QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS TREVOR THANG 1 Supervisors: Dr.

Transcription

1 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS TREVOR THANG 1 Supervisors: Dr. Eugene Wong 2, Dr. Rob Bartha 1 Department of Medical Biophysics 1, Department of Physics 2 Western University

2 INTRODUCTION Cancer is one of the leading causes of death in Canada. It is an endogenous disease which arises from the spontaneous mutation of normal tissue cells causing them to divide rapidly. About 40% of Canadians will develop cancer in their lifetime and over half of them will die from it. [1] Thus, cancer detection and treatment research has become very prevalent in today s society and is needed to help fight off this disease. One technique used to treat cancer is the use of Radiation Therapy. The protocol for radiotherapy is first acquiring a Computed Tomography (CT) scan of the tumor area and then creating a treatment plan which irradiates the tumor. Radiation can come in the form of x-rays or gamma rays, and the goal is to have these ionizing rays damage and destroy the DNA of the tumor cells, to prevent them from multiplying further. However, radiation cannot be perfectly targeted so some healthy cells are also irradiated in the process. Therefore, in order to avoid a lethal dose of radiation to healthy tissue, the process is spread out over a few weeks. [1] The long length of treatment allows for many variables to change throughout the treatment process. These factors include: patient position, patient weight, organ movement, tumor movement and tumor aggressiveness. Currently, once a treatment plan has been created by medical physicists and oncologists, the plan rarely get revised. For example, if the patient loses a lot of weight (a side effect of chemotherapy), instead of changing the dose to account for the weight loss, an entirely new treatment plan is create; increasing costs, wasting time and resources. If the tumor moves slightly to the left, the patient is moved slightly to the right in order to compensate. This is the established norm which is designed to maximize patient care and hopefully extend the life of the patient. 2 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

3 One key element missing from the radiation therapy is functional information. Currently, treatment plans only use anatomical and x-ray attenuation data from the CT image. Functional imaging could give insight towards the aggressiveness of certain areas of the tumor, as well as potentially allow for the detection of smaller tumors that can t be seen using traditional CT imaging. Also, this data can be tracked throughout the treatment process. Any changes seen can be reflected in the treatment plan by an appropriate adaption which best treats the immediate needs of the patient. However, one obstacle that all imaging scientist must always consider is whether the changes seen on the image are true biological changes or a result of inconsistent imaging environments. Of the many forms of functional imaging, the one being explored is Magnetic Resonance Imaging (MRI), specifically, Magnetic Resonance Spectroscopy (MRS). MRS can be used to detect in vivo metabolites, some of which may be flags for the aggressiveness of cancerous cells. The drawback to MRS is how many more variables the technology introduces. The biggest factor and the factor to be explored is the homogeneity of the main magnetic field (B 0 ). Furthermore, MRS data requires specialized programs to process and display them properly. Thus, it would be ideal to avoid actually processing and analyzing the MRS before first determining whether or not it is good data. The idea proposed is to use more conventional MRI data as surrogates for the quality of MRS data. Whether this can be done and how it can be done will be explored in this experiment. The overall purpose of the project is to calculate quality assurance metrics which can used as thresholds which govern the validity of the imaging data. This experiment will contribute to this project by determining a surrogate metric which can flag the presence of an 3 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

4 inhomogeneous magnetic field (B 0 ), thus giving insight towards the accuracy of the Magnetic Resonance Spectroscopy data. THEORY Magnetic Resonance Imaging (MRI), involves nuclear resonance within a strong magnetic field. The main magnetic field (B 0 ) generated by the MRI machine can range from 1.5T 7T. This magnetic field is typically generated by a super-cooled solenoid and by convention, the field generated is pointing in the ẑ direction of a Cartesian plane. Magnetic Dipole Physics In order to understand Magnetic Resonance (MR) physics, it is important to understand how a singular dipole exists in a magnetic field, and thus, this report will explain the physics of one dipole in a magnet. Though samples contain multiple dipoles, their effects will not be explained as it does not facilitate further understanding to the report. MR data is acquired by taking advantage of the magnetic dipole of certain molecules, mainly 1 H, 31 P, and 13 C. This magnetic dipole (µ) is produced by electrons orbiting the atomic nucleus. (Figure 1) [2] When the magnetic field (B 0 ) is turned on, the atomic dipole within the machine begins to align with the magnetic field. This is due to the fact that this is the most energetically favorable position for the dipoles to exist. (Figure 2) Figure 1: Atomic Magnetic Dipole created by Orbiting Electron 4 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

