3 Strategies for Dose Reduction and Improved Image Quality in MSCT

3 Strategies for Dose Reduction

and Improved Image Quality in MSCT

M. Kachelriess, S. Schaller, W. A. Kalender

M. Kachelriess, PhD

Institute of Medical Physics, Krankenhausstrasse 12, 91054 Erlangen, Germany

W. A. Kalender, PhD

Professor, Institute of Medical Physics, Krankenhausstrasse 12, 91054 Erlangen, Germany

S. Schaller, PhD

Siemens Medical Solutions, Siemensstrasse 1, 91301 Forchheim, Germany

CONTENTS

3.1 Introduction 35

3.2 Dose and Image Quality 36

3.3 Dose Reduction with Multidimensional Adaptive Filtering 36

3.4 Dose Reduction with Automatic Exposure

Control 38

3.5 Dose Reduction by User Education 40 3.6 Tube Voltage and Filtration 41 3.7 Penumbra Effects 42 3.8 Dose Information 43 3.9 Discussion 44

References 45

3.1 Introduction

Since the introduction of the fi rst computed tomog- raphy (CT) scanner in 1972, the importance of CT for the medical community has increased drastically.

Presently, whole-body scans with isotropic sub-mil- limetre resolution are acquired routinely during a single breath-hold.

The decisive step towards three-dimensional isotropic data was the introduction of spiral CT in 1989 by W. A. Kalender. This scan mode is based on a continuous rotation of the gantry while simul- taneously translating the patient along the axis of rotation. The scan trajectory is a spiral and allows for truly 3D data acquisition. The z-interpolation step allows one to select the longitudinal position (z-position) of the reconstructed images arbitrarily and retrospectively. The continuous axial sampling

allows for high-quality 3D displays and led to a renaissance of CT (Kalender 2001). The introduc- tion of multi-slice spiral CT (MSCT) in 1998, which allows simultaneous scanning of currently up to 16 slices, further improved the scanner’s volume coverage, z-resolution and scan speed. For example, typical chest exams are carried out with collimations of 1 ×5 mm in 36 s with single-slice, 4×1 mm in 30 s with 4-slice and 16 ×0.75 mm in 10 s with 16-slice scanners (Fig. 3.1).

Recently, patient dose has become an increasingly more important issue: increasing the information density of a scan, e.g. the spatial resolution, requires to increase the photon density (absorbed photons per volume element) by at least the same factor if the signal-to-noise ratio (SNR) is to be maintained. Cur- rently, less than 5% of all X-ray exams are performed with CT. The cumulative dose of CT, however, is of the order of 40% of all X-ray exams. Dose is espe- cially critical in paediatric CT (Brenner et al. 2001).

Practical hardware- and software-based solutions to reduce patient dose are strongly required.

Fig. 3.1. Three generations of spiral CT scanners. The sizes of

the detector elements indicate the slice thickness of single-

slice, 4-slice and 16-slice scanners for typical exams. These

slice dimensions (or “logical” detector dimensions) may differ

from the physical dimensions since detector slices are often

made up by electronically combining several detector rows

This paper gives an overview of CT dose issues and of strategies to reduce the patient dose or, alter- natively, to improve the image quality for a given dose level. Note that in all attempts to reduce dose, image quality is crucial since the diagnostic outcome must not be jeopardized.

3.2 Dose and Image Quality

Image quality is a function of dose. We can quantify image quality in CT mainly by giving two parameters:

(a) the pixel noise σ that is measured as the standard deviation of the pixel values within a homogeneous image region; and (b) the spatial resolution. Due to the circular symmetry of CT scans one usually dis- tinguishes between the in-plane resolution ∆r and the axial- or z-resolution S _eff (effective slice thickness). For example, an image noise of σ = 30 HU, an in-plane reso- lution of ∆r=0.7 mm and a z-resolution of S eff =0.8 mm are typical values for the present standard exams. Please note that the resolution values are not identical to the voxel size of the image data which may be chosen to be arbitrarily small during image reconstruction.

