Dose optimization on a GE Discovery™ PET/CT 600

Amália Nogueira (amalia.nogueira@sapo.pt), Frederik A A de Jonge, Teresa C Ferreira

Unidade de Medicina Nuclear, HPP Lusiadas

         

The activity in the Department started in 2010. Our PET-CT system – GE Discovery ™PET/CT 600 Positron Emission Tomography Scanner - a fully integrated 3D PET scanner combined with a Bright Speed 16- slice CT scanner. This CT scanner is: -  multi-detector array helical CT (MDCT) with 16 slices in the longitudinal direction, (along the length of the patient lying on the table) per rotation of the gantry and a third generation CT in which the arc of detectors and the x-ray tube rotate together allowing -  rapid scanning (increased scanning speed due to a shorter rotation time and a wider beam having the ability to cover the entire scan volume within a single breath-hold reducing motion artifacts) -  wider scan field (scanning thin slices in one single dataset which can simultaneously be used for images with either high or standard longitudinal resolution improving the spatial resolution in the longitudinal direction without the drawback of extended scan times).  

According to GE Manual this equipment has special CT aspects: Ø Collimator - two independently controlled tungsten cams - the rotation of the cams provides continuous variable slice thickness and Z-axis position. Ø  The Scan Geometry -  gantry aperture - 70 cm. -  focal spot to Iisocenter distance – 54 cm, -  focal spot to detector distance – 95 cm. Ø  Matrix Detector -  24 detector cells in the Z direction; -  outside, four rows located on each side of the detector are 1.25 mm in the Z direction -  center, 16 rows are 0.625 mm in the Z direction. -  up to sixteen signals are gathered per gantry rotation - each signal can be collected from an individual detector row or a combination of detectors (the 4 channels at the beginning and end are not used due to physics of multi-slice scanning and helical view weighting algorithms). - 16 signals can be taken from 24 detector cells (or 16 slices per rotation of the gantry).

Other characteristics: - automatic exposure control (AEC) through automatic tube current modulation (ATCM) reducing the dose to the patient by altering the mA:

-  longitudinal (z) mA modulation (AutomA) -  angular or xy modulation (SmartmA).

The mA modulation is determined from the attenuation and shape of the scout scan projections of the patient and it is imposed by the selection of the noise index value (NI) chosen to the reconstructed image. This NI, an operator-selected variable, is a parameter indicative of the x-ray quantum noise contained in the scan data used to reconstruct the image and it is an important factor of the quality of a CT image. This NI will approximately equal the standard deviation of pixel values in the central region of the image when a uniform phantom* (with the patient's attenuation characteristics) is scanned and reconstructed using the standard reconstruction algorithm. * A 30 cm diameter water phantom or a 35 cm diameter low density polyethylene phantom.

Theoretically, the X ray noise (quantified as the standard deviation of the CT number in Hounsfield units) is inversely related to the square root of the number of photons and the number of photons is proportional to the slice thickness, slice acquisition time and mA. The most significant factor influencing the quantum noise in the scan data is the x-ray attenuation of the patient section being scanned. X-ray attenuation is related to the size and tissue composition of the patient section. Longitudinal (z) mA modulation (AutomA) varies the mA along the z axis to produce relatively uniform noise levels across to the various regions of anatomy. The goal of AutomA is to make all images contain similar x-ray quantum noise independently of patient size and anatomy. From this we can understand how important it is to center well the patient on the scan field of view (SFOV) because the bowtie filter projects maximum x-ray intensity at isocenter since this is the region of maximum attenuation if the patient is centered.

The images are acquired with thin sections but for axial viewing we use thicker slices (usually 5 mm*) and it is very important to understand that, in this system, the first prospective slice thickness is used to calculate AutomA. Generally we must set the NI for the thicker slice images because this choice conditions the value of mA used in the acquisition and consequently the value of the dose. This thicker series must be the our first choice (after AC series) when planning a protocol (even if we pretend 0,625 mm slices for 3D reformatting). (Those reconstructed images can be seen as a sandwich of data from multiple 0.625 mm detector rows and If we chose a value of NI for those reconstructed images, the signal-to-noise ratio (SNR) in any of those CT images will increase proportional to the square root of the number of detector rows composing the reconstructed image). *This value of 5 mm seems enough when we are working in PET/CT and the PET resolution is between 4 to 5 mm

Angular or xy modulation (SmartmA) addresses the variation in x-ray attenuation around the patient by varying the mA as the x-ray tube rotates around the patient. SmartmA (angular or xy modulation) adjusts the tube current to minimize X-rays over angles that have less importance in reducing the overall image noise content. Anatomically, there are highly asymmetrical structures such as the shoulders where x-rays are significantly less attenuated in antero-posterior (AP) direction than in the lateral direction. The operator chooses the initial mA value, and, through SmartmA, this mA value is modulated upward or downward, through the equalization of the photon flux to the detector, within a period of one gantry rotation. (So SmartmA only after AutomA.)