5 B 0 Figure 2: Magnetic Dipoles in a Sample Left: No B 0 Right: Large B 0 Once the dipole is aligned with the magnetic field, the MRI machine uses radiofrequency (RF) pulses to excite the dipole into a higher energy state such that it no longer aligns with the magnetic field but instead points at some angle away from B 0. (Figure 3) The most common RF pulse is a 90 flip angle pulse which rotates the dipole 90 or into the xy plane, where the magnetic field is defined as the direction of ẑ. Based on magnetostatics, the magnetic dipole (µ) will begin to precess about the z-axis because of a torque created by the magnetic field. This force and direction of this precession is defined by Figure 3: RF Pulse Rotating Magnetic Dipole (µ) Equation 1. 5 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

6 The frequency at which a dipole precesses (Larmor Frequency) is one of the most fundamental principles of MRI. The Larmor Frequency is dependent on each dipole s ( 1 H, 31 P, and 13 C) gyromagnetic ratio ( ) which relates the precession frequency and the magnetic field. The gyromagnetic ratio of 1 H is MHz/T. Torque of a Magnetic Dipole in a Magnetic Field Larmor Frequency Equation 1 is the torque vector on the dipole is the magnetic dipole vector is the magnetic field vector Equation 2 is the angular frequency (rads/sec) is the precession frequency (cycles/sec) is the gyromagnetic ratio is the magnetic field strength The precession about the z-axis however does not remain in the xy plane. As the dipole emits an electromagnetic signal, there is a loss of energy. As a result,, the magnetic dipole cannot remain in this excited state and thus begins to tilt back towards the main magnetic field. As it tilts back, there is a recovery of longitudinal magnetization and a loss of transverse magnetization. The rates of longitudinal recovery and transverse decay can be expressed using the exponential time constants T1 and T2 respectively. Summary: After a RF pulse (90 flip angle pulse), the magnetic dipole is excited into the xy plane and begins to precess at the Larmor frequency as defined by Equation 1 and 2. As the dipole begins to lose its energy, it must return to its ground state (aligned with the magnetic field). The rate at which the loss of magnetization in the transverse plane (xy plane) and the gain 6 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

7 of magnetization in the longitudinal direction (z) can be expressed as the exponential time constants T2 and T1 respectively. Magnetic Resonance Spatial Imaging [3] Magnetic Resonance Imaging is used to measure a variety of variables concerning magnetic dipoles. Common variables are T1 and T2 values and atomic diffusion. However, these signals from the sample have to be labeled in order to create a three dimensional image and this is done through the use of magnetic gradients. These gradients along the x, y and z axis are used to encode spatial data into the spinning dipole by change the magnetic field strength along any axis. Slice Encoding (Z Coordinate) A magnetic gradient is applied in the direction. This gradient causes slight variations in the magnetic field along the z axis and thus, there is a slight energy difference between the ground and excited dipole states. Thus, when a 90 flip angle RF pulse, with the chosen energy is applied to the sample, only the dipoles with that exact energy gap (between ground and excited) will precess and give off signal. As a result, any signal picked up from the sample must have been emitted from the known xy plane. It is important to note that the magnetic field is always pointing in the z direction; gradients simply change the magnitude of the magnetic vector. Phase Encoding (Y Coordinate) A magnetic gradient is applied along the direction. The slight changes in magnetic field affect the precession frequency of the dipoles as explained by Equation 2. This gradient is only applied for a short period of time so that the dipoles are no longer in phase with one another; instead, only dipoles in the same y co-ordinate will have the same phase. 7 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

8 Frequency Encoding (X Coordinate) The final magnetic gradient is applied along the x direction. This causes the Larmor frequency to change and thus some dipoles will rotate faster than others. When this final gradient is turned on, the signal is acquired thus allowing the frequencies to be encoded. Summary: In order to encode spatial information, MRI machines use 3 magnetic gradient encode a three dimensional Cartesian coordinate into the magnetic dipoles. The z coordinate is selected with the use of a specific RF pulse which only excites one slice of the sample. The y coordinate is encoded in the dipole phase which resulted from a temporary magnetic gradient, dephasing the dipoles. Finally, the x coordinate is encoded in the dipole frequencies in the presence of a third magnetic field. All of this information is summarized by Figure 4. Figure 4: Magnetic Resonance Imaging and Spatial Encoding 8 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

