other adjacent structures.
score. We routinely perform such a low-dose non-contrast enhanced scan, prospectively gated, single-phase examination in all patients (typically at 100 mA). Not only does this give information about the degree of calcification present (which may affect the tube voltage and tube current used in the CTA) but also allows accurate identification of the coronary arteries to be made, which means scan volumes can be kept to a minimum. In patients who are being examined for evaluation of their graft patency, it is important that the type of grafts used at operation are known as this will effect the coverage required (Fig. 6). Patients with internal mammary grafts require coverage from the lung apex, whereas those with only saphenous vein grafts require a reduced coverage as these invariably arise from the ascending aorta.
Recent developments Iterative reconstruction techniques Conventional CT reconstruction techniques involve manipulation of the raw data utilizing simple mathematical models. The commonest algorithm used in cardiac CTA is known as filtered back projection. This requires half a gantry rotation of data to reconstruct an image and thus doubles the temporal resolution compared to techniques requiring a 360 gantry rotation. All reconstruction techniques until recently have assumed a point focal spot, a point detector, point voxel, a pencil beam, and perfect sampling of a simple object. Such assumptions have been necessary due to the high requirements on computing power for image reconstruction. Of course, in reality, the
Figure 5 The use of padding with the step and shoot technique. On the left zero padding is applied. This reduces radiation dose to its minimum but only provides enough data to reconstruct one phase of imaging. It can be used in patients with heart rates below 65 b.p.m. with a regular rhythm. On the right maximum padding is demonstrated (200 ms). The centre of the padding is placed at 55% of the RR interval to allow reconstruction of end systolic and end diastolic data. This technique is applicable to faster and more irregular rhythms.
866
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
Table 1 Tube current and tube voltage comparisons between garnet detector and conventional detector systems. BMI
<20 20e25 25e30 30e35 >35
Conventional scanner
Garnet detector scanner
Tube voltage (kV)
Tube current (mA)
Tube voltage (kV)
Tube current (mA)
100 120 120 120 140
450 500 600 650 650
100 100 120 120 140
270 330 360 390 390
focal spot, x-ray beam, detector, and voxel have definitive size and sampling is susceptible to statistical variation. Such assumptions result in a noisy final image, which have to be compensated for by the use of a higher photon flux and therefore dose to the patient. The ever-increasing speed of computer processing has now meant that more complex calculations are possible and more realistic, complex, calculations can be used without significantly compromising image reconstruction times. In addition, computational modelling programs allow the system to compare the data received to that which was expected, and produce further iterations of the image. These techniques reduce artefacts (for example from calcium and coronary stents) and are inherently less noisy.28 Therefore, it follows that by using iterative reconstruction, a comparable image quality can be obtained at reduced dose.29 The degree of iteration that can be performed is variable as a post-processing technique, but clearly the tube current applied (and typically reduced) by using increased iterative reconstruction can only be decided prior to scanning.
Empirically we have found that greater reductions of tube current (with subsequent increasing degrees of iterative reconstruction) can be performed with inherently high contrast examinations (for example, CT coronary angiography) than examinations where low contrast resolution is important (for example, the detection of liver metastases). Clearly diagnostic accuracy must not be compromised in the pursuit for lower dose. A multi-centre vendor sponsored trial is underway to help answer such questions but research in this area will be ongoing for some time.
Garnet detector technology When a photon hits a detector it produces scintillation. The efficiency of the detector is dependant on how quickly it reacts to the photon (primary speed) and how quickly the material recovers after being hit by the radiation (afterglow). Primary speed is critical to image quality. When the decay is slow, there is more information carry-over from view to view and from voxel to neighbouring voxel. The information carry-over blurs the image and reduces resolution. Software correction is employed to deal with this problem, but there is a penalty of increased noise or reduced signal-to-noise ratio. Afterglow is the secondary time component of the exponential decay of the output after the x-ray source is turned off. In practical terms, afterglow refers to the remaining light from the scintillator at several milliseconds or longer since the x-ray source is turned off. Afterglow carries a part of the signal from one view into the next view during scanning. Therefore high afterglow without correction will cause arching artefacts, extending from low attenuation anatomy into areas of higher attenuation and will decrease in-plane spatial resolution. In order to achieve good image quality, low afterglow is desired. For decades the detectors used in CT systems have been GOS (Gd2O2S). However, a new garnet-based detector has been developed, which has a primary speed 100 times faster than GOS and an afterglow 25% of that of GOS. This detector can be used in high-definition mode to improve resolution of structures30 but also, because of reduced inherent noise, should allow dose reductions when compared to standard detectors. In addition, its improved contrast resolution appears to allow greater use of low tube voltage techniques in patients with higher BMIs with the subsequent dose savings (see Table 1). The combined use of iterative reconstruction and garnet technology can result in a halving of dose compared to conventional detector technology.31
High-pitch spiral acquisition
Figure 6 CTA of a graft study. Note the patient has a left internal mammary graft, as well two saphenous vein grafts, which required an increase in the scan coverage to ensure full assessment. Had this not been performed the occluded subclavian would have been missed.
Dual-source CT has been available for several years and has allowed improved temporal resolution in both spiral and axial scanning; however, a new variation has recently been described. Instead of performing a conventional retrospective spiral scan, the scanner is linked to the patient’s ECG so the scan is triggered at the start of diastole. However, at this point the heart is scanned as a continuous dataset without phase information (as it would in any other type of scan) so it is imaged in various phases of the cardiac cycle. However, a very large pitch (3.2) is used. This allows a very fast table
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
feed. This, in turn, means that in patients with a slow heart rate the whole heart can be imaged whilst still in diastole (even though different slices will be at different parts of it). This technique brings dose reductions because of the high pitch used. This technique has shown significant dose reductions.32 As with the other new techniques discussed, its place in cardiac imaging is as yet to be defined.
