SAHLGRENSKA ACADEMY

EVALUATION OF SPECT/CT WITH TWO

BED POSITIONS FOR 177LU TREATMENTS

Felicia Halleby

Essay/Thesis: 30 hp

Program and/or course: Medical Physicist Programme

Level: Second Cycle

Semester/year: Spring 2017

Supervisor: Jakob Himmelman, Tobias Rydén and Peter Bernhardt

Examiner: Magnus Båth

Abstract

Essay/Thesis: 30 hp

Program and/or course: Medical Physicist Programme

Level: Second Cycle

Semester/year: Spring 2017

Supervisor: Jakob Himmelman, Tobias Rydén and Peter Bernhardt

Examiner: Magnus Båth

Keyword: 177Lu, SPECT/CT, Multi-field-of-view SPECT, OSEM

Purpose: The aim of this study was to evaluate the impact on image quality when different

acquisition parameters and reconstruction methods are used. Another purpose was to

investigate the possibilities of reducing the SPECT (single photon emission computed

tomography) examination time for the patients treated with 177Lu.

Theory: In radionuclide therapy 177Lu is utilized for treatment of neuroendocrine tumours and

prostate cancer. 177Lu emits primarily beta particles but also a small amount of gamma

photons which makes it possible to visualize the uptake of 177Lu with a gamma

camera. Tomographic images (SPECT) covering 40 cm of the patient length, are

generally collected for quantification of the uptake of 177Lu in tumours and risk organs

at Sahlgrenska University Hospital. In many cases it would have been preferable if the

SPECT images covered a bigger part of the patient, but that would also increase the

examination time to an unacceptable length if the acquisition protocol is not

optimized.

Method: Phantom measurements were performed to evaluate the impact on image quality when

time per projection angle, the number of projection angles, energy window setting

(one or two energies) and reconstruction algorithm was changed. IRACRR (iterative

reconstruction with resolution recovery), IRACRR with TEW (triple energy window)

scatter correction and SARec (Sahlgrenska Academy Reconstruction code) were the

algorithms used for the reconstructions. Image quality parameters were calculated and

compared.

Result: The highest signal to background ratio was measured in the TEW (IRACSCRR) and

the SARec images reconstructed with one photon energy. The superior signal to noise

ratio was calculated from the TEW with one energy and IRACRR with two energies.

Signal to background ratio did not change with altered acquisition time or number of

projection angles. The noise increased when the acquisition time was reduced.

Abbreviations

BSA Bovine Serum Albumin

CPU Central Processing Unit

CT Computed Tomography

DEW Dual Energy Window

FOV Field of View

GPU Graphics Processing Unit

HU Hounsfield Units

IRACRR Iterative Reconstruction with Attenuation Correction and Resolution Recovery

IRACSCRR Iterative Reconstruction with Attenuation Correction, Scatter Correction and Resolution

Recovery

MC Monte Carlo

MEGP Medium Energy General Purpose

MIRD Committee on Medical Internal Radiation Dose

MLEM Maximum Likelihood Expectation Maximization

NET Neuroendocrine Tumour

OSEM Ordered Subset Expectation Maximization

PET Positron Emission Tomography

PSMA Prostate Specific Membrane Antigen

PVE Partial Volume Effect

SARec The Sahlgrenska Academy Reconstruction code

SBR Signal to Background Ratio

SNR Signal to Noise Ratio

SPECT Single Photon Emission Computed Tomography

TEW Triple Energy Window

VOI Volume of Interest

4

Table of content

Abbreviations .......................................................................................................................................... 3

Introduction ............................................................................................................................................. 5

177Lu treatments ................................................................................................................................... 5

177Lu imaging....................................................................................................................................... 5

Image reconstruction and corrections .................................................................................................. 6

Imaging of patients during lutetium treatments at Sahlgrenska University Hospital .......................... 8

Phantom measurements of 177Lu ......................................................................................................... 8

Aims ........................................................................................................................................................ 9

Materials and Methods .......................................................................................................................... 10

Adhesion measurements .................................................................................................................... 10

Phantom measurements ..................................................................................................................... 10

Image reconstruction and evaluation ................................................................................................. 12

Results ................................................................................................................................................... 14

Discussion ............................................................................................................................................. 19

Conclusions ........................................................................................................................................... 22

Acknowledgements ............................................................................................................................... 23

Reference list ......................................................................................................................................... 24

Appendix ............................................................................................................................................... 25

5

Introduction

In nuclear medicine, radioisotopes are used for both imaging and treatment of primarily cancer [1].

