Designing iron oxide nanoparticles for image guided thermal medicine applications
The impact of magnetic iron oxide nanoparticle (IONP) core design (size and shape) on magnetic resonance imaging (MRI) and heating characteristics is being actively pursued by many research groups. This work evaluates MRI relaxation and the specific absorption rate (SARFe) properties of IONPs as a function of diameter (6 -32 nm). SARFe increased with IONP size up to a maximum of 628 ± 11 W/g Fe for the largest IONP studied (33 nm). IONPs with a diameter > 10 nm were demonstrated to have transverse relaxivity (r2) values suitable for low-concentration IONP imaging applications, such as cell tracking. On the other hand, the optimal IONP diameter for achieving longitudinal relaxivity (r1) was dependent on magnetic field strength. Measurements of r1 values performed at a clinically relevant magnetic field (1.4 T) indicate all IONP diameters were more sensitive than gadolinium-based contrast agents (r1 ≥ 2.9 mMFe -1 s -1 ) widely used in clinical imaging. However, such a large r1 was not observed when measured at a larger (pre-clinical) magnetic field (9.4 T). We conclude that the ideal IONP diameter for image guided heating applications is dependent on the magnetic field strength of the MRI for the intended application.
Introduction<!>Heating<!>Transverse Relaxivity (r2)<!>Longitudinal Relaxivity (r1)<!>Heating and Relaxivity Characteristic Relationship<!>R1 Image Guided Heating<!>Conclusions<!>Materials<!>IONP Synthesis and Characterization<!>Sample Preparation<!>Heating Characterization
<p>Magnetic iron-oxide nanoparticles (IONP $ s) have considerable potential as a clinical theranostic platform capable of combining magnetic resonance imaging (MRI) contrast with hyperthermia therapy [4][5][6] . Clinically, IONPs have been applied for anemia treatment, cancer diagnosis, and are currently in clinical trials for cancer treatment through hyperthermia 7 . Therefore, the impact of IONP design on MRI and heating characteristics continues to be actively explored [8][9][10][11][12][13][14][15][16][17][18] . An important question we consider herein is if a single IONP can be optimized for diagnostic MR imaging and therapeutic heating applications within the same concentration range.</p><p>Traditionally, IONP-based MRI contrast agents are used to provide a hypo-intense signal (negative contrast) caused by rapid signal decay from transverse (T2) relaxation of the water protons 19 . For example, Feridex and Resovist were approved for clinical diagnosis of cancer lesions through negative contrast 20,21 . Similarly, GastroMARK was approved for diagnosis of masses within the gastrointestinal track as an agent that produced negative contrast 22 . Unfortunately, competing imaging methods, gadolinium-based contrast agents (GBCAs) in particular, resulted in the removal of these products from most markets 7 . However, recent safety concerns with GBCAs as contrast agents have resulted in a reemergence of IONP-based contrast agents. Specifically, Feraheme, an anemia treatment, is currently undergoing clinical trials for applications in cancer diagnosis (pancreatic, bone, and metastatic), central nervous system disorders, and angiography 7 . All of these applications focus on the use of echo-based pulse sequences, such as gradient echo and spin echo sequences. The negative contrast produced with echo-based MR imaging methods is capable of localizing IONPs and can quantify IONPs existing in low concentrations (≤ 0.018 mMFe) through the measurement of the transverse relaxation rate constant, R2 = 1/T2 23 . Recent developments of very small IONPs and echoless pulse sequences have produced positive contrast with longitudinal (T1) relaxation 19,24 . Positive contrast can provide IONP concentration at higher concentrations (< 60.7 mMFe) through the measurement of the longitudinal relaxation rate constant, R1 = 1/T1 4,19,25,26 .