Dose is usually quantifi ed by specifying the organ dose or the effective dose (values in mSv). For our pur- poses, it suffi ces to regard the effective tube current time product (I·t) _eff , i.e. the “effective mAs value”, as being the characteristic dose parameter since this parameter is proportional to the physical dose values. It is defi ned as the product of the tube current with the time under which a given z-position (slice) is exposed to X-rays;

therefore, the effective mAs depends on the scan’s pitch value p and on the rotation time t _rot as follows:

Recall that the pitch value is defi ned as

where d is the table increment per rotation, S is the collimated (nominal) slice thickness and M is the number of simultaneously scanned slices. Its inverse p ^–1 gives the number of rotations that cover a given z-position. Consider a scan with a table increment of 6 mm and a collimation of 16 ×0.75 mm. This situa- tion corresponds to a pitch of 0.5 and, consequently, each z-position is exposed to X-rays during two full p = M . S

d

rotations (twofold overlapping scan). Further assume a rotation time of 0.4 s and a tube current of 200 mA.

Then, according to the defi nition above, the effective mAs value is 160 mAs. It is this effective mAs value, and not the tube current itself, that is directly related to the image noise σ. The present scanners make use of this fact by allowing the user to specify the effec- tive mAs value. Internally, the scanner computes the correct tube current from the specifi ed effective mAs value and the selected pitch. Consequently, dose and image noise and the effective mAs value do not depend on the spiral pitch, a fact that has often been misunderstood.

Quantitatively, the proportionality

(1)

holds. The constant of proportionality depends in a complicated way on the patient size and density, on the tube voltage and on the prefi ltration.

Prior to the scan, the user specifi es the three parameters effective mAs value, effective slice thick- ness and the in-plane resolution (via the name of the reconstruction kernel) at the scanner console. He is thereby aiming at his preferred image quality. The image noise σ is the only dependent parameter since it cannot be controlled directly.

Equation has important consequences: increasing the spatial resolution requires to dramatically increase the effective mAs value and thereby the dose given that the image noise σ is to be held constant. For example, doubling spatial resolution (i.e. reducing S _eff and ∆r to 50%) requires to increase patient dose by a factor of 16 if the pixel noise shall remain at its original level. Since such dose increases are hardly acceptable, scan proto- cols must carefully balance between spatial resolution and image noise; therefore, many applications are designed to either aim at good low-contrast resolution (low image noise) or to produce high spatial resolution.

The fi rst is the case in tumour detection tasks where lower spatial resolution is acceptable. In contrast, lung exams and, especially, exams of the inner ear, require to scan with the highest spatial resolution available.

3.3 Dose Reduction with Multidimensional Adaptive Filtering

To allow for even higher image quality in the future, new software- and hardware-based methods that σ ² ∝

(I . t) _eff . S _eff . ∆r ³ 1

(I . t) _eff = I ₀ . t _rot

p .

make better use of the applied dose have been fur- nished.

This includes so-called adaptive fi ltering approaches that seek to reduce image noise either by directly manipulating the reconstructed images or by perform- ing dedicated raw-data preprocessing prior to image reconstruction (Eklundh and Rosenfeld 1981;

Lauro et al. 1990; Keselbrener et al. 1992; Hsieh 1994, 1998) . Directly manipulating images turns out to be less effi cient and more critical, even if the fi lters are applied locally (adaptively), since respective algo- rithms tend to smooth parts of the image; therefore, raw-data-based approaches, as they are often provided by the manufacturers, are preferred.

A highly effective method is multi-dimensional adaptive fi ltering (MAF) which is a raw-data-based algorithm using two detector dimensions plus the view dimension (rotation direction) to locally smooth data.

The MAF allows to reduce image noise in cross sections of high eccentricity such as the pelvis, thorax or shoul- der without affecting spatial resolution. In contrast to the said image-based smoothing approaches (includ-

ing the selection of smooth reconstruction kernels), the raw-data-based MAF operates only on a tiny frac- tion of the acquired raw data. The MAF selects those attenuation values for fi ltering that suffer from photon starvation due to high attenuation. Typically, 5% of the data are modifi ed; 95% of the acquired data remain unchanged (Kachelriess and Kalender 2000;

Kachelriess et al. 2001b).