The biological and chemical properties of either the radioisotope itself or the radiopharmaceuticals are

utilized in nuclear medicine for targeting of the tumours. Depending on what type of tumour cells that

is supposed to be irradiated, the properties of the radiopharmacy are designed to give a higher uptake

in the tumour cells than in normal tissue. The radiopharmacy is generally injected intravenously and is

distributed in the body by the bloodstream. A tumour targeting technique makes it possible to treat

metastases - something that is difficult in for example external radiotherapy [2]. Utilization of

radioisotopes in imaging with positron emission tomography (PET) or gamma cameras gives

functional information that cannot be visualized with other imaging methods [1]. Depending on what

type of radiation that the radionuclides are emitting, they have different applications. Radionuclides

that emit alpha or negative beta particles when they decay are more suitable for treatments, while

annihilation or gamma photons are more desirable in imaging because of their ability to penetrate the

body. In therapy situations the aim is to deposit energy locally in the target and radionuclides with

emission of alpha particles or negative beta particles are therefore the optimal choice since they have a

short range in tissue and are more likeable to force the tumour cells into any form of cell death [2].

177Lu treatments 177Lu is a lutetium isotope with a half-life of 6.7 days, it emits primarily beta particles and is used in

radionuclide therapy. In radiotherapy of neuroendocrine tumours (NETs) it is common to use

radiopharmaceuticals that are labelled with 177Lu. NETs derive from neuroendocrine tissue and are

most frequently found in the gastrointestinal tract and the bronchopulmonary system [3]. The cells in

NETs are generally overexpressing somatostatin receptors and that makes it possible to treat

these tumours with radiolabelled somatostatin analogues. One radionuclide therapy of NETs is

performed by utilizing the chelator DOTA to label lutetium with the somatostatin analogue (Tyr3)-

octreotate (TATE), building the complex 177Lu-DOTATATE. (Tyr3)-octreotate binds to the

somatostatin receptors on the cell surface and the 177Lu-DOTATATE-complex is internalized into the

cell through endocytosis [4, 5].

Another radionuclide therapy where 177Lu can be used is in combination with the PSMA-617, an agent

that attaches to PSMA (Prostate Specific Membrane Antigen). As the name indicates, PSMA is an

antigen in the membrane of prostate cells. PSMA is overexpressed in prostate cancer cells and are

rarely found in the normal cells outside the prostate which makes PSMA a distinguished target for

imaging and treatment of prostate cancer metastases [6]. For patients with castration-resistant prostate

cancer, therapy with 177Lu-PSMA-617 would most probably increase the length and quality of life [6].

Even though tumour cells are the target in radioisotope therapy some of the radioactivity will be

distributed in other organs and tissues. The absorbed doses to the bone marrow and the kidneys are

often of concern in radionuclide therapy since the radiopharmaceuticals are distributed in the blood

and excreted through the kidneys. In 177Lu-PSMA-617 therapy, the salivary glands also are organs at

risk since they have high activity uptake [4, 6]. For areas with augmented absorbed doses, the

dosimetry is of immense importance to prevent adverse effects.

177Lu imaging

An advantage of using 177Lu in therapy prior of many other radionuclides is that there is a small

amount of gamma emissions in the 177Lu decay. Primarily two gamma energies are emitted: 113 keV

6

and 208 keV with decay incidences of 6.4 % and 11.0 % per decay [7]. The gamma decays make it

possible to visualize the uptake of lutetium with a gamma camera. During treatment with 177Lu-

DOTATATE, both planar and tomographic acquisitions are made where the latter are managed with

Single Photon Emission Computed Tomography (SPECT), combined with a CT. The SPECT/CT

images are required to quantify the uptake in three dimensions. Due to the unique biokinetics of all

individuals a quantification tool like SPECT/CT is essential for dosimetry and diagnostic precision

[8].

Image reconstruction and corrections

Iterative reconstruction algorithms are recommended for quantitative images [5]. One effective

iterative method is ordered subset expectation maximum (OSEM) [9, 10]. OSEM is based on the

maximum likelihood expectation maximum (MLEM) method [2] with the advantage that the

reconstruction is accelerated by calculating a number of projections simultaneously in subsets. The

basic idea of iterative reconstructions is that an estimate of the projection image is forward projected

into its sinogram which is compared to the measured sinogram. Henceforth, the estimate is updated

with regard to the sinogram differences and the procedure is repeated until the estimated sinogram has

converged enough towards the measured sinogram [2]. The MLEM updates are mathematically

described in equation 1.

𝑓�̅�

𝑘+1 =𝑓�̅�

𝑘

∑ 𝑎𝑖𝑗𝑛𝑖=1

∑𝑔𝑖

∑ 𝑎𝑖𝑗𝑓̅𝑗𝑘𝑚

𝑗=1

𝑎𝑖𝑗.

𝑛

𝑖=1

1

Where 𝑓�̅�𝑘+1 is the updated estimate, 𝑓�̅�

𝑘 is the old estimate, 𝑎𝑖𝑗 is the probability of detecting a photon

from pixel j in bin i and 𝑔𝑖 is the measured sinogram.