Conclusions CT brings undoubted patient benefit and because of this the associated small risks have not been the primary focus of radiologists developing CT techniques. However, cardiac CTA began its life as a high-dose examination, viewed with suspicion by clinical cardiologists, partly because of the radiation dose it required compared to conventional coronary angiography. This focused both the clinicians and manufacturers involved into developing dose-reduction strategies. Such efforts have brought about remarkable achievements in a short space of time. We routinely perform CTA at a tenth of the dose we did 5 years ago at the same time as achieving significant improvements in image quality. Although some dose-reduction techniques are specific to CTA, there are others that are applicable to CT in general. With increasing public awareness of the risks of exposure to ionizing irradiation, it is vital that we all examine the protocols we use for CT examinations and are satisfied we can justify the parameters used. This article aims to help those performing cardiac CTA to ensure that they optimize the techniques used.
References 1. World Health Organisation’s International Agency for Research on Cancer. Overall evaluations of carcinogenicity to humans, list of all agents evaluated to date, http//monographs.iarc.fr/ENG/classification/ listagentsalphorder.pdf. 2. Picano E. Sustainability of medical imaging. Education and debate. BMJ 2004;328:578e80. 3. Berrington de Gonzalez A, Darby S. Risk of cancer from diagnostic x-rays: estimates for the UK and 14 other countries. Lancet 2004;363:345e51. 4. Berrington de Gonzalez A, Mahesh M, Kim K, et al. Projected cancer risks from computed tomographic scans performed in the United States in 2007. Arch Intern Med 2009;169:2071e7. 5. Pierce DA, Shimizu Y, Preston DL, et al. Studies of the mortality of atomic bomb survivors. Report 12, part 1: cancer. Radiat Res 1996;146:1e27. 6. Cardis E, Vrijheid M, Blettner M, et al. The 15-country collaborative study of cancer risk among radiation workers in the nuclear industry: estimates of radiation related cancer risks. Radiat Res 2007;167: 396e416. 25. 7. Brenner D, Elliston C, Hall E, et al. Estimated risks of radiation-induced fatal cancer from pediatric CT. AJR Am J Roentgenol 2001;176:289e96. 8. Brenner DJ. Radiation risks potentially associated with low-dose CT screening of adult smokers for lung cancer. Radiology 2004;231:440e8. 9. Committee to Assess Health Risks from Exposure to Low Levels of Ionizing Radiation; Nuclear and Radiation Studies Board. Division on Earth and Life Studies, National Research Council of the National Academies. Health risks from exposure to low levels of ionizing radiation: BEIR VII Phase 2. Washington, DC: The National Academies Press; 2006. 10. Thompson DE, Mabuchi K, Ron E, et al. Cancer incidence in atomic bomb survivors, part II: solid tumors, 1958e1987. Radiat Res. 1994;137(Suppl. 2):S17e614. 2007.
867
11. Martin CJ. Effective dose: how should it be applied to medical exposures? Br J Radiol 2007;80:639e47. 12. Brenner DJ. Effective dose: a flawed concept that could and should be replaced. Br J Radiol 2008;81:521e3. 13. National Research Council. BEIR VII: health risks from exposure to low levels of ionizing radiation. Washington, DC: The National Academies Press; 2006. 14. Einstein AJ, Henzlova MJ, Rajagoplana S. Estimating risk of cancer associated with radiation exposure from 64-slice computed tomography coronary angiography. JAMA 2007;298:317e23. 15. Mollet NR, Cademartiri F, van Mieghem CA, et al. High-resolution spiral computed tomography coronary angiography in patients referred for diagnostic conventional coronary angiography. Circulation 2005;112: 2318e23. 16. Kaufmann PA. Triggering: low-dose computed tomography coronary angiography with prospective triggering: a promise for the future. J. Am. Coll. Cardiol 2008;52:1456e7. 17. Hausleiter J, Meyer T, Hermann F, et al. Estimated radiation dose associated with cardiac CT angiography. JAMA 2009;301:500e7. 18. Hausleiter J, Meyer T, Hadamitzdy M, et al. Radiation dose estimates from cardiac multislice computed tomography in daily practice: impact of different scanning protocols on effective dose estimates. Circulation 2006;113:1305e10. 19. Hsieh J, Londt J, Vass M, et al. Step-and-shoot data acquisition and reconstruction for cardiac x-ray computed tomography. Med Phys 2006;33:4236e48. 20. Scheffel H, Alkadhi H, Leschka S, et al. Low-dose CT coronary angiography in the step-and- shoot mode: diagnostic performance. Heart 2008;94:1132e7. 21. Husmann L, Valenta I, Gaemperli O, et al. Feasibility of low-dose coronary CT angiography: first experience with prospective ECG gating. Eur Heart J 2008;29:191e7. 22. Maruyama T, Takada M, Hasuike T, et al. Radiation dose reduction and coronary assessability of prospective electrocardiogram-gated computed tomography coronary angiography: comparison with retrospective electrocardiogram-gated helical scan. JACC 2008;52:1450e5. 23. Shuman WP, Branch KR, May JM, et al. Prospective versus retrospective ECG gating for 64-detector CT of the coronary arteries: comparison of image quality and patient radiation dose. Radiology 2008;248:431e7. 24. Kalender WA, Wolf H, Suess C. Dose reduction in CT by anatomically adapted tube current modulation. II. Phantom measurements. Med Phys 1999;26:2248e53. 25. Kalender WA, Wolf H, Suess C, et al. Dose reduction in CT by on-line tube current control: principles and validation on phantoms and cadavers. Eur Radiol 1999;9:323e8. 26. Greess H, Wolf H, Baum U, et al. Dose reduction in computed tomography by attenuation-based on line modulation of tube current: evaluation of six anatomical regions. Eur Radiol 2000;10:391e4. 27. Pflederer T, Rudofsky L, Ropers D, et al. Image quality in a low radiation exposure protocol for retrospectively ECG-gated coronary CT angiography. AJR Am J Roentgenol 2009;192:1045e50. 28. Isola AA, Ziegler A, Schäfer D, Köhler T, Niessen WJ, Grass M. Motion compensated iterative reconstruction of a region of interest in cardiac cone-beam CT. Comput Med Imaging Graphics 2010;34:149e59. 29. Yang C, Dazhi G, Cong N, et al. Bayesian statistical reconstruction for low-dose x-ray computed tomography using an adaptive-weighting nonlocal prior. Comput Med Imaging Graphics 2009;33:495e500. 30. Min JK, Swaminathan RV, Vass M, et al. High-definition multidetector computed tomography for evaluation of coronary artery stents: comparison to standard-definition 64-detector row computed tomography. J Cardiovasc Comput Tomogr 2009;3:246e51. 31. Gosling O, Loader L, Venebles P, Roobottom C, Bellenger N, MorganHughes. A comparison of radiation doses between state of the art multislice computed tomography coronary angiography with iterative reconstruction, standard multi-slice computed tomography coronary angiography and invasive diagnostic coronary angiography. Heart 2010;96:922e6. 32. Lell M, Marwan M, Schepis T, et al. Prospectively ECG triggered highpitch spiral acquisition for coronary angiography using dual source CT: technique and initial experience. Eur J Radiol 2009;11:2576e83.