The MAF can be used to remove the well-known streak-like noise artefacts that are oriented along the direction of highest attenuation (typically the lateral direction). Low-contrast objects are often completely hidden by this kind of structured and correlated noise. Increasing the tube current to avoid these artefacts would be the wrong strategy since global improvements are not required. The effect of apply- ing MAF is illustrated in Fig. 3.2. The patient exhibits severe streak artefacts in the shoulder region due to photon starvation. They are largely removed by MAF. Even more, the difference images show no patient structure. This proves that MAF does not simply smooth the image. The MAF instead man-

Fig. 3.2. Noise in the original data (top row) is reduced by up to 50% with multi-dimensional adaptive fi ltering (MAF; middle

row). Since streaks are oriented laterally, the transaxial slice and the coronal MPR show great improvements, whereas there are

no modifi cations in the anteroposterior direction and, consequently, the sagittal reformation is hardly modifi ed. The subtraction

images (bottom row) prove that MAF does remove noise but not structure details

ages to reduce noise and to remove the correlated noise without affecting spatial resolution; thereby, the visibility of low-contrast objects can be greatly improved (Baum et al. 2000).

3.4 Dose Reduction with Automatic Exposure Control

Similar to conventional radiography, an automatism to obtain a defi ned image quality is needed for CT. The present CT scanners allow to specify the spatial resolu- tion, but it is not possible to specify the image noise.

One must instead set the effective mAs product which infl uences image noise only indirectly. The outcome is strongly dependent on patient size, density and shape.

The dependence on patient size and its implica- tions on dose values are best illustrated using a numerical example: a 20 cm water-equivalent object (representing an adult head) attenuates the X-rays by a factor of approximately 50 and only 2% of the pho- tons reach the detector. For a small object of half size (representing a child head), the attenuation factor decreases to 7 and it becomes likely that too many quanta reach the detector. Due to this exponential behaviour, it is hard to manually extrapolate scan parameters from the standard adult patient to small children. This fact is one of the major causes for using too high doses in paediatric CT.

A fi rst step towards AEC has been the introduc- tion of the dose modulation technique. It modulates the tube current during each rotation as a function of patient anatomy to compensate for the different attenuation properties; typically, for anteroposterior projections the tube current is decreased and for lat- eral projections the tube current is increased (Gies et al. 1999; Kalender et al. 1999b).

It can be shown that the optimal way to modulate the tube current as a function of the rotation angle α is proportional to the square root of the attenuation p( α) measured at that angle (Gies et al. 1999; Harpen 1999) :

To achieve optimal modulation, attenuation values must be known. The most effective (in terms of dose utilization) implementations use attenuation values that have been measured 180° in advance to control the tube at the current projection on-line. Figure 3.3 illustrates this principle by giving typical quantum number values. Typical mAs-reduction values that can be achieved with this on-line dose modulation technique are given in Table 3.1, taken from Kalen- der et al. (1999c) and Greess et al. (2000).

These mAs-reduction values have been achieved without increasing image noise. It further turns out that the noise structure of the images acquired with dose modulation is less pronounced and less corre- lated than the noise structure of the standard acquisi- tion, thereby indicating improved image quality. Even more, Monte Carlo calculations of the organ dose and effective dose with and without dose modulation have shown that the mAs-reduction values under- estimate the physical dose reduction (Kalender et I ₀ (α) ∝ e ^{1 2 p(α)} .

Fig. 3.3. Scans with constant tube current (red) suffer from extremely different attenuation properties within eccentric objects. Here, photon starvation occurs in the lateral direction while more than suffi ciently many photons are measured anteroposteriorly. Tube current modula- tion (green) greatly reduces these dis- crepancies by modulating the tube cur- rent as a function of anatomy. In total, it reduces the mAs product (here by 30%) and, consequently, patient dose without increasing image noise

Table 3.1. Typical mAs-reduction values that can be achieved with on-line dose modulation technique

Region mAs reduction (%)

Head 18

Shoulder 53

Thorax 22

Abdomen 15

Pelvis 25

Extremity 39

al. 1999c; Schmidt and Kalender 2001, 2002): the values given in the table are only a conservative esti- mate of the true dose reduction achieved with tube current modulation.

Besides the on-line dose modulation techniques, other implementations make use of topogram (or scout view) data from an anteroposterior and a lateral view (Kopka et al. 1995). Sinusoidal curves are fi tted to the topogram data to predict the angle-dependent attenuation for all α; however, besides the need for two topograms, the mAs-reduction values reported are signifi cantly lower than with the on-line method (Kalender et al. 1999c).