In gamma camera imaging, there are some physical and geometrical factors that affect the measured

activity, which gives a false distribution evaluation if they are not corrected for. Partial volume effect

(PVE) occurs when activity “spills over” from or to the structure of interest, and the spill over is

generating a false activity distribution in the image. The PVE gives a low contrast between areas of

high and low activities and it is an outcome of the resolution limitations of the detecting system. There

are several ways to correct for PVE but one way is to implement the collimator-detector response

(CDR) function into the reconstruction [8]. The CDR function describes the blur caused by the

intrinsic properties and the collimator properties of the detector system, factors that affect the blur are

for example the length and diameter of the collimator holes.

Numerous gamma photons will interact with tissue in the patient or with the camera material during

imaging. Some photons will be attenuated in the patient body and the measured counts in the camera

detectors will consequently decrease, which creates image artefacts. If a CT is connected to the

gamma camera system an attenuation correction can be done by converting the Hounsfield units (HU)

to attenuation coefficients for the current gamma energy. Motion of the patient between emission and

transmission in a SPECT/CT also leads to artefacts but this can be corrected for manually with post-

registration of the motion [11]. The detector system presumes that the detected photons origin from a

location perpendicular to the detection position. If a photon first is scattered and changes its direction

before it reaches the detector, it gives a false measured activity distribution since the detection does

7

not occur perpendicular to the source of the decay [8]. Scattering windows can be applied to correct

for the scattering effects, the basic idea of this method is to use multiple acquisition windows to

estimate the scatter amount. Several scattering window techniques have been developed, but two of

the most frequently used are dual energy window (DEW) and triple energy window (TEW). In the

DEW method one extra window below the main window is set and in the TEW method two narrow

windows are set on both sides of the main window (see Figure 1). However, scatter correction with

DEW or TEW in 177Lu imaging is not an optimal choice since it is a method that is considered to

amplify the noise [12].

Figure 1: Visualization of the TEW technique. The blue solid line represents a photo peak with surrounding scattered

photons. The greens lines separate the three acquisition windows. The main window is placed centred over the photo peak

and two scattering windows are placed below (Low) and over (High) the main window. The solid red line under the photo

peak represents the scattered photons in the peak and the dashed blue line under the peak represents the amount of counts

that are removed when scatter corrected.

Monte Carlo (MC) is a powerful tool when it comes to scatter correction through simulation of the

particle transport from decay to detection [13]. The limitation with MC is that the calculations are

demanding and the reconstructions often take too much time for clinical use. With accelerated

calculations though, it is possible to improve the image quality (in the clinic) using MC corrections

implemented into an OSEM reconstruction of the image [14]. The Sahlgrenska Academy

reconstruction code (SARec) [14] is a MC based code developed for reconstruction of SPECT/CT

images using OSEM iterations. In SARec, a collimator-detector model is implemented to correct for

the PVE.

A considerable source to the long computation times in MC reconstruction is the amount of forward

projections required for the reconstruction. If the code is written for the central processing unit (CPU)

the calculation velocity is reduced due to the limitations of parallelization of the code. In SARec the

graphic processing units (GPUs) is utilized to parallelize the code in a more effective way than what is

possible in a CPU code, the parallelization gives a major reduction of the reconstruction time [14]. The

change in reconstruction time depends on several factors, like the properties of the GPU and the CPU

that are used, but a rough estimation is that the reconstruction time could be reduced by a factor of 100

if SARec is used compared to a CPU based MC-algorithm. The previous simulation performance of

8

SARec was 3200 million photons/s, but the GPUs have been exchanged to ones with higher capacity

since that measurement [14].

Imaging of patients during lutetium treatments at Sahlgrenska University Hospital

A SPECT/CT examination for a 177Lu-therapy patient at Sahlgrenska University Hospital today takes

approximately 30 minutes. The acquisition protocol for the lutetium examination includes 120

projection angles á 30 s, a 128×128 matrix, an acquisition window ±10 % around the 208 keV energy

peak and a Field of View-length (FOV) of 40 cm. The images are reconstructed with IRACRR

(iterative reconstruction with attenuation correction and resolution recovery) in Xeleris 3.1108 (GE

Healthcare) with 10 subsets and 2 iterations. In the IRACRR algorithm, iterations with OSEM are

performed and resolution recovery is used to correct for the resolution loss due to the CDR. Because

of the low incidences of gamma decays the collection time becomes long. There is a high amount of

scattered photons with energies around the lower energy peak, which results in that clinics generally

choose to only collect the photons of 208 keV [5]. Depending on the locations of tumours and risk

organs a FOV of 40 cm is not always enough to cover all the interesting structures. One way to

increase the FOV is to perform two SPECT-acquisitions with two different bed positions. A multiple

bed SPECT results in an improved activity quantification compared to planar scans [15]. The major

drawback with two bed SPECT is that it also doubles the examination time if there are no optimization

of the procedure. Factors that affect the examination time are the number of projection angles and the

acquisition time per angle. With more statistics by sampling two energy peaks or with an improved

reconstruction method, the SPECT examination time might be reduced.