Contents lists available at ScienceDirect
Clinical Radiology journal homepage: www.elsevierhealth.com/journals/crad
Review
Radiation-reduction strategies in cardiac computed tomographic angiography C.A. Roobottoma, b, *, G. Mitchellb, G. Morgan-Hughesb a b
Peninsula Medical School, University of Plymouth, Plymouth, UK Derriford Hospital, Plymouth, UK
art icl e i nformat ion Article history: Received 2 February 2010 Received in revised form 7 April 2010 Accepted 21 April 2010
Ionizing radiation has long been known to increase the risk of cancer. X-rays and g-rays are officially classified as a carcinogen by the World Health Organization’s International Agency for Research on Cancer.1 Of the 5 billion imaging investigations performed worldwide two-thirds employ ionizing radiation.2 Diagnostic x-rays are the largest man-made source of radiation exposure to the general population, and computed tomography (CT) represents the largest proportion of these.3 Diagnostic CT has seen a dramatic increase in applications in the last two decades, not least in the higher dose applications. Whilst the increased use of CT has undoubtedly been of patient benefit, it inevitably will be associated with an increase in malignancy due to medical exposure. In fact a recent study from the USA has estimated that the CT examinations performed in 2007 could result in 29,000 future cancers based on current risk estimations.4 Whilst the numbers in the UK will be less (only 4 million examinations are performed compared to 70 million), it is clear that it is the responsibility of all radiologists to carefully examine their CT techniques and protocols with the aim to reduce the dose of examinations without compromising their accuracy. Cardiac computed tomographic angiography (CTA) initially was a very high dose application. However, both clinicians and CT system manufacturers have done a large amount of work to reduce dose. Dramatic changes have been achieved and the aim of this review is to highlight these. However, such developments are not exclusively applicable to cardiac CTA and many can be utilized in CT in general. Ó 2010 The Royal College of Radiologists. Published by Elsevier Ltd. All rights reserved.
Biological effects of radiation Biological effects of ionizing radiation are classified as deterministic or stochastic. Deterministic effects, such as skin injuries and cataract formation, occur predictably when dose exceeds a certain threshold. Stochastic effects (for example, the induction of cancer and germ cell mutations) occur with a probability that increases with dose. Deterministic effects * Guarantor and correspondent: C.A. Roobottom, Derriford Hospital, Derriford Road, Plymouth, PL17 7DY, UK. Tel.: þ44 0 1752 792185. E-mail address: [email protected] (C.A. Roobottom).
are exceptionally rare at doses delivered by CT systems and will not be considered further. Stochastic effects are believed to have no radiation dose threshold and therefore are associated with the low radiation doses delivered during CT. These effects have a long latent period, occur randomly and are dependant on the level and type of irradiation delivered, the tissue irradiated and the age of the patient. Stochastic risks are believed to be cumulative, with an increasing risk occurring with successive exposures. However, the risks for an individual CT examination must be taken in the context of natural background irradiation, which varies widely but is approximately 2.5 mSv per year in the UK.
0009-9260/$ e see front matter Ó 2010 The Royal College of Radiologists. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.crad.2010.04.021
860
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
Data on the risks of radiation has largely been drawn from the follow-up of 25,000 individuals exposed to atomic bomb explosions in 1945 5 and to 407,000 radiation workers from 15 countries.6 Whilst controversy still remains about the interpretation of the data, it is generally accepted that a linear no threshold risk for radiation exposure exists and, as radiologists, our aim should always be to minimize dose wherever possible. The influence of age and sex has been examined in those individuals exposed to nuclear explosion.5,7 Decreasing age of exposure clearly increases risk. At all ages women have approximately twice the risk than men who receive the same exposure. There is a strong relationship between age and exposure of breast tissue, with higher risk for those below 40 years.8
Measurement of radiation dose Absorbed dose is the mean energy imparted to the matter in a volume by ionizing radiation, divided by the mass of the matter in the volume. The SI unit of absorbed dose is the gray (Gy). The biological effect of a given absorbed dose varies depending on the type and quality of radiation emitted by the radionuclide or external radiation field and the organs that it traverses. Current International Commission on Radiological Protection (ICRP) terminology uses a radiation-weighting factor (wR) to normalize for the radiation effect. In cardiac imaging, as emissions are photons (nuclear cardiology), and external radiation (CT and CTA), the wR is 1. However, it has been suggested that for x-rays the weighting may be an underestimation and could be double this.9 In addition to the absorbed dose and type of radiation, the probability of stochastic effects varies depending on the organ/tissue irradiated. A second weighting factor, the tissue-weighting factor (wT), is used to normalize for this effect. Equivalent dose multiplied by wT is termed weighted equivalent dose. The sum of weighted equivalent dose over all organs or tissues in an individual is termed the “effective dose.” The ICRP has calculated the weighting threshold for each organ. This weighting has been adjusted in 1991 and 2007 based on current evidence. The major difference in the recent publication compared to previous ones is a higher wT for the female breast and a lower factor for the gonads. Gender-specific dosimetry has been lacking, as is the effect of body habitus. In addition, radiosensitivity of many organs, such as the breast, has been observed to decrease with age.10 Moreover, a long lag time is typical from acute radiation exposure to the development of malignancy; e.g., a 12-year minimum latency period from radiation exposure to excess breast cancer risk has been described in Japanese atomic bomb survivors, therefore relative risk and weighting factors need to be age, sex and build specific and so dose estimations will continue to evolve. CT system manufacturers use dose data derived from measurements made in head and body phantoms to determine a weighted CT dose index (CTDI) for each CT system model at all available selections of tube voltage (kVp), tube current (mA), and rotation time. The selected pitch value is then incorporated to produce a CT dose index called the