All dose-modulation methods have in common that there are only local (rotation-wise) adaptations of the current. Variations of the mean attenuation properties in the longitudinal direction are not taken into account as the tube current’s mean or maximum value is retained after each rotation.

The remaining step towards AEC is to combine the tube current modulation with a tube current control.

The tube current control slowly varies the current’s mean value according to the patient anatomy and thus is able to compensate for varying image noise along the z-direction. The AEC comprises a method to “foresee” the image noise as a function of anatomy.

Then, the tube current can be controlled to achieve a given target noise. Again, there are topogram-based implementations and more sophisticated raw-data- based implementations (Kachelriess et al. 2001a).

In the future, scanners will make use of AEC to assist the user in adapting the tube current to patient anat- omy or even to replace the tube current parameter in favour of directly specifying image noise and hence image quality.

Figure 3.4 shows the tube current curve and the resulting image noise as a function of the anatomi- cal level during a spiral 16-slice scan from neck to thorax. The red and blue graphs correspond to the conventional case: the mean level of the tube current is held constant. The noise curves (red is hidden by blue) show excellent image quality in the neck and insuffi cient image quality in the shoulder. Without AEC, image quality can only be increased by increas- ing the tube current. With AEC, in contrast, it is pos- sible to accept more noise in the neck and to reduce noise in the thorax at the same time. Assume that the user specifi es a constant noise level corresponding to the green horizontal curve. Then, AEC controls and modulates the tube current to achieve the desired level (green oscillating curve). The running mean (green slowly varying curve) gives the local mAs- reduction value as a function of the z-position.

The phantom experiment of Fig. 3.5 demonstrates the performance of AEC. The step phantom consists of three elliptical cylinders of decreasing diameter.

The standard acquisition (constant tube current) suf- fers from varying image quality along the z-axis: high noise in the large cylinder; acceptable noise in the medium cylinder; and unnecessary low noise in the small cylinder. The three axial images further suffer from streak artefacts due to correlated noise. The AEC scan, in contrast, is as specifi ed: constant noise for the complete scan range and reduced noise correlations in the axial images. Besides constant noise, one may also specify other noise profi les, if desired.

Fig. 3.4. Tube current (top) and resulting image noise (bottom) for a scan from neck to shoulder. Red: constant tube current (conventional scan). Blue: tube current modulation. Green: tube current control and modulation (AEC). The thick lines in the top image are the local averages of the modulated graphs and illustrate the dose savings as a function of the z-position. The red graph is hidden by the blue graph in the bottom image

Fig. 3.5. This phantom experiment demonstrates that AEC

achieves constant image noise for all z-positions and that

noise structure can be signifi cantly reduced

Even more potential for dose reduction lies in the combination of AEC and MAF (Kachelriess et al.

2002) . Both methods are complementary: AEC glob- ally infl uences complete projections and MAF is able to locally adjust the noise statistics of single detec- tor readouts. The synergetic effects are illustrated in Fig. 3.6. The AEC reduces the tube current at the lung level where the attenuation is quite low. The increased noise is acceptable. In the shoulder, however, the tube current is globally increased by AEC. Here, MAF fur- ther reduces noise in the most central lateral rays.

An interesting specialization to the tube current modulation and tube current control technique has been recently introduced to reduce dose of retrospec- tively gated cardiac CT exams (Kachelriess and Kalender 1998) : The so-called ECG pulsing tech- nique reduces the tube current for those ECG phases that are unlikely to contribute to the reconstructed images (Jakobs et al. 2002).

3.5 Dose Reduction by User Education

When installed at a location, the CT system is usually adjusted to standard factory settings to give good image

quality. This does not automatically mean the lowest possible radiation dose. Learning how to customize and optimize the scan parameters (which constitute the scan protocols) with regard to dose and image quality is therefore a promising strategy to reduce dose.

However, customizing scan protocols to optimally meet all criteria is a diffi cult task. This is due to the complex dependence of image noise on the patient cross section (including size, weight, etc.). Finding the optimal mAs value, for example, is an iterative process that requires profound experience and, in principle, the possibility to repeat scans with different param- eters until the preferred image quality is achieved.

Dose considerations, however, forbid to repeat scans on a given subject. Phantom or cadaver scans may help, but the results are often too far from reality.