9 Chemical Shifts The examples and discussion used throughout this section will pertain to using 1 H as the dipole. However, the basic concepts remain the same for any dipole. Magnetic Resonance Spectroscopy (MRS) is a technique used to gather metabolic data in-vivo from a sample. It can detect protons within slightly different chemical environments and thus protons belonging to specific molecules can be identified and quantified. Typically, magnetic dipoles do not exist by themselves; instead, they are bound to a molecule. When bound to a molecule, the effects of neighboring atoms and their electrons are especially influential on the dipole and its magnetic environment. Electrons can partially shield nuclei from the magnetic field thus changing the effective magnetic field felt by the dipole. However, this effect is very minute: approximately 0.1% of the B 0. Though small, this change in magnetic field is what causes a slight change in Larmor Frequency which is the key to differentiating nuclei from one another. [2] This slight change in frequency can be quantified and must be experimentally determined. This change in frequency is called the chemical shift and is defined by Equation 3. Chemical Shifts Equation 3 represents the chemical shift (ppm) represents the change in frequency due to altered chemical environment (Hz) represents the frequency of the reference molecule (Hz) (for 1 H, the reference molecule is TMS, trimethylsilane) NOTE: The chemical shift values are independent of the main magnetic field strength. 9 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

10 For example, examine the molecule, choline below. (Figure 5) Figure 5: Choline with Highlighted Hydrogen Atoms All hydrogen atoms colored in with the same color are chemically identical due to identical bonding sequences. Thus, it would be expected that the signal received from the blue protons would by nine times the signal received from the green protons. [2] The chemical shifts for each set of protons have been determined by experimentally, the most important chemical shift is the one for the blue protons. This is because there is significantly more signal available and thus, is more noticeable on the trace. (Figure 6) The chemical shift for the blue protons is approximately 3.24ppm. [4] Figure 6: Data from a Prostate Cancer Patient illustrating the Choline Peak at 3.2ppm 10 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

11 Summary: Based on the peaks seen on an MRS trace, metabolites can be identified based on the frequency shifts due to their chemical environments. These shifts are quantified as chemical shift. METHODS In order to generate thresholds, it is important to understand how the data is currently being acquired and the condition/accuracy of the images taken. From London s Cancer Institutes of Health Research (CIHR), multiple types of MR data (T2w, DW-EPI) were gathered from prostate cancer patients who were imaged just prior to their prostatectomy. These images were acquired at Robart s Research Institute using a 3T General Electric (GE) MRI Machine. Of the many patients data files (~ 45), only a small subset (11) were identified to have had MRS data along with the standard MR images. Two types of images were identified to have the largest variability in accuracy: Diffusion Weight Echo Planar Imaging (DW-EPI) and Magnetic Resonance Spectroscopy (MRS). It was suspected that the two resulted in poor data because of inhomogeneous magnetic fields as supported by the MR Physics described in the Theory section of this report. MRS data is deemed poor if the peaks are not clearly resolved. This can result from having thick peaks that have a large Full Width at Half Max (FWHM) (Figure 7). Large FWHM can arise due to many variations in the magnetic field. The peaks blur together and thus, specific metabolites cannot be differentiated from one another. 11 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

12 Figure 7: Large FWHM & Blurring of Peaks 1 & 2 DW-EPI data is deemed poor when here is warping of the image. This is a result of poor spatial encoding. As mentioned previously, the spatial encoding is coded with the use of magnetic gradients. If the magnetic field is non-uniform, multiple additional gradients may exist within the machine which the computer processes unknowingly. The result is a distorted image, such as Figure 8 below. Figure 8: Large Deformation in DW-EPI Since both imaging modalities (MRS and DW-EPI) are theoretically susceptible to an inhomogeneous magnetic field, it is hypothesized that the quality of the two may be correlated. 12 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