</p><p>Relaxation rates as a function of IONP concentration are known as relaxivities (r1 or r2). Relaxivities are defined as the slope of the linear regression of the relaxation rate constant (Rx, where x = 1, 2) and the concentration of the contrast agent ([IONP]):</p><p>where Rx0 is the relaxation rate constant of the media without the IONP and rx is the relaxivity. Positive (T1weighted) contrast is obtainable with echo-based MRI sequences when the contrast agent has a low r2/r1 ratio (e.g., r2/r1 = 2) 24 . However, the ability to produce positive contrast with echo-based MRI sequences is limited for commercially available IONPs with typical ratios r2/r1 = 59 -1429 26 . Recent work has demonstrated an ability to perform positive contrast IONP quantification, even at high r2/r1 ratio and high IONP concentration (as high as 60.7 mMFe = 3.4 mg Fe/mL), through the use of ultra-short echo or echoless pulse sequences, such as sweep-imaging with Fourier transform (SWIFT), which measure R1 19,[27][28][29][30] .</p><p>Magnetic fluid heating (MFH) is a technique where heat, measured as the specific absorption rate (SARFe), is produced by the interaction between an alternating magnetic field and IONPs 31,32 . MFH has been applied to both hyperthermia cancer therapy and nanowarming of cryopreserved tissues 5,12,33,34 . Both applications benefit from a good estimate of thermal dose to avoid unintentional injury to surrounding tissues. Magnetic particle imaging (MPI) shows great promise as a means to image IONP localization and control heating, but it is still under development for use in the clinic 35,36 . Computed tomography (CT) is capable of imaging high IONP concentrations (> 18 mMFe, 1 mg Fe/mL); however, as is the case with the high heating IONPs described herein, concentrations below 1 mg Fe/mL are capable of causing unintended tissue damage through heating 37 . Measurement of R1 through echoless pulse sequences have been shown to correlate linearly with thermal response in agarose samples and tissues 4,25,38 .</p><p>The mechanisms by which IONPs produce heat via MFH or cause MR relaxation are different, although both relate to the fundamental magnetic properties of the IONP. MFH produces heat by exposing IONPs to a kHz magnetic field in the absence of a background magnetic field. The applied alternating magnetic field causes the magnetization of the IONP to change orientation, whereas the amount of heat produced is dependent on the magnetic hysteresis of the material at the applied AC magnetic field strength and frequency 39,40 . From a simplified view, MRI is a spectroscopic technique exploiting the Zeeman energy splitting of protons created with a static magnetic field, typically ≥ 1 T. MRI signal is generated by applying a radio frequency pulse in the MHz range. In the static magnetic field used in MRI, the IONP is magnetically saturated and therefore will not respond to the kHz magnetic fields typically applied during MFH. On the other hand, the presence of the IONP accelerates the MRI relaxation processes. Certain physical properties, like magnetic susceptibility, affect both MRI relaxation and heating; therefore, a single particle might be designed to perform well for both of these applications.</p><p>Many IONPs have been developed and assessed for both MRI contrast and MFH 4,9,19,25,38,[41][42][43] ; however, these applications are typically considered separately and not within the same concentration range. Additionally, IONP size is known to have an impact on magnetic properties impacting both MRI and MFH 9,18 . Herein, we take a comparative look at the heating and imaging properties of IONPs as a function of IONP core size while maintaining a faceted structure. There results herein enable us to optimize IONP size for the application of magnetic resonance image guided magnetic fluid hyperthermia.</p><!