Recently, industry has started to provide tools that retrospectively simulate arbitrary scan parameters (Leidecker et al. 2002). Basically, these tools consist of simulating arbitrary tube currents or, equivalently, mAs products. The resolution can be varied retro- spectively by selecting different reconstruction ker- nels and different effective slice thicknesses.

Two classes of simulation methods are available.

One approach adds random noise to reconstructed CT images. The standard deviation of the random values is a function of the voxel grey value, the source

Fig. 3.6. The combination of AEC with MAF currently is the optimum dose-reduction strategy: due to its tube current modula-

tion and control, AEC increases noise in regions of low attenuation (arrowheads) and decreases noise and noise correlation in

regions of high attenuation and high eccentricity. In contrast to tube current modulation techniques that infl uence the photon

statistics of complete projections, MAF can modify single-detector element readings and further reduce the noise in the shoulder

and pelvis region (arrows). The difference in image indicates that no structure has been smoothed by the procedure

mAs and the desired destination mAs; however, this simple method does not correctly consider the X- ray physics which demand to add noise to the X-ray photon attenuation instead of adding noise to image pixels. Consequently, it cannot reproduce noise struc- ture and artefacts characteristic to CT images.

The second class of algorithms correctly simulates the physics behind CT imaging. To do so, it requires CT raw data and the complete image reconstruction pipe- line. Noise is added to the projection values to account for the desired destination mAs product. Figure 3.7 shows respective results using a raw-data-based tube- current simulator. As expected, image noise increases with decreased tube current and the noise structure generated is physically correct.

3.6 Tube Voltage and Filtration

As we have seen, there is much potential in optimiz- ing the dose usage by optimizing the tube current (or, equivalently, the effective mAs value). What remains to be considered are the effects of the tube voltage U and of the prefi ltration on image quality and on patient dose. It is well known that a high tube voltage yields a high-energy X-ray spec- trum and that the X-ray attenuation coeffi cients µ decrease with increased energy. In other words:

fewer photons are absorbed in the patient. The dose itself scales with both the number and the mean energy of absorbed photons. The number of pho- tons available is proportional to the tube current but is a complicated function of U.

Although these relationships are already complex enough, we have not yet included image quality into these considerations. Since the attenuation coeffi - cients vary with the tube voltage, the reconstructed grey values will also do so ( µ is a function of U);

therefore, it does not suffi ce to regard image noise σ alone as the indicator of image quality. Equation (1)

must instead be extended to put the noise σ and the signal µ into relation:

Here, D is the object diameter and the numerator accounts for the total object attenuation. Low values of ( σ/µ) are desired. Its inverse µ/σ is commonly known as the signal-to-noise ratio (whenever signal differences ∆µ are of interest the SNR is given as

∆µ/σ). Image quality increases with increasing SNR.

The complex dependence between the SNR and the tube voltage cannot be handled analytically, however.

Monte Carlo simulations or phantom experiments are required to analyse the effect of tube voltage on image quality and dose.

For example, a recent study evaluated iodine contrast as a function of tube voltage to improve the scan protocol design for paediatric CT (Schaller et al. 2001). Iodine contrast ∆µ = µ I – µ H

O and image noise were measured in a thorax phantom equipped with an iodine insert for 80, 120 and 140 kV. Dose was estimated using the CT dose index (CTDI) measured in a 16 cm CTDI phantom. The CTDI is defi ned as the integral over the axial dose profi le normalized to the nominal collimation:

The study resulted in the following values (Table 3.2):

image quality, given by the SNR, must be normalized with the dose value: Quality=SNR ² /Dose serves as an adequate image quality measure. The Table 3.2 clearly shows the strong dependency of image quality on the tube voltage: the quality of the 80 kV scan differs from that of the 140 kV scan by a factor 2.9. This means that the SNR at 80 kV is 70% better than at 140 kV when the same dose is applied. Alternatively, one can obtain the same image quality at 80 kV with only 34% of the dose.

Fig. 3.7. The dose tutor helps to decrease the tube current from its original level to the desired low-dose protocol (left: original 160 mAs, right: simulated 40 mAs). Noise structure is correctly reproduced. The user can experience with realistic examples which tube current is acceptable

(σ/µ) ² ∝

µ ² (I . t) _eff . S _eff . ∆r ³ e ^µD

CTDI _100mm = dz D (z) M . S

1

50mm

–50mm冕

These results, however, must be taken with care and cannot be applied to the general case. They instead emphasize the potential that lies in varying the X-ray spectrum. Note that all parameters have been correctly taken into account: contrast; noise;

spatial resolution; and dose. Here, comparisons were made at equal spatial resolution.