CTDIvol. Once the scan length is determined from the topogram, the appropriate CTDIvol is combined with the actual length scanned in the patient to calculate the doseelength product (DLP). Manufacturer’s estimates are potentially prone to error and physicists onsite should check these measurements periodically to ensure reliability. This is done by the use of a pencil ionization chamber placed in a cylindrical polymethylmethacrylate phantom (16 cm for head simulations and 32 cm for body). As it is impossible to measure the absorbed dose to individuals’ organs, all effective dose calculations are estimates. Estimates are made either with the use of physical phantoms or by mathematical simulations (the commonest being the Monte Carlo simulation). Software is available to estimate dose derived from these models for current CT systems with the use of specific imaging protocols. Although such methods are the most accurate available, they are time-consuming and assume an “average” build for the patient. Because of the time-consuming nature of the above calculations, the most common method of dose estimation uses conversion factors. These are region-specific weighting factors that are determined by Monte Carlo methods and averaged for multiple CT systems. These conversion factors utilize the organ-specific weighting factors for a given anatomic area. Where effective dose (E) is obtained by the product of doseelength product (DLP) and the relevant conversion factor (EDLP). E ¼ EDLP DLP A committee determines tissue-weighting factors by a subjective balance of cancer incidence, cancer mortality, life shortening, and hereditary risk. Weighting factors are averaged over sex and age and effective dose risk assessment is only appropriate to a 30-year-old hermaphrodite. As a result, the estimated risk of cancer may be a factor of 3 higher or lower when applied to a reference patient and will be more variable when applied to an individual.11 Another shortcoming of this approach is the assumption that there is a uniform weighting across a scan volume. For example, in cardiac imaging the chest conversion factor is commonly used. Yet it is usually the lower chest and upper abdomen that is scanned. This scan volume includes a disproportionately large volume of breast tissue and often a significant amount of liver. As both of these organs are radiosensitive (and the weighting factor for breast has recently been increased) the current common method of calculation is likely to be an underestimate. Another important issue, which is particularly relevant to cardiac CT, is the increasing use of sequential axial imaging (the so-called “step and shoot” technique). This will be described in detail below but in brief involves the use of an axial rather than a spiral technique for just a small portion of the cardiac cycle in end diastole (when the heart is relatively stationary). Using this technique the scanner rotates in a fixed position and only turns the tube current on in end diastole. The time the tube is on is just over half a gantry rotation. The position relative to the patient is based on the
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
electrocardiogram (ECG) not the position of the tube relative to the patient and as such the half gantry rotation could begin and end in any position. This is important as if the half gantry rotation is anterior it will irradiate the breasts, therefore the effective dose will be high (due to the breasts relative radio sensitivity). In contrast if the gantry is posteriorly placed, the breasts will receive a lower dose. As yet the weighting factors take no count of this and it is very difficult to know at what gantry position the irradiation was on from the data provided by the CT systems. So effective dose, although easy to calculate, is an imperfect dose descriptor and is likely to be superseded.12 An alternative approach has been described that uses measured or calculated organ radiation doses, applies them to age and sex-specific organ risk estimates from the BEIR (biological effects of ionizing radiation) VII report,13 and calculates an effective risk from the examination. Effective risk attempts to estimate the risk of developing cancer from the partial-body irradiation of the examination. Although this is a new approach requiring further evaluation, it has the potential to improve the quantification of risk from CT radiation exposure to individuals. One such analysis suggested that the lifetime cancer risk estimate of a cardiac CT (performed retrospectively without dose modulation) could be as high as 1:143 for a 20-year-old woman, compared to 1:3621 for an 80-year-old man.14 Such figures, which are not unique to cardiac CT, highlight the ongoing need for appropriate patient selection and dose-reduction techniques. These will be discussed below.
Technical considerations Cardiac CT is the most demanding of all CT applications. The coronary arteries are small (1e5 mm) and are almost continuously moving. Therefore, it follows that their accurate visualization requires both high temporal and spatial resolution. Spatial resolution is dependant on focal spot and detector size and the number of reconstructed views performed [increased by the use of “wobbling” the focal spot in either the z-axis (Siemens) or the xey axis (GE with their HD scanner)]. This resolution, as such, is fixed for any individual scanner (although the scanner may have a highresolution mode). Temporal resolution is fundamentally linked to gantry rotation. Currently 180 of gantry rotation are required to obtain sufficient information to reconstruct an image. This is obtained either rotating one-detector/tube combination through 180 , or a two-tube/detector combination (with dual headed systems) through 90 . It follows that increasing gantry rotation will increase temporal resolution. However, detector tube combinations are extremely heavy (and have got even heavier with the introduction of 256 and 320 detector array systems) and the ability to increase gantry rotation above a third of a second is limited with current technology. Temporal resolutions of 175 ms are thus typically achievable with a single source system (with a 350 ms gantry rotation) when all the data are obtained from one cardiac cycle. This is usually adequate for patients with heart rates below 65 beats per minute (b.p.m.). Obtaining the
861
relevant data over two or more cardiac cycles can increase temporal resolution further. These partial segmental reconstruction techniques (in two segment mode) collect half the data from one cardiac cycle and the rest from the second and, in so doing, halve temporal resolution compared to single segment mode (where all the data are from one cardiac cycle). However, such techniques rely on a very stable RR interval. If this is variable, data from different parts of the cardiac cycle are mixed and result in image degradation.