Phantom measurements of 177Lu

In previous phantom measurements performed by PhONSA (The medical Physics, Oncology &

Nuclear medicine research group at Sahlgrenska Academy), BSA (Bovine Serum Albumin) has been

added in the phantoms to prevent adhesion of the activity on the phantom walls. In the reconstructed

images from these phantom measurements some inhomogeneities have been identified in volumes

with homogeneous distributed activity. One hypothesis is that the BSA is causing these

inhomogeneities and therefore the wish is to perform phantom measurements without BSA but with no

increased adhesion of lutetium on the phantom walls.

9

Aims

• To investigate the influence of the window settings, acquisition times and the number of

collection angles on image quality.

• To reduce the SPECT examination time by finding the optimal acquisition parameter values

and reconstruction method.

• To generate an examination protocol for clinical use at the Sahlgrenska University Hospital

with two bed positions for studies with 177Lu-DOTATATE and 177Lu-PSMA.

10

Materials and Methods

Adhesion measurements

A solution containing 0.5 L deionized water and 1 g Bovine Serum Albumin (BSA) was prepared,

creating a BSA concentration of 2 g/L. The mixing of the solution was done during 10 minutes in a

magnetic stirrer with no heat on. BSA concentrations of 0.2 g/L and 0.02 g/L were mixed through

dilution of some of the original solution. A sampling tube was filled with deionized water and 177Lu-

DOTATATE was added, forming an activity concentration of 0.48 MBq/mL. The activity was drawn

up with a syringe, the activity in the syringe was measured in a dose calibrator ion chamber (CRC-15,

Capintec, IA, USA) before and after it was emptied in the sampling tube and the difference was used

to calculate the concentration. A total of 8 cuvettes made of PMMA (same material as the phantoms

used later in this study) was filled with 0.5 mL BSA-solution or just deionized water, two cuvettes was

filled with each of the four different BSA concentrations (2 g/L, 0.2 g/L, 0.02 g/L and 0 g/L). 0.5 mL 177Lu activity solution was added in all the cuvettes giving a total activity concentration of 0.24

MBq/mL and BSA concentrations of 1 g/L, 0.1 g/L, 0.01 g/L and 0 g/L respectively. The counts per

minute (cpm) in the filled cuvettes was measured with a gamma counter (Wallac 1480 Wizard® 3”

NaI (Tl), Wallac Oy, Turku, Finland). Two hours after the filling of the cuvettes they was emptied and

rinsed with tap water before they were measured with the gamma counter for a second time. Each

cuvette was measured during 5 minutes at a time and the ratio between the measured counts before

and after the emptying was calculated. No decay correction was performed due to the long half-life of 177Lu.

Phantom measurements

Three different phantoms where used in this study; the Jaszczak SPECT phantom with hot spheres and

cold rods and secondly the Lung-Spine phantom with hot spheres, lung cavities filled with

Styrofoam® beads and a spine insert of Teflon®. Finally, the Triple Line phantom with three line

sources was used. All phantoms used in this study are shown in Figure 2, the heart insert visualized in

the lung phantom (centre) was not used. The cold rods were not used for image evaluation but for

radiation safety issues. The rods fill up some of the phantom volume and less activity must be handled

and added to the rest of the volume to gain the wanted activity concentration. The hot spheres where

simulating tumours of varied sizes and the lung phantom was used to emulate the heterogeneous

composition of densities in the upper torso. The lung cavities contained a mixture of Styrofoam®

beads and sodium azide solution to mimic lung tissue with a density of 0.3 g/cm3. The line sources

were used to gain a long object to image and for profile intensity measurements.

Figure 2: Jaszczak SPECT phantom (left), Lung-Spine phantom (centre) and Triple Line phantom (right).

11

The Jaszczak hot spheres had volumes of 0.5, 1, 2, 4, 8 and 16 mL respectively. Also, a hot sphere of

16 mL was placed between the lungs and another sphere of 4 mL was placed adjacent to the spine in

the lung phantom. All the spheres were filled with an activity concentration of 0.41 MBq/mL. The line

sources where filled with an activity concentration of 42 MBq/mL. The background volumes of the

Jaszczak and lung phantoms, i.e. the volume outside of the spheres, were filled with activity

concentrations of 0.046 MBq/mL and 0.047 MBq/mL respectively, this constitutes sphere-to-

background ratios (R) of 8.9 in the Jaszczak phantom and 8.7 in the lung phantom. The activity

concentrations were chosen to mimic the activities in tumours (spheres) and normal tissue

(background) in a patient at a lutetium SPECT examination. Sodium azide giving a concentration of

0.2 g/L was added to the activity volumes to avoid growth in the phantoms. Due to the results from the

adhesion measurements and after consultation with the manufacturer, BSA was not used in the

solution. Beakers were filled with deionized water and the sodium azide was added, a magnetic stirrer

with no heat on was used in 10 minutes to resolve the mixtures. To maintain the wanted activity

concentration for the various phantoms and spheres, a syringe was used to extract activity from a vial

containing 177Lu-DOTATATE. Henceforth, the activity was added to an external container with water

to mix the concentrations for the spheres and lines or directly into the phantoms that was prefilled with

mostly of the sodium azide solution. The activity of the syringe was measured with a dose calibrator

(VIK-202, Comecer, Netherlands) before and after the emptying, the difference was used to get a good

estimation of the activity added.