3.7 Penumbra Effects

In addition to the dose reduction strategies that have been mentioned, it is of great importance to know about fundamental differences between single-slice and multi-slice scanners. Theoretically, there should be no difference in dose usage if comparisons are performed at equal spatial resolution and at equal image noise. Practically, however, penumbra effects in the axial-dose profi le must be taken into account.

Due to the fi nite size of the X-ray tube’s focal spot, the desired rectangular distribution of primary X-rays in the z-direction cannot be achieved. In reality, the dose profi le has to be approximated by a trapezoidal function. The shape of this function is a function of scanner-dependent geometrical param- eters: the size F of the focal spot; the distance D of the pre-patient collimation to the focal spot; and the distance R _F of the focal spot to the centre of rotation (isocentre). The relations are illustrated in Fig. 3.8.

The profi le’s triangular tails are called penumbra, its constant rectangular region is called umbra.

To make full use of the exposure to the patient it is of interest to use not only the constant umbra region, but to include also the penumbras into the measure- ment. This can be easily achieved with single-slice scanners by opening the post-patient collimation until the penumbra is included. For multi-slice scanners (M >1), however, it is not possible to include the pen- umbra since the outermost slices (penumbra region) would record a different intensity than the innermost slices (umbra region) and image artefacts occur for this case; therefore, the penumbra region has to be cut

off by the post-patient collimation for scanners with two or more slices (Fig. 3.9). A dose increase relative to single-slice scanners is to be expected.

A simple geometric consideration shows that the width W of the penumbra at the isocentre is related to the size F of the focal spot as

Inserting values typical for modern MSCT scan- ners (e.g. R _F =600 mm, D=200 mm, F=0.5 mm) yields a penumbra size of W=1 mm. To estimate the “geo- metrical” dose increase of multi-slice scanners as compared with single-slice scanners one must regard

Table 3.2. Value results of the study

Protocol Dose Iodine Image Quality

(mGy) contrast noise (mGy)

(HU) (HU)

80 kV, 120 mAs 4.9 528 15.0 252 120 kV, 37 mAs 4.5 336 14.4 121 140 kV, 26 mAs 4.4 284 14.5 086

Fig. 3.8. Finite-sized focal spots yield trapezoidal primary dose profi les, whereas rectangular profi les are desired. The penum- bra size W is constant, whereas the size of the umbra varies with the collimation

Fig. 3.9. Differences in the post-patient collimation yield a slight dose increase with multi-slice scanners due to the exclu- sion of the penumbra region. This increase becomes negligible with wider cone angles and a higher number of slices

W = F .

D

R _F – D

the ratio of the total dose profi le area (umbra plus penumbra) to the umbra area (umbra only). This ratio is given by

Consequently, the dose increase itself is given by the term W/M·S which decreases with an increasing number M of slices and with increasing slice thickness S. Table 3.3 gives examples for 4- and 16-slice scanners with typical collimations assuming W=1 mm.

Apparently, the dose increase compared with single-slice CT becomes smaller for larger collima- tions and will become negligible with the introduc- tion of scanners with even larger cone angles (e.g. 32 slices or more).

3.8 Dose Information

Individual dose assessment will play a signifi cant role in the future. For example, the European Union Direc- tive (European Communities 1997) demands the spec- ifi cation of the individual patient dose for every CT examination. To directly assess scanner- and organ- specifi c dose values, dose calculators have become 1+ .

M . S W

Table 3.3. Examples for 4- and 16-slice scanners with typical collimations assuming W=1 mm

Collimation (mm) Geometrical dose increase (%)

04 × ^{1 25}

04 × ^{2.5 10}

16 × ^{0.75 08}

16 × ^{1.5 04}

available. Based on Monte Carlo dose calculations (Schmidt and Kalender 2002), they allow to com- pute organ-dose and effective-dose values as a func- tion of the scanner type and scan protocol (Fig. 3.10).