13 The quality of the images can be quantified using the Full Width at Half Max (FWHM) for MRS and Volumetric Distortion for DW-EPI. Gathering the FWHM Water and Lipid The MRS data gathered from the MRI is imported into a computer as a P-File, a specially encoded data format which GE uses to protect their data. Using a program developed by Dr. Rob Bartha called fitman, the P-File is converted in a more appropriate format which fitman is able to open and display. A sample raw data is presented in Figure 9. A noise filter is then applied in order to reduce the noise as shown in Figure 10. The largest peak is from water (4.7ppm) within the prostate. The FWHM of this peak is acquired. Afterwards, in order to gather information about smaller peaks, such as the lipid peak, the water peak must be removed from the spectrum using another filtering tool. (Figure 11) Finally, the FWHM of the Lipid peak (1.5ppm) can also be acquired. Figure 9: Raw MRS data from Prostate Figure 10: Noise Filtered MRS Water FWHM Figure 11: Water and Noise Filtered MRS data Lipid FWHM Quantifying the Magnetic Distortion There are many ways to quantify the amount of distortion within the image but the method used in this experiment is a volumetric percent change. Using the Cancer Treatment 13 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

14 Planning System, Pinnacle, manual contours of the prostate were made in two sets of MR images: DW-EPI and T2 Weighted Fast Spin Echo (T2w-FSE) (Figure 12). Unlike DW-EPI which is susceptible to an inhomogeneous magnetic field, T2w-FSE is not due to a special pulse sequence known as a 180 Refocusing Pulse. Since the T2w-FSE volume is closest to the truth, it will be used as the gold standard which the DW-EPI will be compared with. The volumetric percent change was calculated using Equation 4. Figure 12: Manual Contours on Pinnacle of DW-EPI (left) and T2w-FSE (right) Volumetric Percent Change Equation 4 represents the volumetric percent change represents the T2w-FSE Volume of the Prostate represents the DW-EPI Volume of the Prostate The Statistics After completing these important appropriate measurements, statistical analysis was performed to determine whether or not a correlation exists. Also, based on the resulting trends, a suitable threshold was determined. 14 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

15 Volumetric Percent Change (%) RESULTS The first correlation was done with the FWHM of the water peak and the Volumetric Distortion. As the water peak is typically the largest peak, it is a good place to start. The correlation is shown below in Figure Volumetric Distortion as a Surrogate for MRS Quality - Water %Δ = 1.03 FWHM R² = FWHM of Water Peak (Hz) Figure 13: Correlating FWHM of the Suppressed Water Peak with the Volumetric Distortion Plotting FWHM of Suppressed Water Peak vs. Volumetric Percent Change shows a positive correlation with an R 2 value of ( N=11) It is important to note that the x-intercept is not 0. This is due to MR Physics which even in a perfectly uniform magnetic field, prevents a infinitely thin peak from existing. The second largest peak, lipid peak was also analyzed. This was done for consistency to ensure that the FWHM dependency was not associated with certain metabolites. The results of the lipid correlation are shown below in Figure QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

16 Volumetric Percent Change (%) Volumetric Distortion as a Surrogate for MRS Quality - Lipid %Δ = 0.52 FWHM Point A R² = FWHM of Lipid Peak (Hz) Figure 14: Correlating FWHM of The Lipid Peak with the Volumetric Distortion Plotting FWHM of Lipid Peak vs Volumetric Percent Change shows a posistive correlation with an R 2 value of ( N=11) One outlier is present amoungst the data. Though the correlation between the 2 variables is slightly weaker (R 2 = 0.63), the correlation is definitely still present. The reason for a weaker correlation is due to an outlier present (Point A). There are many factors which could have resulted in such a large volumetric percent change (ie. patient movement). Thus, in order generate a threshold, outliers need to be ignored. Therefore, Point A was removed from this data set as shown in Figure QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

17 Volumetric Percent Change (%) Volumetric Distortion as a Surrogate for MRS Quality Lipid (Revised) %Δ = 0.67 FWHM R² = FWHM of Lipid Peak (Hz) Figure 15: Correlating FWHM of The Lipid Peak with the Volumetric Distortion - Revised Plotting FWHM of Lipid Peak vs Volumetric Percent Change shows a posistive correlation with an R 2 value of ( N=10) This change results in a much stronger correlation with an R 2 value of However, with only 10 data points, the correlation would be strengthened if the number of patients available were doubled. 17 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