><p>Previous work demonstrated that larger diameter IONPs (33 nm) produced SARFe that is much higher than that of other published IONPs 2,9,14,41 . This observed high heating is attributed to the stability from the DSPE-PEG2000 coating and the ferromagnetic nature of these IONPs 9 . Analysis of all IONP diameters was performed at the University Minnesota with a lower frequency setting (20 kA/m and 190 kHz) to allow a direct comparison with heating measurements performed with commercially available Ferrotec EMG-308 25 . The SARFe was observed to increase dramatically between 11 -30 nm, similar to the trend described by Tong et. al. (Figure 1) 9 . The change in experimental parameters, such as background matrix, magnetic field strength, and frequency, explain the observed lowered SARFe 2,39 . Furthermore, the SARFe observed for the 32 nm diameter IONPs is three times higher than commercially available IONPs, such as EMG-308 (Ferrotec), BNF-Starch (Micromod), nanomag ® -D-spio (Micromod), and Feridex (Advanced Magnetics) 14,25 (p < 0.001). The increased SARFe allows for a reduction in necessary IONP concentration for magnetic fluid hyperthermia and nanowarming applications.</p><!><p>At the preclinical (9.4 T) field strength, the r2 is observed to increase with IONP diameter. A similar trend was observed at the clinical (1.4 T) field strength; however, a plateau was measured for IONP diameters above 10 nm (Figure 2). Furthermore, r2 is anticipated to be constant with magnetic field strength because the saturation magnetization for each IONP is far below (< 5 mT) 1.4 T 18 . The differences observed between 1.4 and 9.4 T can be attributed to the difference in temperature for the experimental setup (37 and 22 o C, respectively). The temperature dependence of r2 has been demonstrated to be stronger for larger IONP diameters 3,18 .</p><p>The r2 values shown at 1.4 T for diameter > 6 nm demonstrate an r2 which is larger and therefore has an improved sensitivity when compared with EU-approved Resovist (61 mMFe -1 s -1 ) and FDA-approved Feridex (41 mMFe -1 s -1 ) 3 . In addition to applications as an MRI contrast agent, IONPs with a high r2 are of particular interest at low concentrations for labeling molecular markers for disease diagnostic MRI. Molecular marker applications measure the migration of cells over time by loading the cells with IONPs before injection 44 . In general, r2 ≥ 324 mMFe -1 s -1 is considered a reasonable relaxivity to facilitate the measurement of single cells 45 . However, higher r2 is better for molecular marker applications because the necessary IONP uptake to allow cell tracking is reduced 41 . Therefore, IONPs > 6 nm would be feasible for cell tracking at 1.4 T and highly applicable at preclinical field strengths (9.4 T).</p><!><p>Gadolinium-based contrast agents are used clinically to provide positive (T1-based) contrast with typical r1 values ranging from 2.9 -4.7 mM 3 . In general, IONPs often are capable of providing more sensitivity with a higher r1. However, their clinical use is limited to r2 due to their higher r2/r1 ratio (see Table S1). Some IONPs have been specifically designed to mimic the r2/r1 ratio (~1) of GBCAs to produce positive contrast 3 . Similar r2/r1 ratios (2) have only been demonstrated for IONPs with very small diameters (3 nm). These IONPs have low r1 (5.2 mMFe -1 s -1 ) and r2 (10.5 mMFe -1 s -1 ), impacting their sensitivity 24 , although they remain comparable to clinical GBCAs.</p><p>Ultrashort-echo and echoless pulse sequences have been developed with zero or negligible acquisition delay, allowing for the detection of signals with rapid T2 relaxation to obtain T1-weighted images 4,18,19,29,30,[46][47][48] . Acquisition of T1-weighted images allows for IONP design which can focus on optimization of r1. At 9.4 T, the echoless sequence SWIFT has demonstrated a quantitative range from 0.1 to 35.7 mMFe (0.05 to 3.0 mgFe /mL) using R1 (=1/T1) mapping 26 .</p><p>The comparison between r1 acquired at 1.4 and 9.4 T demonstrates a much lower r1 at the higher magnetic field (Figure 3). Of note, at 1.4 T, IONP core sizes > 10 nm maintain an r1 above values typically observed for clinical GBCAs at 1.5 T 3 . The observed change in r1 with magnetic field strength is consistent with theory described by Gillis et. al. and observations by Smolensky et. al. 18,49 . As higher magnetic fields are applied, one of the mechanisms of relaxation, Néelian relaxation, is inhibited because the magnetic moment of the IONP is locked into the same orientation as the main magnetic field. Therefore, the relaxation rate constant and relaxivity decreases with field strength 50 .