This CT system can acquire in three different modes: Axial, Cine and Helical. In helical or spiral mode the x-ray tube and Digital Acquisition System (DAS) exposes and rotates continuously through 360º while the patient is passing through the area of exposure at a preset rate of movement (pitch, ratio of table travel per rotation in millimeters divided by the beam collimation, is another important parameter in the dose to the patient; dose varies inversely to the pitch value). This information is then reconstructed into images at the chosen slice thickness and interval. The system has also the ability to manage dose, slice profile, and helical artifacts through two recon modes: Full and Plus mode Plus mode has up to 20% wider slice profile, uses additional views of data to reconstruct, so that the exposure time increases slightly but needs less 15-20% mA with equal noise (lower helical artifacts and is only available for 16 slice helical scans). Full mode provides a thinner slice but requires 10-15% more mA with equal noise. Both modes can be used prospectively and retrospectively.

The system also allows other additional reconstructions of acquired scan groups, using all or any portion of each group, changing prospectively some of the scan parameters: Display Field of View (DFOV), Algorithms, Recon Modes, and Window Width/ Window Level, Image Center, Start and End Locations for each group, image interval, Slice Thickness and Series Description. Recon Type Useful for Soft tissues with similar densities

Std routine exams like chest, abdomen and pelvis scans

Lung interstitial lung pathology

Detail hybrid tissue detail and bone edges

Bone high resolution exams and sharp bone detail

Edge small bone in the head and high resolution scans

Chest

mediastinum and lung detail

Since the beginning of PET/CT examinations in the Department we were concerned with the importance of exploring the capacities of the system and the need to look for the conditions that could decrease the dose applied to the the patients whithout loosing the diagnostic capability. In this work we had in mind the recommendations of some of the international authorities in radiation protection, namely: ICRP, Managing Patient Dose in Multi-Detector Computed Tomography (MDCT) (2006) IAEA Basic Safety Standards, Radiation Protection in Newer Medical Imaging Techniques: PET/CT (2008) IAEA TECDOC 1621, Dose Reduction in CT while Maintaining Diagnostic Confidence: A Feasibility/Demonstration Study (2009) IAEA Safety Report Series, Radiation Protection in Peadiatric Radiology (2012)    

In GE manual it is stated: “GE Healthcare strongly suggests reducing radiation dose to as low as reasonably achievable (ALARA) in all patients, especially pediatric and small patients, whenever it is determined that a PET/CT scan is necessary. PET/CT is an extremely valuable tool for diagnosing injury and disease, but its use is not without risk.” That concern was also important in terms of “collective dose” because CT and PET/CT are techniques with an increasing use. Understanding the CT acquisition and reconstruction parameters was very important to maintain image quality while imparting low doses to the patients. On the other hand, we have to be careful and not reduce the dose to a point that could affect image quality and decrease lesion detection.      

From the mentioned features of this CT system we can understand that the dose received by the patients due to CT examination is greatly dependent from the choice of the acquisition parameters. Automatic tube current modulation (ATCM) systems can greatly help to reduce patient dose if we are able to manage our scan protocols using measures related to image quality. When motion artifacts are separated, image noise and image contrast are the most important descriptors of image quality. Image noise, or quantum mottle, is most directly related to the radiation dose used for CT scanning. Image contrast is determined by a more complex relationship to the scan and reconstruction parameters - it is dependent on the x-ray tube potential (peak photon energy, kVp), but it is independent of the photon fluence (mAs). A decrease in kVp, decreasing the penetration of x-rays, can decrease radiation dose but increases image contrast, whereas an increase in kVp decreases image contrast.