18 DISCUSSION Since there are strong correlations between MRS FWHM and DW-EPI distortions, quality assurance thresholds can be generated. The R 2 values for the water peak (0.82) and for the lipid peak (0.87) suggest that there is likely to be an underlining cause which is damaging both imaging techniques. It is hypothesized that it may be an inhomogeneous magnetic field, but this was not proven during this experiment. Based off of numerous scientific reports, the typical FWHM for a suppressed water peak should be approximately 8 Hz. [5] Thus, based on the line of best fit in Figure 13, an appropriate threshold for Volumetric Percent Change would be 7%. Of the 11 patients (2 data points are identical), 7 were deemed to be acceptable since they were above the threshold. A similar result was seen for the revised lipid peak data. With a threshold of 7% and the corresponding lipid FWHM threshold of 35Hz, the same 4 patients were deemed unacceptable. In regards to omitting the Point A, the goal of this project is to create a set of thresholds that can be used to quickly check the validity of the MR data. Thus, when generating each threshold, it is not necessary for each threshold to individually differentiate all the good data from the bad. It is hopefully, a cumulative effort of all the quality assurance metrics which gives the best complete picture of the image accuracy. Although, it is expected that even though by our threshold, Point A would be deemed unacceptable (greater than 7% change), other quality assurance metrics could help identify it as being good MRS data (lipid FWHM less than 35Hz). There are many revisions that can be made to improve the results, such as how volumetric distortion was quantified. In this experiment, a volumetric percent change was employed but may not reflect a true value for distortion as shown in Figure 16. Alternatives 18 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

19 include Surface Area-to-Volume ratio and Percent Overlap. Correlations for these new metrics should also be done. Figure 16: Equal Volumes (0% Percent Change) with Large Distortion The scale of this project is very large and thus the results presented in this paper are merely the first of many potential quality assurance metrics. In order to validate this threshold, Receiver Operator Curve analysis must be performed, more patient data must be gathered to strengthen the correlation value. Furthermore, this threshold may be replaced in the future as there are special pulses sequences have been developed that can directly measure the homogeneity of the magnetic field. These techniques include Phase Difference Method ( [6] and Bandwidth-Difference ( ) [7] However, at the time of this research, this data was not being acquired at Robart s and thus a surrogate had to be used. Whether or not these new types of data are beneficial is up to further research which must balance the amount of new information gained with the increased patient scanning time. CONCLUSION In this experiment, DW-EP images were used as a surrogate for MRS quality. The reason for this is due to the hypothesized effect of an inhomogeneous magnetic field on the two data types. Using data acquired from the London s CIHR project, images taken at Robart s were processed and metrics were calculated to quantify the quality of both the DW-EPI and MRS images. Volumetric percent change and FWHM metrics were shown to have a strong correlation. An R 2 19 QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS

20 value of 0.82 for the water peak and 0.87 for the lipid peak (0.87) allowed a threshold of 7% volumetric percent change to be calculated. Of the 11 patients, 4 were deemed to have unacceptable data. Further studies for this project include alternative quantification of distortion, analyzing the FWHM of functional metabolites such as citrate and choline, and exploring the effectiveness of new pulse sequence which can directly measure the magnetic field. As more of these quality assurance metrics are determined, medical physicist and oncologist will have more flexibility in their treatment plan. As long as their treatment plan fits within these metrics, the physician is free to adapt the plan giving more power and meaning to the world of Adaptive Radiotherapy. REFERENCES [1] Canadian Cancer Statistics 2011: Featuring Colorectal Cancer, Canadian Cancer Society. Accessed: Feb [2] Dr. Neil Gelman. MEDBIO 3505F Lecture Notes: Fall Western University [3] Functional Magnetic Resonance Imaging. Scott A. Huettel. Sinauer Associates, Inc. Sunderland, Massachusetts, USA. [4] MR Spectroscopy Quick Guide. Siemens Medical. [5] Prostate cancer metabolite quantification relative to water in 1 H-MRSI in vivo at 3 Tesla. Mclean. [6] Magnetic Resonance Imaging Quality Control Manual (American College of Radiology), Weinreb, 1991 [7] Routine Testing of Magnetic Field Homogeneity on clinical MRI Systems. Hua-Hsuan Chen, QUALITY ASSURANCE OF MAGNETIC RESONANCE IMAGING FOR ADAPTIVE RADIOTHERAPY: PRELIMINARY INVESTIGATIONS