Effects of the cardiac cycle In order to minimize movement blur it is fundamentally important to reconstruct images of the coronary arteries when they are most stationary. This occurs when the ventricles are undergoing iso-volumetric contraction and relaxation, which occurs at end systole and end diastole. If two successive R waves in the ECG are defined as 0 and 100%, these windows occur around 40% for end systole and 75% for end diastole. Iso-volumetric contraction is relatively constant but the time the heart is stationary in diastole is indirectly proportional to the heart rate. The slower the heart rate, the longer this time window is and so the less chance of movement blur. This explains why good heart rate control improves image quality and the key reconstruction windows for coronary CT are centred around 75% and 40% of the RR interval.
Retrospectively gated cardiac CT Retrospectively gated cardiac CT utilizes a spiral acquisition to obtain the relevant data for reconstruction. The data required have to be sufficient to not only reconstruct an image, but to reconstruct an image at any phase of the cardiac cycle. For this reason the pitch required is low (0.2e0.5) and is varied according to the heart rate. The data are stored in conjunction with the ECG data and thus any image from any phase of the cardiac cycle can be reconstructed if required. This has a number of advantages. Images can be reconstructed through out the cardiac cycle to obtain functional data. The different parts of the cardiac cycle can be analysed (either manually or by automatic software) to visualize the coronaries (and other cardiac structures) when there is least movement blur. There is also usually enough data present to allow the removal of data (which is usually when an ectopic occurs) and still be able to reconstruct an image. For these reasons the majority of cardiac CT up until recently has been performed using this technique. However, the downside to such flexibility is the increased radiation dose. As dose is inversely proportional to pitch, the dose of cardiac CT (whose pitch may be as low as 0.2) can be five-times greater than a standard helical scan with a pitch of one. This has been reflected in the reported doses for cardiac CT, where doses up to 15.2e21.4 mSv have been reported.15 Such high radiation doses, particularly when applied to patients of low to intermediate risk (where the incidence of disease is low), have made workers question the riskebenefit balance of its use in such patients.16 For this reason radiation reduction strategies have been a priority for workers using cardiac CT and major advances
862
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
in technique have been made. Such advances have not necessarily been taken up by all clinicians using cardiac CT. This was highlighted by the recent PROTECTION 1 trial; a multicentre observational study of 50 sites performing cardiac CT.17 Despite a high proportion of university sites (42%) and possible selection bias from the centres willingness to take part in such a research study, there were dramatic differences in dose. Median doses ranged from 5e30 mSv with wide variation in the use of dose reduction strategies. For this reason it is appropriate to go through such techniques in detail.
Radiation reduction strategies ECG-linked tube current modulation
visualized at 75% of the RR interval and as such this is the only phase that will be required. Maximum tube current therefore should be centred at this point (i.e. 75%) and a narrow window should be used (say 10%). If this is done the image quality will be good for phases 70, 75, and 80% but be noisy elsewhere. However, it is important to stress that there is still phase data throughout the cardiac cycle, and it is possible to obtain ejection fractions and view wall-motion abnormalities by creating multiphase images throughout the cardiac cycle even on the noisy datasets, as the structures being evaluated are large (Fig. 2). By minimizing the maximum tube current window you are effectively reducing the dose to the patient to its maximum with ECG-dose modulation. Such tactics are effective in patients with slow regular heart rates. If a patient has a tachycardia (and in cardiac CT terms this means above 65 b.p.m.) or variation in the RR interval (as indicated by a changing heart rate on the ECG monitor) the first thing undertaken should be to reduce the heart rate if possible. This is most commonly done with the use of either oral or intravenous metoprolol. If the patient cannot receive beta blockade, or it is ineffectual, then the tube modulation protocol will need modification. In patients with faster heart rates not all coronary segments are likely to be seen in one phase. A combination of both end diastolic and end systolic images are likely to be required for full evaluation. For this reason maximum tube current will be required for 35% of the RR interval as well as 75%. Centring at 55% of the RR interval and widening the maximum tube current window to 40% of the cardiac cycle can achieve this. Such tube modulation is thus used for patients with heart rates above 65 b.p.m. or with variation of heart rate during breath-hold.
Retrospective cardiac CT was initially performed with a constant tube current. However, it is possible to vary the current that is applied relative to the ECG (Fig. 1). The rationale is that in patients with a slow steady heart rate it is possible to predict that all coronary segments will be free of motion blur in end diastole (that is around 75% of the RR interval). Therefore, maximum tube current is only required around this part of the cardiac cycle and can be reduced during the remainder. Typically maximum tube current has been applied during mid diastole (60e80% of the RR interval) and is reduced by between 50e80% for the remainder. Using such a strategy, radiation exposures can be reduced by 37% on 64-section CT systems18 in patients with stable heart rates while maintaining diagnostic accuracy. The degree of ECG-linked tube modulation can be varied. The amount of modulation required will be dependant on the patient’s absolute heart rate and the stability of the RR interval. In patients with a slow stable heart rate, with a stable RR interval, it is probable that all the coronaries will be
Prospective axial gating
Figure 1 Illustration of ECG dose modulation. The tube current in this case is maximum throughout diastole but reduced to the minimum possible for the rest of the cardiac cycle.
The main disadvantages of retrospective gating are that the tube current is always on, even with the most aggressive forms of dose modulation and the requirement for a low pitch. In order to overcome this, a more radiation-efficient method of acquisition was developed.19 This technique is performed axially rather than spirally and on a 64-section system involving obtaining a 4 cm block of axial images during a predetermined part of the cardiac cycle. During the next cardiac cycle the table moves on an increment of 3.5 cm and during the cycle after that obtains a further 4 cm block (Fig. 3) and so on till the appropriate volume has been obtained (usually three to four acquisitions for a coronary study). This so-called “step and shoot” technique has a number of advantages. Being an axial scan, the z-axis resolution is improved; the pitch is dramatically reduced (effectively to 0.875 due to the overlapping datasets); and the tube current is zero outside the predetermined window. The latter results in dramatic reductions in dose compared to axial scanning techniques and doses well below 5 mSv are regularly achieved.20e23 These results were also borne out in clinical practice in PROTECTION 1.17 The only disadvantages of the technique are the fact multi-sector reconstruction techniques (which use data from more than one cardiac cycle to create the image) have not been developed. This means the patient’s heart rate
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
863
Figure 2 These two images demonstrate the effect of dose modulation. The image on the left demonstrates image quality with full tube current (in end diastole), whilst the image on the right is the same section but with minimum tube current (at end systole). The image has much more quantum mottle but the cardiac chambers can easily be assessed to perform an ejection fraction.