</p><p>At both field strengths, we observe a local maximum and then a decrease in r1 with IONP diameter (Figure 3). The change between the values at high field strength is small enough (0.2 -0.8 mMFe -1 s -1 ) to be considered invariant. Smolensky et. al. compared faceted IONPs with spherical IONPs demonstrating different trends 18 . Our IONPs have a different surface coating for stabilization, but were observed to be faceted (see Figure S1). Our IONPs follow a trend similar to the faceted IONPs synthesized by Smolensky et. al. within IONP diameters of 4 -18 nm. However, the increase in r1 with IONP diameter is no longer maintained once ferromagnetic character is exhibited. The ferromagnetic character for the IONP synthesized has been demonstrated previously for diameters larger than 18 nm based on magnetic hysteresis characterization 9 . Therefore, the decreasing trend observed at 9.4 T for IONPs >18 nm matches the expectations for an IONP with a high anisotropy as is observed with ferromagnetic particles 18,51 .</p><!><p>Our results provide the capability to evaluate the correlation between SARFe and r1 or r2. The SARFe has been demonstrated to have a direct correlation with IONP diameter, independent of the change in field strengths used by Tong et. al. and within this manuscript 9 . Similarly, r2 demonstrates an increasing trend with IONP diameter despite the background magnetic field. In contrast, the IONP diameter which corresponds to the maximum r1 shifts with magnetic field strength. Therefore, a correlation between SARFe and r2 is conceivable and worth further consideration.</p><p>The dependence on SARFe and r2 on experimental parameters, such as magnetic field strength and frequency, create challenges for directly comparing results with other research groups. Our dataset demonstrates the challenge regarding the changes in r2 with field strength and differences in SARFe of the same IONP measured at both the University of Minnesota (Figure 1) and Rice University  41,42 . However, the direct numerical comparison is complicated by changes in the experimental setup and IONP synthesis and, therefore, magnetic characteristics. Changes in IONP crystal structure cannot be observed through SQUID, Mössbauer spectroscopy, TEM, or x-ray diffraction analysis, but can be measured with small angle neutron scattering have been demonstrated to still have an impact on SARFe 52 . Furthermore, changes in surface anisotropy caused by the attachment of a surface coating are known to impact both SARFe and r2 8,38,53 . Therefore, a robust demonstration of the correlation between SARFe and r2 requires a comprehensive study which must incorporate small angle neutron scattering to fully characterize the shape and crystalline structure of the IONP; that study is beyond the scope of work presented here and will be the subject of future research.</p><!><p>Several groups have demonstrated multifunctional IONPs capable of both heating and imaging; however, the imaging capability focused on negative T2-weighted detection 41,42 . Additionally, others have reported r1 properties; however, they do not endeavor to use both heating and imaging within the same concentration range 43 . Herein, our goal is to analyze IONPs using R1 mapping to provide quantitative information within a reasonable IONP concentration for magnetic fluid hyperthermia 4 . The heightened heating produced by these IONPs provide an improved overlap in IONP concentration range of R1 mapping and measurable heating.</p><p>The feasibility of using an R1 map to predict the expected temperature change after 3 minutes of heating is shown in Figure 4. The change in temperature observed for samples at concentrations above 16 mMFe (0.896 mg Fe/mL) are high enough to damage tissues 1,2 . Agarose samples are a simplification compared to the complex milieu of the intended biological systems for magnetic fluid hyperthermia. However, previous work has demonstrated the ability to work within an in vivo system to create a calibration between an R1 map and heating capabilities 4 . Therefore, this simplified system informs the size selection of this IONP before moving forward to pre-clinical studies connecting imaging and heating parameters.