In fact, image noise and image contrast have to be considered to adapt scanning parameters for managing radiation dose. Tube amperage (mA) is approximately proportional to radiation dose and usually this value can be adjusted for body mass and configuration. The dose varies approximately with the square of kVp. Tube voltage is less frequently adjusted for body mass than mA although GE manual advises to establish protocols utilizing 100 or 80 kVp (but accompanied by an increase of mA in order to keep the noise level of the image) in thin patients and children. Considering all this we decided to limit our work for managing the dose to the exploration of mA influence, besides the application of the general measures of “good practice”:

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§  Perform Only Necessary PET/CT Examinations §  Scan Only the Anatomical Region Indicated §  Properly Center All Patients in the Gantry We also kept a register of the values of the CTDivol and DLP values of all our studies (including the values of the FDG-F18 administered). Those values enable us to compare our results with other departments and even to validate them in terms of international reference levels

CTDi vol (mGy) N. examinations

2010 3.8 ± 0.45 158

2011 5.4 ± 0.71 258

2012 5.7 ± 1.77 754

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As a starting point of exploration of the CT scanner features we used a protocol optimised to the goal of our studies mainly directed to whole body studies of oncological patients. This protocol was designed with the help of a GE applications specialist taking in consideration the clinical requirements of the CT in our department – the attenuation correction of PET images and the better localization of the lesions (better morphological information). The exploration of the CT’scanner features, was performed taking as starting point this optimised protocol and having in mind the importance of going further in decreasing the doses of our examinations but whithout loosing the necessary quality of the image.

Scan Type

Thick speed

Interval (mm)

SFOV kV mA

Helical Image thickness

(3.75)

3.270 (Ped. Head; Head; Small) Large

(80;100; 140) 120

mA range 80-150

Full Table speed (13.75)

NI 17.0

0.5 sec Pitch (1.375:1)

Acquisition conditions used:

We knew the importance of the flux of photons produced on the x-ray tube for the quality of the image and how this quality is influenced by the quantum noise of the reconstructed image. Theoretically this flux of photons depends directly from the mA of the tube. (GE equipment instead of the “noise” uses another parameter related with it and called “noise index” (NI)). Knowing the importance of the NI value on the value of the dose imparted to the patient we wanted to understand better the relationship between NI and radiation dose when we varied the values of mA and different values of NI. Taking as starting point the protocol optimised by the applications specialist, the only changes introduced were the values of NI and the limits of mA. We always used the ACTM features.

Imagem  do  phantoma  

The CT examinations were performed using the CT body phantom (32-cm-diameter polymethyl methacrylate phantom)

Test identification

kV mA NI Series description

CTDivol mGy

Phantom

NEMA 12 120 80 - 150 12 AC- ST –Lung

5.43 CT

NEMA 17 120 80-150 17 Lung

4.04 CT

NEMA 18 120 50-100 18,85 AC- ST –Lung

2.70 CT

The values of CTDivol were recorded from the CT scanner console.

the NI imposed on the protocol is the NI of the first series after the serie called “AC series” (used for attenuation correction). This calculation was performed applying the mentioned GE definition: This calculation was performed applying the mentioned GE Reading the standard deviation of pixel values on a circular definition: region of interest (ROI), aproximately 1.4 mm2, centered on an uniform slice of the phantom

1,25”, those designations mean that the first series is taken from the contributions of 4 slices of 0.625 mm each and the other from the contribution of 2 slices of 0.625. (Those calculations for all the acquisitions performed are still under study)

(Those calculations for all the acquisitions performed are still under study)

On the next set of images from those 2 series we can see that the noise of the Lung series(4) is higher than the St series (3) the Lung series(4) is higher than the St series (3)

3

4

After those CT tests we performed also a study of PET Phantom using the CT acquisition conditions from NEMA 12 e NEMA 18. The next images show the results

We decided to explore our results a bit further and we made a lot of new CT acquisitions of images using different associations of mA, but keeping a constant difference of 50 mA, and NI values; always with the same protocol and the CT phantom

Test identification

mA NI

1,;3; 4; 5; 6 50-100; 40-90; 30-80; 20-70 17,0

11;13;14;15;16 50-100; 40-90; 30-80; 20-70 18,85

21; 23; 24; 25; 26 50-100; 40-90; 30-80; 20-70 20,09

31; 33; 34; 35; 36 50-100; 40-90; 30-80; 20-70 40,46

In terms of dose we got those results showing that we could really lower the dose. According to the limits chosen to mA we succeeded to lower the dose varying from 28 to 51 %.