The effective dose is regarded as the best indicator of stochastic risk such as induction of malignancy. It expresses the sum dose value over all organs weighted by their radiation sensitivities. The weighting factors are recommended by the International Commission on Radiation Protection (ICRP 1990).

Dose calculators should also be used to answer jus- tifi ed questions about dose and risk values. Here it is of advantage to know the annual effective dose caused by natural internal and external radiation sources (e.g. 2.1 mSv in Germany, between 2.0 and 7.0 mSv in Europe, and 3.0 mSv in the United States; NCRP 1988;

BMU 2000). When reporting to a patient, effective dose values should be put into relation in that context. State- ments such as “the exposure of this examination cor- responds to two years of natural background radiation

Fig. 3.10. A typical dose

calculator computes

organ dose values and

the weighted effective

dose as a function of the

anatomical region and

the scan parameters. It

can be integrated into

the scanner environment

(Kalender et al. 1999a)

in the United States” are more informative than stating that the effective dose is 6.0 mSv.

3.9 Discussion

The demand for higher spatial resolution and lower image noise is limited by the increasing patient dose values according to Eq (1). Nevertheless, we have seen a number of strategies that help to compensate for the dose increase. This occurs either by making more effi cient use of the available dose or by adequate training programs.

On the one hand, manufacturers’ efforts will fur- ther improve dose utilization. Automatic exposure control, dose modulation, and intelligent adaptive fi l- tering algorithms bear the highest potential for dose reduction. The design of optimized scan protocols adds to these improvements.

On the other hand, there will always be enough degrees of freedom available to the radiologist to signifi cantly infl uence patient dose (in the positive or in the negative sense). The exposure to the patient should be adjusted to obtain the required diagnostic

information and not to get the best image quality possible. Training tools to learn about the infl uence of scan parameters help to iteratively seek for the pre- ferred image quality. Figure 3.11 shows a fully quali- fi ed image reconstruction software (VAMP GmbH, Möhrendorf, Germany). It optionally allows to arti- fi cially increase the mAs product from the original level to any value below by adding noise to the raw data and performing an image reconstruction. One may further start a multidimensional adaptive fi lter- ing algorithm and study the combined effects.

The decisive step in the near future will be the introduction of automatic exposure control which allows to directly specify the image quality (e.g. in terms of s). The mAs parameter may then vanish from the user interface to be replaced by one or more “image quality” parameters. Alternatives are the introduction of an “equivalent mAs value” corre- sponding to some normalized patient or the develop- ment of sophisticated user interfaces that graphically display reference images as a function of the selected parameters. These would illustrate the expected image quality before the scan is started.

Combining these strategies will help to limit the patient dose to reasonable values even if CT further goes to higher spatial resolution.

Fig. 3.11. The software-

based dose-reduction

algorithm MAF (left

group of the open tab)

and dose-training tools

that retrospectively allow

for a virtual decrease of

patient dose (right group

of the open tab) are read-

ily available in medical

applications such as the

VAMP Syngo Explorer

workstation

References

Baum U, Lell M, Kachelriess M, Greess H, Kalender WA, Bautz WA (2000) Rawdata-based 3D adaptive fi ltering for CT scans of the cervicothoracic region: Clinical evaluation.

Radiology 217:413

BMU (2000) (Bundesministerium für Umwelt, Naturschutz und Reaktorsicherheit) Umweltradioaktivität und Strahlenbelastung im Jahr 2000. Unterrichtung durch die Bundesregierung

Brenner DJ, Elliston CD, Hall EJ, Berdon WE (2001) Estimated risks of radiation induced fatal cancer from pediatric CT.

Am J Roentgenol 176:289–296

Eklundh JO, Rosenfeld A (1981) Imaging smoothing based on neighbor linking. IEEE Trans Pattern Anal Mach Intell 3:

679–683

European Communities (1997) Council Directive 97/43/

Euratom of 30 June 1997 on health protection of individu- als against the dangers of ionising radiation in relation to medical exposure, and repealing Directive 84/466/Euratom.