The phantoms were ordered in a stack on the patient bed of a SPECT/CT camera (Tandem Discovery

670 Pro, GE Healthcare, Waukesha, WI) simulating the patient length of interest. Figure 3 shows the

arrangement of the phantoms.

Figure 3: Phantom arrangement during the image acquisition.

SPECT/CT acquisitions based on the protocol for 177Lu-DOTATATE examinations at Sahlgrenska

University Hospital were performed but extended to two FOVs. Following parameters was studied,

changing one at a time: acquisition time per projection angle, the number of projection angles, the

number of main energy windows (one or two) and scatter correction with TEW. For every image a

range of 360°, a 128×128 matrix, a pixel size of 4.42 mm, a slice thickness of 4.42 mm and zoom 1

was applied. The projection angles were homogeneously distributed over the 360° range. When the

number of projection angles was reduced, the time per angle was increased to keep the total number of

12

detected photons constant. For the acquisition with time and angle parameters identical to the former

lutetium protocol at Sahlgrenska University Hospital, reconstructions with both one energy, two

energies and one energy with scatter correction (TEW) were performed, and for the other cases only

the information from the 208 keV peak were used. Performed SPECT acquisitions are showed in

Table 1. A MEGP (Medium Energy General Purpose) collimator was used for all acquisitions.

Guidelines from MIRD pamphlet no 26 [5] and parameter values from [5] were considered during the

parameter choices. All acquisitions were performed the same evening and no decay correction was

performed due to the half-life of 177Lu.

Table 1: SPECT acquisition parameters.

Number of angles

Time per angle (s)

Main window (keV)

Scattering window (keV)

Total acquisition time (min:s)

120 30 208 ±10% - 66:10

120 30 113 & 208 ±10% - 66:10

60 60 208 ±10% - 64:12

30 120 208 ±10% - 62:44

120 15 208 ±10% - 36:10

120 10 208 ±10% - 26:10

120 30 208 ±10% 182.6 ±2.2% 234.2 ±2.1%

66:10

Image reconstruction and evaluation

The images were reconstructed with OSEM in both Xeleris (3.1108) and SARec. The OSEM updates

were performed with 10 subsets and 4 iterations in the Xeleris reconstructions because that is the

parameters used in GE’s quantification software. For the SARec reconstructions 10 subsets and 5

iterations were used since the image convergence appears after approximately 5 iterations [14]. No

post filtration was applied for the original image quality measurements, but for a comparison of image

quality and for visualization of the results a Butterworth filter was applied. In the Xeleris images there

are an undefined smoothening of the image in the reconstruction but no additional filtration was

performed.

Intensity profiles were collected over the largest sphere (16 mL) in the Jaszczak phantom and over one

of the line sources for all reconstructions and acquisition approaches. Results are shown in the

Appendix.

Segmentation of the spheres was done with manually placed spherical volumes of interest (VOIs) with

diameters corresponding to the inner diameters of the spheres, (31, 25, 20, 16, 12 and 10 mm). 20

background VOIs with the same size as the hot spheres were positioned in the background volume.

The placement of the VOIs was visually evaluated based on the CT images that was matched with the

SPECT images. The background VOIs were placed in volumes where the activity and density

distributions were homogenous. For the lung phantom, a homogeneous volume could be found

beneath the lung cavities.

13

The normalised signal-to-background ratio (SBR) and the signal-to-noise ratio (SNR) between the

sphere and background volumes were calculated with equations 2-3.

𝑆𝐵𝑅 =

�̅�

�̅� ∙ 𝑅 ,

2

Where SBR is the normalised signal-to-background ratio, SNR is the signal-to-noise ratio, �̅� is the

mean voxel value in the sphere, �̅� is the mean voxel value in the 20 background VOIs, R is the true

activity ratio and �̅�𝑏 is the mean standard deviation between the background VOIs. The voxel values

for the VOIs are expressed in counts/mm3.

A two-sided paired t-test for the six spheres in the Jaszczak phantom was performed on the SBR and

SNR between the acquisition with 120 projection angles and 30 s per angle and the acquisitions with

changed acquisition approaches.

𝑆𝑁𝑅 =

|�̅� − �̅�|

�̅�𝑏.

3

14

Results

The measurements of adhesion on PMMA showed no correlation between BSA concentration and

adhered 177Lu-DOTATATE, the results are to be found in the Appendix.The image artefacts shown in

previous studies were not present in the reconstructions in this work.