needs to be well controlled (below 65 b.p.m.). As the data are reconstructed from several separate blocks of data (albeit with 5 mm overlap) step artefacts can also occur and so has been one of the drivers for very large detector CT systems. In reality they are rarely a problem and look very different to stenotic disease (Fig. 4). Although only applicable in patients with steady, slow heart rates, aggressive use of beta blockade opens this technique up to the majority of patients. A recent audit in our department showed 86% of patients were scanned in prospective mode. It is a common misconception that prospective gating does not allow multi-phase data analysis. This is incorrect. In order to reconstruct a single-phase image, 180 of gantry rotation is required. However, if gantry rotation is extended more phase data can be obtained. The larger the degree of “padding” the more phase data are obtainable. Just as with dose modulation, the centre of reconstruction and the amount of padding can be varied (up to 200 ms). Thus, one can use a small amount of padding centred around end diastole in patients with slow heart rates (to obtain small phase increments, say to reduce problems of step artefact) or in patients with higher or variable heart rates, move the centre to 55% and have maximum padding to obtain end systolic (35%) and end diastolic data (75%; Fig. 5). The penalty of such flexibility, of course, is increased radiation dose. The dose of maximum padding is roughly double that of using zero padding. This is because the tube is turned on in each location for approximately double the time. However, the use of such a protocol still offers significant dose savings over the use of retrospective spiral scanning, even with aggressive dose modulation.
Patient specific protocols Tube voltage and tube current Spiral imaging in the majority of CT applications benefits from the use of attenuation dependant, tube current modulation. This is the tube current, and therefore the photon flux is
adjusted relative to the attenuation of the different parts of the patient. For example, the attenuation of the patient is higher when the tube/detector combination is scanning in the coronal plane and is lower in the anteroposterior (AP) plane. To address this manufacturers have introduced programs that adjust the tube current depending on the attenuation of the object in both the transverse (x, y) and the longitudinal (z) directions to minimize either photon-starved or photon-rich projections. Such techniques have been shown to significantly reduce dose without degradation of image quality 24e26 and are used routinely in general imaging protocols. Unfortunately such techniques are not available for cardiac imaging at present. Therefore, if one protocol is used for cardiac patients it will be set up with a high enough tube current to cover large patients. Therefore, it follows that smaller patients will receive unnecessary doses. To avoid this, patient-specific protocols must be used in routine clinical practice. At present the most reliable method of doing this is the use of the body mass index (BMI), which is derived from the patient's height and weight, which should always be measured prior to imaging (patients are notoriously inaccurate, particularly when it comes to weight!). Although the BMI should generally be used to work out the appropriate tube voltage and tube current, the clinician always has the option to vary this according to the patient’s specific build. The size of the patient around mid-thorax is the most important, so ideally fat distribution, muscle mass, and breast size should be taken into account. Once the BMI has been calculated a specific tube voltage/ tube current combination should be prescribed. This will be dependant on the system used [for example, we have found that we can use a lower tube voltage technique in larger patients with our GE HD750 system than we can with our GE VCT-XT (because of the garnet detector technology in the HD750-see below)]. Therefore, these will need evaluation by each centre. However Table 1 serves to provide typical figures as used in our department.
864
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
Figure 3 The “step and shoot” technique. In contrast to retrospective imaging the step and shoot technique is performed using sequential axial scans with a small degree of overlap to construct the cardiac image.
The operator has the option of varying both the tube current and the tube voltage. Dose is proportional to tube current in a linear fashion and to the square of tube voltage. This means that tube voltage should be reduced in preference to tube current when possible. Typically, CT protocols use 120 kVp as the standard tube voltage. However, diagnostic image quality using 100 or 80 kVp has been used in small adults and paediatric patients18 and has been shown to be effective in cardiac CTA.27 The limitations of decreased tube voltage scanning are decreased penetration of the beam with a resultant need to increase the tube current. This means the technique is not applicable in larger patients. The use of low tube voltage techniques also results in an increase in subject contrast (as the beam has more photons approximating to the k edge of iodine). This allows a reduction of the iodine load to the patient and requires
that image visualization be performed with a wider window width.
Patient specific scan length
As dose a CTDI length scanned, it follows that the larger the volume of the patient that is scanned, the larger the dose received by the patient. In the PROTECTION 1 trial the average dose for a CTA was 12 mSv and the average scan volume was 12 cm.17 Therefore, for every centimetre of volume added (at least with retrospective scanning) a further millisievert of dose is given to the patient. The volume of acquisition must therefore be individually “prescribed” for each individual patient. When only the native coronaries are being examined the scan volume can either be “guessed” from the topogram or more accurate delineation can be made from the use of a low-dose calcium
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
865
Figure 4 Two examples of step artefact. This artefact is not easily confused with stenotic disease because of its typical appearance and the fact that it affects not only the coronary arteries but also all other adjacent structures.
score. We routinely perform such a low-dose non-contrast enhanced scan, prospectively gated, single-phase examination in all patients (typically at 100 mA). Not only does this give information about the degree of calcification present (which may affect the tube voltage and tube current used in the CTA) but also allows accurate identification of the coronary arteries to be made, which means scan volumes can be kept to a minimum. In patients who are being examined for evaluation of their graft patency, it is important that the type of grafts used at operation are known as this will effect the coverage required (Fig. 6). Patients with internal mammary grafts require coverage from the lung apex, whereas those with only saphenous vein grafts require a reduced coverage as these invariably arise from the ascending aorta.