In further explorations we tried to design those curves of NI versus mA when the values of NI where lower than the value adopted on the optimised protocol

Test identification

mA NI

00A;00B…;00H 80-130; 70-120; 60-110; 50-100; 40-90;30-80; 20-70;10-60 5,27

0A;0B;0C…..0H 80-130; 70-120; 60-110; 50-100; 40-90;30-80; 20-70;10-60

12,35

1A; 1B;1C;…;1H 80-130; 70-120; 60-110; 50-100; 40-90;30-80; 20-70;10-60

17,0

2A;2B;2C;…;2H 80-130; 70-120; 60-110; 50-100; 40-90;30-80; 20-70;10-60

18,85

3A;3B;3C;…;3H 80-130; 70-120; 60-110; 50-100; 40-90;30-80; 20-70;10-60

30,2

4A; 4B;4C;…; 4H 80-130; 70-120; 60-110; 50-100; 40-90;30-80; 20-70;10-60

40,46

And we got those results in terms of dose (CTDivol)

•  From those curves we can conclude that it is not worth to acquire with values of NI lower than 12,35 and we were able to understand why this value is called the “reference value”.

•  The decrease in dose for those different mA conditions between values of NI=12,35 to NI=17,00 varied in an approximately regular way from 26 to 0 % (23; 26; 24; 19; 15; 9; 3;0)

•  The decrease in dose for those different mA conditions between values of NI= 17 to NI=40 also varied regularly from 12 to 58 % (12; 14; 18; 27; 35; 45; 53; 58)

In terms.of images we have the next results…….

Comparing the high resolution images with lower NI and the extreme conditions in terms of mA values

NI = 17,00

80-150 mA 10-80 mA

Comparing the low contrast detectability of the images with lower NI and the extreme conditions in terms of mA values 80-150 mA 10-80 mA

NI = 17,00

Comparing the high resolution of the images with higher NI and the extreme conditions in terms of mA values

80-150 mA 10-80 mA

NI = 40,46

80-150 mA 10-80 mA

Comparing the low contrast detectability of the images with higher NI and the extreme conditions in terms of mA values

NI = 40,46

Later on we also tried to understand what happens when we change only the lower level of mA.

or the higher level…..

We have performed a lot of tests and we got some interesting conclusions: 1. We succeeded to better understand the system and to prove the possibility of lowering the doses without compromising the diagnostic 2. When using AutomA and SmartmA and varie the mA limits: ü We can increase the values of NI from 12 (reference value) to 17 gaining in terms of dose (CTDi vol) up to 26 %.

ü In terms of dose the variation of the higher limit of mA has a very small importance when NI varies from 17 to 40.

ü On the contrary the variation of the lower limit can give differences from 14 to 65%, when NI varies from 17 to 40, and the mA limit varies from 80 to 10.

Even with these preliminary results, we have a lot of points that we want to verify or clarify: •  Are those calculated values of NI on the acquired images influenced by the imposed values of mA limits?

•  What is the result of those low values of mA limits on the images of PET phantom*?

•  What is the limit in terms of contrast to decrease those low mA limits? (Partial volume effect can be a problem but perhaps we can compensate this problem. With the reduction in image thickness, the magnitude of partial volume averaging also decreases. Thus, the CT number - image brightness - associated with objects that occupy less than one voxel increases as the voxel size decreases. For objects with z-axis dimensions less than one image width, the contrast of the object improves with reduced slice thickness, whereas quantum noise increases with reduced slice thickness. If a narrow image thickness is used, the contrast-to-noise ratio and visibility of small details can improve despite increased noise )

•  We recorded always the dose values from the scanner but we intend to check those results with measurements with an ionization camera. *Those verifications on PET phantom are completely determinant but they are also more difficult because we need activity to perform them.

We also need to proceed our explorations on other important parameters that we kept constant on our tests: •  The influence of kVp in lowering the dose: variation in the tube potential causes a substantial change in CT dose as well as image noise and contrast; reduction in kVp can result in a substantial increase in the image noise and it can impair image quality if the patient is too large or if the tube current is not appropriately increased to compensate for the lower tube voltage. Thus, when implementing reduced kVp protocols, it is imperative that appropriate mAs values are determined as a function of patient size. For very large patients, relatively higher tube voltage is almost always needed to obtain diagnostically adequate studies. •  The interactions of the different reconstruction filters and protocols on the quality of the images and dose. (image post-processing filters have been designed to decrease image noise in scans acquired with reduced radiation dose). •  Decide the best approach to change mA or kVp according to the size of the patients

With all these results I hope to get my final goals: •  To be able to create different protocols adjusted to the conditions of the patients and imparting lower doses to the patients

•  To convince the Physicians in the Department that we can safely apply them without loosing the diagnostic quality of the images

•  To lower the DRL’s of the Department