Decreasing sphere size gave reduced SBR and SNR for both the Xeleris and SARec reconstructions in

this study. The comparison of SBR and SNR for one and dual energy peaks and scatter correction with

IRACSCRR (TEW) are presented in Figure 4 a-b. In Figure 4 c-d the SBR and SNR for filtered

SARec images are compared with the unfiltered SARec and IRACRR. The highest SBR was acquired

with SARec and IRACSCRR (TEW) reconstructions for 208 keV and the superior SNR results were

calculated from the 208 keV IRACRR and IRACSCRR images. For the two reconstructions including

both the 113 and 208 keV energy peaks, the SBR decreased in comparison with the single peak

reconstructions. When two energies were collected, the SNR increased in the SARec images while the

SNR for IRACRR decreased for most of the spheres.

a)

b)

c)

d)

Figure 4: SBR and SNR for different reconstructions and sphere sizes. SBR in the Jaszczak phantom (a), SNR in the Jaszczak

phantom (b). SBR and SNR for SARec images post filtered with a Butterworth (BW) filter of order 2 and a cut off frequency

of 0.03 cycles/mm are compared with unfiltered IRACRR and SARec (c-d).

The SBR and SNR results for the spheres in the Lung-Spine phantom are compared with the equal

sized sphere in the Jaszczak phantom in Figure 5. In general, the recovery was lower and the noise

higher the in the Lung-Spine phantom.

15

a)

b)

c)

d)

Figure 5: Comparison of the SBR and SNR between the homogenous Jaszczak phantom and the heterogeneous Lung-Spine

phantom. SBR in the 16 mL sphere (a), SNR for the 16 mL sphere (b), SBR for the 4 mL sphere (c) and SNR for the 4 mL

sphere are presented for different reconstructions.

In Figure 6, the SBR and SNR values for the reconstructions with altered number of projection angles

and acquisition times are presented. According to the t-test, the different number of angles did not

have any impact on the image quality parameters calculated. Shorter acquisition time did not have a

significant effect on the SBR values in the reconstructions, but there was a reduction in SNR for both

the IRACRR and SARec images.

A comparison of filtered and unfiltered SARec images for 208 keV, 120 angles and 30 s per projection

is presented in Figure 7. Figure 8-11 show images for the different angles and times approaches, the

SARec images are post filtered with a low pass butterworth filter of order 2 and with a cut off

frequency of 0.03 cycles/mm to mimic the smoothness in the IRACRR images.

16

a)

b)

c)

d)

Figure 6: SBR for different number of projection angles (a), SNR for different number of projection angles (b), SBR for

different acquisition times per projection angle (c) and SNR for different acquisition times per projection angle (d) for the

different sized spheres in the Jaszczak phantom. An asterisk (*) over the bar means that the difference from the

reconstruction with 120 angles and 30 s per projection is significant with a P-value ≤ 0.05. For the bars with no asterisk

there are no significant change of the image quality parameters.

Figure 7: 208 keV SARec reconstruction of the Jaszczak phantom; post processed with a Butterworth filter of order 2 and

with a cut of frequency of 0.03 cycles/mm (left) and unfiltered (right).

17

a)

b)

c)

d)

e)

f)

Figure 8: IRACCRR reconstructions for different acquisition times. Slices from the Jaszczak phantom are shown in a-c and

slices from the Lung-Spine phantom are shown in d-f. The different acquisition times are: 30 s (a and d), 15 s (b and e) and

10 s (c and f).

a)

b)

c)

d)

e)

f)

Figure 9: IRACCRR reconstructions for different number of projection angles. Slices from the Jaszczak phantom are shown

in a-c and slices from the Lung-Spine phantom are shown in d-f. The different number of projection angles are: 120 (a and

d), 60 (b and e) and 30 (c and f).

18

a)

b)

c)

d)

e)

f)

Figure 10: SARec reconstructions for different acquisition times. The images are post filtered with a Butterworth filter.

Slices from the Jaszczak phantom are shown in a-c and slices from the Lung-Spine phantom are shown in d-f. The different

acquisition times are: 30 s (a and d), 15 s (b and e) and 10 s (c and f).

a)

b)

a)

d)

e)

f)

Figure 11: SARec reconstructions for different number of projection angles. The images are post filtered with a Butterworth

filter. Slices from the Jaszczak phantom are shown in a-c and slices from the Lung-Spine phantom are shown in d-f. The

different number of projection angles are: 120 (a and d), 60 (b and e) and 30 (c and f).

19

Discussion

The results in this work show that both IRACRR and SARec reconstructions with two main energy

windows, over the 113 keV and 208 keV peaks, result in a lower SBR than if just a single main energy

window over the 208 keV peak is used, this outcome was predictable because of the high amount

of down scattered photons in the lower energy window. The expectation was that the scatter could be

properly handled with MC and it is shown in these results that SARec with two energies gave

improved results in comparison to IRACRR, but the 208 keV SARec and IRACSCRR (TEW)

reconstructions still gives the superior SBR. Nevertheless, one thing that can be noticed is that the

noise is lower for the SARec reconstructions where two energy peaks are sampled, this is probably

because of the increased number of photons in the image. To the author’s knowledge, TEW for two

peaks was not available at the workstation used but that might give even better recovery than the one

peak TEW. If the TEW scatter correction could handle the scattered photons around the 113 keV peak

properly the information from the acquisition would be utilized in a more effective way.