Recent developments Iterative reconstruction techniques Conventional CT reconstruction techniques involve manipulation of the raw data utilizing simple mathematical models. The commonest algorithm used in cardiac CTA is known as filtered back projection. This requires half a gantry rotation of data to reconstruct an image and thus doubles the temporal resolution compared to techniques requiring a 360 gantry rotation. All reconstruction techniques until recently have assumed a point focal spot, a point detector, point voxel, a pencil beam, and perfect sampling of a simple object. Such assumptions have been necessary due to the high requirements on computing power for image reconstruction. Of course, in reality, the
Figure 5 The use of padding with the step and shoot technique. On the left zero padding is applied. This reduces radiation dose to its minimum but only provides enough data to reconstruct one phase of imaging. It can be used in patients with heart rates below 65 b.p.m. with a regular rhythm. On the right maximum padding is demonstrated (200 ms). The centre of the padding is placed at 55% of the RR interval to allow reconstruction of end systolic and end diastolic data. This technique is applicable to faster and more irregular rhythms.
866
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
Table 1 Tube current and tube voltage comparisons between garnet detector and conventional detector systems. BMI
<20 20e25 25e30 30e35 >35
Conventional scanner
Garnet detector scanner
Tube voltage (kV)
Tube current (mA)
Tube voltage (kV)
Tube current (mA)
100 120 120 120 140
450 500 600 650 650
100 100 120 120 140
270 330 360 390 390
focal spot, x-ray beam, detector, and voxel have definitive size and sampling is susceptible to statistical variation. Such assumptions result in a noisy final image, which have to be compensated for by the use of a higher photon flux and therefore dose to the patient. The ever-increasing speed of computer processing has now meant that more complex calculations are possible and more realistic, complex, calculations can be used without significantly compromising image reconstruction times. In addition, computational modelling programs allow the system to compare the data received to that which was expected, and produce further iterations of the image. These techniques reduce artefacts (for example from calcium and coronary stents) and are inherently less noisy.28 Therefore, it follows that by using iterative reconstruction, a comparable image quality can be obtained at reduced dose.29 The degree of iteration that can be performed is variable as a post-processing technique, but clearly the tube current applied (and typically reduced) by using increased iterative reconstruction can only be decided prior to scanning.
Empirically we have found that greater reductions of tube current (with subsequent increasing degrees of iterative reconstruction) can be performed with inherently high contrast examinations (for example, CT coronary angiography) than examinations where low contrast resolution is important (for example, the detection of liver metastases). Clearly diagnostic accuracy must not be compromised in the pursuit for lower dose. A multi-centre vendor sponsored trial is underway to help answer such questions but research in this area will be ongoing for some time.
Garnet detector technology When a photon hits a detector it produces scintillation. The efficiency of the detector is dependant on how quickly it reacts to the photon (primary speed) and how quickly the material recovers after being hit by the radiation (afterglow). Primary speed is critical to image quality. When the decay is slow, there is more information carry-over from view to view and from voxel to neighbouring voxel. The information carry-over blurs the image and reduces resolution. Software correction is employed to deal with this problem, but there is a penalty of increased noise or reduced signal-to-noise ratio. Afterglow is the secondary time component of the exponential decay of the output after the x-ray source is turned off. In practical terms, afterglow refers to the remaining light from the scintillator at several milliseconds or longer since the x-ray source is turned off. Afterglow carries a part of the signal from one view into the next view during scanning. Therefore high afterglow without correction will cause arching artefacts, extending from low attenuation anatomy into areas of higher attenuation and will decrease in-plane spatial resolution. In order to achieve good image quality, low afterglow is desired. For decades the detectors used in CT systems have been GOS (Gd2O2S). However, a new garnet-based detector has been developed, which has a primary speed 100 times faster than GOS and an afterglow 25% of that of GOS. This detector can be used in high-definition mode to improve resolution of structures30 but also, because of reduced inherent noise, should allow dose reductions when compared to standard detectors. In addition, its improved contrast resolution appears to allow greater use of low tube voltage techniques in patients with higher BMIs with the subsequent dose savings (see Table 1). The combined use of iterative reconstruction and garnet technology can result in a halving of dose compared to conventional detector technology.31
High-pitch spiral acquisition
Figure 6 CTA of a graft study. Note the patient has a left internal mammary graft, as well two saphenous vein grafts, which required an increase in the scan coverage to ensure full assessment. Had this not been performed the occluded subclavian would have been missed.
Dual-source CT has been available for several years and has allowed improved temporal resolution in both spiral and axial scanning; however, a new variation has recently been described. Instead of performing a conventional retrospective spiral scan, the scanner is linked to the patient’s ECG so the scan is triggered at the start of diastole. However, at this point the heart is scanned as a continuous dataset without phase information (as it would in any other type of scan) so it is imaged in various phases of the cardiac cycle. However, a very large pitch (3.2) is used. This allows a very fast table
C.A. Roobottom et al. / Clinical Radiology 65 (2010) 859e867
feed. This, in turn, means that in patients with a slow heart rate the whole heart can be imaged whilst still in diastole (even though different slices will be at different parts of it). This technique brings dose reductions because of the high pitch used. This technique has shown significant dose reductions.32 As with the other new techniques discussed, its place in cardiac imaging is as yet to be defined.
Conclusions CT brings undoubted patient benefit and because of this the associated small risks have not been the primary focus of radiologists developing CT techniques. However, cardiac CTA began its life as a high-dose examination, viewed with suspicion by clinical cardiologists, partly because of the radiation dose it required compared to conventional coronary angiography. This focused both the clinicians and manufacturers involved into developing dose-reduction strategies. Such efforts have brought about remarkable achievements in a short space of time. We routinely perform CTA at a tenth of the dose we did 5 years ago at the same time as achieving significant improvements in image quality. Although some dose-reduction techniques are specific to CTA, there are others that are applicable to CT in general. With increasing public awareness of the risks of exposure to ionizing irradiation, it is vital that we all examine the protocols we use for CT examinations and are satisfied we can justify the parameters used. This article aims to help those performing cardiac CTA to ensure that they optimize the techniques used.