</p><p>Due to the r1 field dependence, the ideal IONP diameter is dependent on the applied field strength. At the preclinical field strength (9.4 T), IONP diameters of 19 and 33 nm were observed to have such a low r1 that the change in R1 based on IONP concentration is negligible. Therefore, a heating prediction cannot be ascertained, and a trade-off is necessary between selecting the highest heating IONP (32 nm, which has a very low r1) and a low-heating IONP (6 nm, which has an acceptable r1). At the clinically relevant field strength (1.4 T), the heating does not need to be sacrificed for better relaxation sensitivity because all of the IONP diameters measured had a sufficiently high r1 to obtain adequate sensitivity for imaging. Thus, for MRI applications at 1.4 T, IONP diameter could be optimized based on heating properties.</p><!><p>This work represents the characterization of IONPs toward an optimization for T1-weighted MR imaging and MFH. IONPs were observed to provide increased heating as the diameter increased, with IONPs with a 33 nm diameter demonstrating three times more heat than commercially available IONPs. In contrast, IONP cores > 10 nm had sufficiently high r2 to provide utility for MRI detection in low concentration, as needed in cell tracking. Further, the r1 values measured at the clinically relevant field (1.4 T) were larger than the r1 values of GBCAs widely used in clinical imaging. However, such a large r1 value was not observed at the pre-clinical field (9.4 T). In summary, IONPs with satisfactory heating and MR imaging performance have been demonstrated herein, however; there remains future opportunities to develop IONPs with optimized performance for both types of applications.</p><!><p>Iron acetylacetonate (Fe(acac)3, 99%), 1,2-tetradecandiol (technical grade, 90%), oleic acid (technical grade, 90%), oleylamine (technical grade, 70%), benzyl ether (98%), hydrochloric acid (> 37%), hydroxylamine HCl, sodium hydroxide, ammonium acetate, and agarose (CAS:9012-36-6) were purchased from Sigma-Aldrich (St. Louis MO) and used as received. All water used was MilliQ filtered water.</p><!><p>Magnetite nanocrystals were synthesized and characterized as previously described 9 . XRD and heating characterization were not anticipated to change significantly between batches. Water-dispersible IONPs were generated by coating the magnetite nanocrystals with DSPE-PEG 2000 using a dual solvent exchange method. Table 1 lists the specific IONP sizes investigated and the corresponding size and polydispersity. Size characterization was performed on each batch of IONPs using transmission electron microscopy (TEM, Figure S1). TEM images were acquired using a transmission electron microscope (Hitachi H-7500) at the Rober P. Apkarian Integrated Electron Microscopy Core at Emory University and a transmission electron microscope (JEOL 1230) at Rice University. The average diameter of the nanocrystals was calculated based on the measurements of at least 300 nanocrystals using image processing software (ImagePro Plus). The iron concentration of the IONPs was determined based on the Ferrozine assay with validation using ICP-OES 54 . Before iron concentration determination, the IONPs were digested in 0.3 M hydrochloric and ascorbic acid for two hours at 60 o C. ICP-OES of the iron concentration was performed with a Thermo Scientific iCAP 6500 dual-view ICP-OES (West Palm Beach, FL, USA) with 1150 W power.</p><!><p>Heating and 9.4 T relaxivity measurements were performed with IONPs suspended in 1% agarose at concentrations ranging from 0.018 to 89.3 mMFe (0.001 to 5 mg Fe/mL). 1.4 T relaxivity measurements were performed with IONPs suspended in water at concentrations ranging from 0.071 to 8.9 mMFe (0.004 to 0.500 mg Fe/mL). The lower concentrations of IONP were necessary at 1.4 T due to hardware limitations related to the measurement of rapidly decaying spins. Heating measurements were performed on all but the smallest IONPs because those below 10 nm produce negligible heat 9 .