Offi cial Journal of the European Communities no. L 180/22 Gies M, Kalender WA, Wolf H, Suess C, Madsen MT (1999)

Dose reduction in CT by anatomically adapted tube current modulation. I. Simulation studies. Med Phys 26:

2235–2247

Greess H, Wolf H, Baum U, Lell M, Pirkl M, Kalender WA, Bautz WA (2000) Dose reduction in computed tomog- raphy by attenuation-based on-line modulation of tube current: evaluation of six anatomical regions. Eur Radiol 10:391–394

Harpen MD (1999) A simple theorem relating noise and patient dose in computed tomography. Med Phys 26:2231–2234 Hsieh J (1994) Generalized adaptive median fi lter and their

application in computed tomography. SPIE Proc 2298:

662–672

Hsieh J (1998) Adaptive streak artifact reduction in computed tomography resulting from excessive X-ray photon noise.

Med Phys 25:2139–2147

ICRP (1990) (International Commission on Radiological Protection) Publication 60. Recommendations of the International Commission on Radiological Protection.

Pergamon, Oxford

Jakobs T, Becker C, Ohnesorge B, Flohr T, Schoepf UJ, Reiser M (2002) Multislice helical CT of the heart with retrospective ECG-gating: reduction of radiation exposure by ECG-con- trolled tube current modulation. Eur Radiol 12:1081–1086 Kachelriess M, Kalender WA (1998) Electrocardiogram-cor- related image reconstruction from subsecond spiral CT scans of the heart. Med Phys 25:2417–2431

Kachelriess M, Kalender WA (2000) Computertomograph mit reduzierter Dosisbelastung bzw. reduziertem Bildpunk- trauschen. Deutsches Patent- und Markenamt. Germany, Patent Specifi cation DE 198 53 143

Kachelriess M, Leidecker C, Kalender WA (2001a) Image qual- ity-oriented automatic exposure control (iqAEC) for spiral CT. Radiology 221:366

Kachelriess M, Watzke O, Kalender WA (2001b) Generalized multi-dimensional adaptive fi ltering for conventional and spiral single-slice, multi-slice, and cone-beam CT. Med Phys 28:475–490

Kachelriess M, Leidecker C, Kalender WA (2002) Patient dose reduction by combining automatic exposure control (AEC) with multidimensional adaptive fi ltering (MAF) for spiral CT. Eur Radiol 12 (Suppl 1):196

Kalender WA (2001) Computed tomography. Wiley, New York Kalender WA, Schmidt B, Zankl M, Schmidt M (1999a) A PC

program for estimating organ dose and effective dose values in computed tomography. Eur Radiol 9:555–562 Kalender WA, Wolf H, Suess C (1999b) Dose reduction in CT

by anatomically adapted tube current modulation. II. Phan- tom measurements. Med Phys 26:2248–2253

Kalender WA, Wolf H, Suess C, Gies M, Greess H, Bautz WA (1999c) Dose reduction in CT by on-line tube current con- trol: principles and validation on phantoms and cadavers.

Eur Radiol 9:323–328

Keselbrener L, Shimoni Y, Akselrod S (1992) Nonlinear fi lters applied on computerized axial tomography. Theory and phantom images. Med Phys 19:1057–1064

Kopka L, Funke M, Breiter N, Hermann K-P, Vosshenrich R, Grabbe E (1995) Anatomisch adaptierte Variation des Röhrenstroms bei der CT. Untersuchungen zur Strahlen- dosisreduktion und Bildqualität. Fortschr Röntgenstr 163:

383–387

Lauro KL, Heuscher DJ, Kesavan H (1990) Bandwidth fi ltering of CT scans of the spine. Radiology 177:307

Leidecker C, Fuchs T, Kachelriess M, Schaller S, Kalender WA (2002) Comparison of different methods for adding virtual noise to measured rawdata in order to estimate the dose reduction potential for clinical protocols in CT. Radiology 225:592

NCRP (1988) (National Counsel on Radiation Protection and Measurement) Exposure of the population in the United States and Canada from natural background radiation.

Report 94

Schaller S, Niethammer MU, Chen X, Klotz E, Wildberger TG, Flohr T (2001) Comparison of signal-to-noise and dose values at different tube voltages for protocol optimization in pediatric. Radiology 221:366

Schmidt B, Kalender WA (2001) Calculation of dose reduction by tube current modulation in pediatric CT. Radiology 221:365

Schmidt B, Kalender WA (2002) A fast voxel-based Monte

Carlo method for scanner- and patient-specifi c dose cal-

culations in computed tomography. Phys Med 18:43–53