References 1. World Health Organisation’s International Agency for Research on Cancer. Overall evaluations of carcinogenicity to humans, list of all agents evaluated to date, http//monographs.iarc.fr/ENG/classification/ listagentsalphorder.pdf. 2. Picano E. Sustainability of medical imaging. Education and debate. BMJ 2004;328:578e80. 3. Berrington de Gonzalez A, Darby S. Risk of cancer from diagnostic x-rays: estimates for the UK and 14 other countries. Lancet 2004;363:345e51. 4. Berrington de Gonzalez A, Mahesh M, Kim K, et al. Projected cancer risks from computed tomographic scans performed in the United States in 2007. Arch Intern Med 2009;169:2071e7. 5. Pierce DA, Shimizu Y, Preston DL, et al. Studies of the mortality of atomic bomb survivors. Report 12, part 1: cancer. Radiat Res 1996;146:1e27. 6. Cardis E, Vrijheid M, Blettner M, et al. The 15-country collaborative study of cancer risk among radiation workers in the nuclear industry: estimates of radiation related cancer risks. Radiat Res 2007;167: 396e416. 25. 7. Brenner D, Elliston C, Hall E, et al. Estimated risks of radiation-induced fatal cancer from pediatric CT. AJR Am J Roentgenol 2001;176:289e96. 8. Brenner DJ. Radiation risks potentially associated with low-dose CT screening of adult smokers for lung cancer. Radiology 2004;231:440e8. 9. Committee to Assess Health Risks from Exposure to Low Levels of Ionizing Radiation; Nuclear and Radiation Studies Board. Division on Earth and Life Studies, National Research Council of the National Academies. Health risks from exposure to low levels of ionizing radiation: BEIR VII Phase 2. Washington, DC: The National Academies Press; 2006. 10. Thompson DE, Mabuchi K, Ron E, et al. Cancer incidence in atomic bomb survivors, part II: solid tumors, 1958e1987. Radiat Res. 1994;137(Suppl. 2):S17e614. 2007.
867
11. Martin CJ. Effective dose: how should it be applied to medical exposures? Br J Radiol 2007;80:639e47. 12. Brenner DJ. Effective dose: a flawed concept that could and should be replaced. Br J Radiol 2008;81:521e3. 13. National Research Council. BEIR VII: health risks from exposure to low levels of ionizing radiation. Washington, DC: The National Academies Press; 2006. 14. Einstein AJ, Henzlova MJ, Rajagoplana S. Estimating risk of cancer associated with radiation exposure from 64-slice computed tomography coronary angiography. JAMA 2007;298:317e23. 15. Mollet NR, Cademartiri F, van Mieghem CA, et al. High-resolution spiral computed tomography coronary angiography in patients referred for diagnostic conventional coronary angiography. Circulation 2005;112: 2318e23. 16. Kaufmann PA. Triggering: low-dose computed tomography coronary angiography with prospective triggering: a promise for the future. J. Am. Coll. Cardiol 2008;52:1456e7. 17. Hausleiter J, Meyer T, Hermann F, et al. Estimated radiation dose associated with cardiac CT angiography. JAMA 2009;301:500e7. 18. Hausleiter J, Meyer T, Hadamitzdy M, et al. Radiation dose estimates from cardiac multislice computed tomography in daily practice: impact of different scanning protocols on effective dose estimates. Circulation 2006;113:1305e10. 19. Hsieh J, Londt J, Vass M, et al. Step-and-shoot data acquisition and reconstruction for cardiac x-ray computed tomography. Med Phys 2006;33:4236e48. 20. Scheffel H, Alkadhi H, Leschka S, et al. Low-dose CT coronary angiography in the step-and- shoot mode: diagnostic performance. Heart 2008;94:1132e7. 21. Husmann L, Valenta I, Gaemperli O, et al. Feasibility of low-dose coronary CT angiography: first experience with prospective ECG gating. Eur Heart J 2008;29:191e7. 22. Maruyama T, Takada M, Hasuike T, et al. Radiation dose reduction and coronary assessability of prospective electrocardiogram-gated computed tomography coronary angiography: comparison with retrospective electrocardiogram-gated helical scan. JACC 2008;52:1450e5. 23. Shuman WP, Branch KR, May JM, et al. Prospective versus retrospective ECG gating for 64-detector CT of the coronary arteries: comparison of image quality and patient radiation dose. Radiology 2008;248:431e7. 24. Kalender WA, Wolf H, Suess C. Dose reduction in CT by anatomically adapted tube current modulation. II. Phantom measurements. Med Phys 1999;26:2248e53. 25. Kalender WA, Wolf H, Suess C, et al. Dose reduction in CT by on-line tube current control: principles and validation on phantoms and cadavers. Eur Radiol 1999;9:323e8. 26. Greess H, Wolf H, Baum U, et al. Dose reduction in computed tomography by attenuation-based on line modulation of tube current: evaluation of six anatomical regions. Eur Radiol 2000;10:391e4. 27. Pflederer T, Rudofsky L, Ropers D, et al. Image quality in a low radiation exposure protocol for retrospectively ECG-gated coronary CT angiography. AJR Am J Roentgenol 2009;192:1045e50. 28. Isola AA, Ziegler A, Schäfer D, Köhler T, Niessen WJ, Grass M. Motion compensated iterative reconstruction of a region of interest in cardiac cone-beam CT. Comput Med Imaging Graphics 2010;34:149e59. 29. Yang C, Dazhi G, Cong N, et al. Bayesian statistical reconstruction for low-dose x-ray computed tomography using an adaptive-weighting nonlocal prior. Comput Med Imaging Graphics 2009;33:495e500. 30. Min JK, Swaminathan RV, Vass M, et al. High-definition multidetector computed tomography for evaluation of coronary artery stents: comparison to standard-definition 64-detector row computed tomography. J Cardiovasc Comput Tomogr 2009;3:246e51. 31. Gosling O, Loader L, Venebles P, Roobottom C, Bellenger N, MorganHughes. A comparison of radiation doses between state of the art multislice computed tomography coronary angiography with iterative reconstruction, standard multi-slice computed tomography coronary angiography and invasive diagnostic coronary angiography. Heart 2010;96:922e6. 32. Lell M, Marwan M, Schepis T, et al. Prospectively ECG triggered highpitch spiral acquisition for coronary angiography using dual source CT: technique and initial experience. Eur J Radiol 2009;11:2576e83.