</p><!><p>A 1 kW Hotshot inductive heating system with a 2.75-turn, water-cooled copper coil (Ameritherm Inc., Scottsville, NY) was used to generate the alternating magnetic field. Reported SARFe data were acquired at 20 kA/m field strength (volume-averaged peak-amplitude across the sample) with an applied frequency fixed at 190 ± 10% kHz 25,38 . SARFe was measured for IONP concentrations ranging from 1.8 -89.3 mMFe (0.1 -5 mg Fe/mL). Each 1 mL sample in a plastic microcentrifuge tube was insulated and centered in the inductive coil and heated for 3 minutes with the temperature recorded every second with a T/Guard 405 Fiber Optic Monitoring System (Qualitron, Quebec, Canada). The SAR was calculated using Box-Lucas fitting for 55 s after heating began.</p><p>Validation of the high heating properties of the 40 nm IONPs was conducted at the University of Minnesota using settings and analysis consistent with practices at Rice University. The SARFe was acquired in water at 20 kA/m and 360 kHz by calculating with the Time-rise fitting for 20 s after heating began. At the University of Minnesota, the SARFe was measured to be 2125 ± 73 W/g Fe, slightly lower than the 2560 W/g Fe reported in Tong et. al. 9 . The slight difference in heating characterization is likely attributed to small differences in laboratory protocol and experimental setup (i.e. the uniformity of the field, higher harmonics in AC field, or heat losses within the system) between the labs at Rice University and University of Minnesota 55 , but the measured SARFe is extraordinary compared to previous work with other IONPs. Unfortunately, the 40 nm IONP was unstable when moved into the 1% agarose medium and could not be compared directly with the data demonstrated herein.</p><p>Relaxivity Characterization at 9.4 T MRI, including R1 and R2 measurement, was performed with a 9.4 T-31cm bore MRI scanner (Agilent Technologies, Santa Clara, CA). All images were acquired with a volume transmit/receive coil having an inner diameter of 3 cm (Agilent Technologies, Santa Clara, CA).</p><p>A multi-slice T2-weighted spin echo (SE) sequence was used to measure R2 maps. Each 2D image was acquired with a repetition time (TR) of 2.4 s, acquisition bandwidth of 50 kHz, an acquisition time of 2.56 ms, a slice thickness of 5 mm, and a resolution of 417 x 417 µm. For R2 mapping, the echo time (TE) was arrayed across 16 points spaced logarithmically between 12 and 800 ms. All 2D images were reconstructed using VnmrJ versions 3.2.</p><p>3D T1-weighted images and R1 maps were acquired using a Look-Locker method together with a MultiBand (MB)-sweep imaging Fourier transform (SWIFT) sequence for readout. The MB-SWIFT flip angle was 1 o , acquisition delay ≈ 2 µs, acquisition bandwidth = 384 kHz, TR = 1.2 ms, gaps = 2, voxel resolution = 194 x 194 x 1172 µm, and total acquisition time ~7 min 19,48 . The field-of view was 50 x 50 x 150 mm 3 with image matrix size = 128 x 128 x 128 x 64 (x,y,z,t). Two different MB-SWIFT settings, which impact the time points acquired, were also adjusted (Nspiral = 128 and 32 and Nv = 1024 and 4096, respectively). The Nv = 1024 setting had 64 time points spaced linearly from 12.7 to 1152 ms. The Nv = 4096 setting had 64 time points spaced linearly from 39.8 to 4596 ms. MB-SWIFT images were reconstructed using an in-house program written in MATLAB (2012b) 48 .</p><p>For each voxel, the time points were fit to the exponential relaxation curve using a three-variable fit 19 . The relaxation rate constants (R1 or R2) were determined with least-squares fitting. The region of interest (ROI) assessed for each tube was approximated as a cuboid with dimensions 2.05 x 2.05 x 5 mm. For SE, a 6 x 6 voxel square ROI was selected in the center of each tube for a single slice. For MB-SWIFT, the 12 x 12 voxel square ROI was selected in the center of each tube across 4 slices. The relaxivity (r1 or r2) was determined by performing a linear least-squares fitting of the relaxation rates as a function of iron concentration.</p><p>Relaxivity Characterization at 1.4T</p>