In previous studies, the SBR in the SARec reconstructions has been closer to 100 % than in this work,

but the relative difference between the SARec and IRACRR is in the same order as in previous results

[14]. The reduction in recovery could have been induced due to insecurities in the activity

measurement, another reason could be that the activities in the phantoms differed from the previous

studies [14]. Increased activity results in more data and the recovery would probably increase. This

effect though, did not appear when the acquisition time was changed in this study.

The SBR and SNR were stable with altered number of projection angles for both IRACRR and SARec

(Figure 6 a-b). Since the time per projection angle was adjusted to not change the total number of

detected photons, the signal should be approximately the same for the different angle approaches and

this corresponds to the results for both reconstruction methods in this study. To assure that eventual

effects on the image quality parameters were not induced because of a lower amount of data but only

the number of projection angles, the time per projection angle was changed in the measurements with

60 and 30 angles.

The SBR did not change when the time per projection was reduced; it is reasonable that the ratio

between the volumes does not change since the relative reduction of detected photons should be the

same in both the background and the hot volumes. There was a significant reduction in SNR when the

acquisition time was shortened for all images, except from the 15 s reconstruction in SARec (Figure 6

d). The test sample was small and the P-value was close to the 5 % limit (5.4 %) for the 15 s SARec

reconstruction, with more data (there was only 6 spheres in this study) the outcome may have differed.

The decrease of SNR was expected since there were a smaller number of detected photons in the

shorter acquisitions. In the SARec reconstructions, there was generally a bigger relative reduction of

SNR with shorter acquisition time. Visually, the 10 s images also gave a noisy impression for both

reconstructions and this could cause inaccuracies in a patient evaluation. For example, in Figure 8 f,

clusters are visible in the mediastinum area, which would make it hard to distinguish between tumours

and normal tissue in a clinical image.

An additional visual concern is that in the IRACRR images the shape of hot spheres often were

deformed, the spheres seemed to be extended towards the centre of the phantom and the shape did not

fully match with the shape in the CT. With shorter acquisition time the spheres seems to be more

deformed. For example, the third largest (4 mL) sphere in Figure 8 changes in shape between the

different time approaches. With less statistics, it appears to be more difficult to determine the location

20

of the hot areas. In dosimetry, incorrect activity distribution in the images could cause underestimated

or overestimated absorbed doses. In therapy with radionuclides radiopharmaceuticals are often

injected in the patient repeated times over a longer period, if the damage in a risk organ is

underestimated for every treatment the organ could lose its function due to the radiation.

Overestimation of the activity could also be problematic if further activity injected is reduced to spare

a risk organ, this would result in a lower absorbed dose to the tumours as well and less tumour cells

would be killed. The shape of the smaller sphere (4 mL) in the Lung-Spine phantom also was irregular

in especially the IRACRR images; this was visual in both the images with varied angles and times.

The boundaries between the small lung sphere and the spine were also vague in both reconstructions.

The irregularities of the small sphere in the lung phantom may be induced because of the

inhomogeneities around it, it is harder for the system to perform an accurate attenuation correction for

a more complex volume. Visually these deformation tendencies mentioned seems to be more

prominent in the IRACRR reconstructions, but no kind of measurement has been performed in this

study to confirm the differences observed. For further studies, it may be of importance to investigate

the deformation more and what impact it has on the quantification.

In SPECT based activity quantification, the desire is that the images reflect the true distribution of

radionuclides in the patient. Therefore, the SBR is of greater importance than the SNR while selecting

acquisition parameters and reconstruction methods. Scatter correction with TEW theoretically

increases the noise in the image but it also generates higher SBR. Using TEW in combination with a

decreased time per angle or number of projection angles would give a higher noise level which could

make it difficult to visually evaluate the images. Since 177Lu emits a low fraction of gamma photons

there already is a lot of Poisson noise in the lutetium images and it is desirable to not increase the

noise level. In this study though, the SNR values for the TEW images was not lower than for the

conventional reconstructions and the exclusion of scattered photons might have a greater effect on the

SNR in this case.

It is shown from the SBRs in Figure 4 c, that the SARec reconstruction are more capable to handle the

density differences in the lung area than the IRACRR, but with scatter correction (TEW) the Xeleris

managed to give results of the same quality as SARec. Both the SARec and Xeleris reconstructions

generally gave lower SBR values for the lung spheres but the differences were bigger in the Xeleris

reconstructions.

The SNR between the reconstruction methods is hard to compare since there is an unknown

smoothening of the Xeleris images in the reconstruction. If the SARec reconstructions are post filtered

the noise impression could be reduced (Figure 7) which makes it easier to identify volumes with high

activity uptake, obviously also the Xeleris images may be post filtered but it might not be necessary.

Drawbacks with filtered images are that they have lower SBR since some signal is ablated from the

image, see Figure 4 c. The quantification becomes poorer with lower SBR, in a clinical situation the

identification should be done based on filtered SPECT images and the CT images, and the

quantification should be done in the unfiltered images to avoid that signal disappears. Also, the SNR

decreased for the filtered image in this study (Figure 4 d) even though the visual impression is less

noisy, this outcome is probably due to the reduced signal. The SNR calculated in this study is based on

the signal concentration in the VOIs (counts/mm3), when the images are filtered some frequencies are

outsourced but that have only a minor impact on standard deviation between the background VOIs. If

filtration of images is performed, the filter must be selected with awareness since information is

separated out from the image. Both in noisy and smeared images it is hard to distinguish small lesions.

21

After a study at Tygerberg Hospital [16], they considered it possible to reduce the acquisition time per

angle from 30 s to 15 s. The image ranking at Tygerberg Hospital was performed by only one

physician, which is not considered enough to make a proper conclusion about the image quality.

Anyway, both in our and the Tygerberg study, the indication is that 15 s per projection angle could be

enough for clinical use but further evaluations must be done before a new protocol can be introduced.

A 10 s collection would probably give too much noise for clinical use. To reduce the number of

projection angles also seems like a viable option, but to gain an apparent saving of time, phantom

measurements where both the time and number of projection are decreased should be performed and

evaluated. The reason why this type of images not was collected and evaluated was because of the

time limitation of this work. In another study bone SPECT images reconstructed from the data

acquired from one and both detector heads were compared and they found no significant difference in

image quality [17]. In a bone SPECT, the circumstances are different from a lutetium examination, for

example 99mTc is used which emits photons with an energy of 140 keV in 100 % of the decays.

Nevertheless, the bone study indicates that both time and number of angles could be reduced in further

protocols. Before a new SPECT protocol can be introduced the approaches for the new acquisition

should be tested on first phantom and then patients to compare the images with those generated from

the original protocol. In addition to computation of image quality parameters, a visual evaluation also

should be done on the patient images to ensure that the reconstructions are good enough for clinical

use. To make two or more acquisitions on a patient is impractical for especially the patient and it is

time demanding. It would be preferable if the SPECT could be collected in list mode, but to the

author’s knowledge this is not possible in the gamma camera system used today at the Sahlgrenska

University Hospital. If desired projection angles could be distinguished and collected from the

sinogram it would facilitate the patient evaluation. To create reconstructions with only a part of the

projection angles is manageable if using SARec for the reconstruction, this will probably be done in

further studies. The applications mentioned above would make it possible to create reconstructions

with a varied time per projection and number of projection angles from only one examination.

22

Conclusions

The IRACSCRR (TEW) and SARec reconstructions of the 208 keV acquisition resulted in the highest

recovery. The highest SNR was calculated from the reconstructions with IRACSCRR (TEW) over the

208 keV peak and with IRACRR over both the 113 and the 208 keV peaks. The SBR decreased and

SNR increased when two energies was collected.

The SBR did not change when the number of projection angles or time per projection was decreased.

The SNR was stable with altered number of projection angles but it decreased with shorter acquisition

time because of the reduced number of detected photons.

All reconstruction methods performed inferior in heterogeneous volumes where both SBR and SNR

was lower than in a homogeneous volume.

Since the recovery was stable for the different acquisition approaches evaluated in this study it will

probably be possible to create a clinical protocol more time efficient than the one used today. The

noise is the primarily limiting factor for what acquisition parameters can be used. Further studies must

be performed, including patients and visual grading, before a new examination protocol can be

introduced.

23

Acknowledgements

I would like to express my gratitude to all the people supporting me during the project! First and

foremost, thanks to my supervisors Jakob Himmelman, Tobias Rydén and Peter Bernhardt for making

this work possible with your expertise and guidance. I would also like to thank Emma Wikberg and

Jens Hemmingsson for your help and advices with the phantoms. Thanks to Sture Lindegren for the

helping me with chemicals and the gamma counter. Finally, thanks to Arvid for all your love and

support through the whole project!

24

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25

Appendix

Figure 12: Intensity profiles over the 16 mL sphere and one of the line sources for the different reconstructions.

26

Table 2: Results over the measurements of adhered activity in PMMA, expressed in counts per minute (cpm).

BSA (g/L) Measured activity,

filled cuvette (cpm)

Measured activity,

emptied cuvette (cpm)

Adhesion ratio (%)

0 600215.6 1633.8 0.2722

0 605986.0 1062.4 0.1753

0.01 550902.7 496.9 0.0902

0.01 528878.2 577.5 0.1092

0.1 571257.3 668.4 0.1170

0.1 552779.4 434.4 0.0786

1 506136.5 2741.9 0.5417

1 529395.9 1499.3 0.2832

a)

b)

Figure 13: Unfiltered SARec reconstructions of the Jaszczak (a) and Lung-Spine (b) phantoms. For 120 projections angles

and an acquisition time per projection of 30 s.

27