France



&lt;br/&gt;Magnetic hyperthermia is a promising cancer therapy gaining great interest in recent years using heat generation ascribed to magnetic particles subjected to a high-frequency magnetic field. To make this therapy feasible in clinical settings, in addition to magnetic particles, important elements include techniques for detecting their position and temperature in a tumor region in determining the effectiveness of therapeutic heating. In previous studies, we developed thermosensitive magnetic micro/nanoparticles with high heating efficiency and considerable permeability change around the therapeutic temperature range of 42-45°C [1]. We also proposed a wireless temperature measurement method for monitoring the tumor temperature using pickup coils to detect the temperature-dependent stray magnetic field caused by the magnetic particles [2].&lt;br/&gt;A challenge remains when the particles injected into the tumor region deviate from the longitudinal axis of drive coil (heating coil). As a result, the heating generation of the particles and the accuracy of the proposed thermometry are decreased. To solve this problem, by taking the advantage of the fact that the induced voltage in pickup coil depends on the position of the particles, we developed a localization system using three pickup coils symmetrically installed inside drive coil. The position of particles could be automatically localized by using a robot arm to operate the unit composed of drive coil and pickup coils in two modes of coarse adjustment (rotary scanning) and fine adjustment (linear scanning) [3].&lt;br/&gt;In this study, we propose a simpler, faster and more cost-effective localization method by using only figure-8 coil instead of three pickup coils (loop type coils) and operating the unit of drive coil and figure-8 coil (hereafter the unit) only in linear scanning mode. Fig. 1 shows the conceptual diagram of the proposed localization method using figure-8 coil. Figure-8 coil consists of two reverse-winding loop coils and is adjusted so that its initial bias (the voltage without sample) close to 0 V and fixed to the center of drive coil. As a result, the induced voltage changes depending on the relative position of the implant (i.e., magnetic particles) as seen from the unit when the implant moves along the longitudinal axis of figure-8 coil (x-axis) on xz-plane. Here, xyz and x’y’z’ are the coordinate systems of the unit (known) and the implant (unknown), respectively. That is, in actual localization, by monitoring the voltage during moving the unit linearly over the existence possibility area of the implant, it is possible to specify the symmetry plane of the implant (e.g., y’z’-plane which is perpendicular to the moving x-axis when the voltage is at the inflection point). Similarly, as shown in Fig. 1(b), by simply rotating the unit 90° counter-clockwise on xy-plane, and then moving the unit along the longitudinal axis of figure-8 coil, it is possible to specify the symmetry plane x’z’. Finally, for specifying the symmetry plane x’y’, rotating the unit 90° around y-axis counter-clockwise, and the resulting three symmetry plane intersect at the position of the implant.&lt;br/&gt;Next, we verified the validity of the proposed method using the constructed experimental setup as shown in Fig. 2(a). The strength and frequency of magnetic field applied at the center of drive coil (pancake coil, 40 turns, internal diameter 28.1 mm, external diameter 113.5 mm) were 2.88 mT and 2 kHz, respectively. The induced voltage in figure-8 coil (each loop 1 turn, diameter 10 mm) without the sample was subtracted from the induced voltage with the sample. The sample (1 g of magnetic particles dispersed in 1mL of deionized water) used in the following experiments was our previously-developed micromagnetic particles with a Curie point of 45°C [1].&lt;br/&gt;As shown in Fig. 2(b), when the sample moved manually along the x-axis (z = 10 mm; y= 0 mm or -10mm), the average voltage (all experiments were repeated three times) changed depending on the position of sample, and the voltage was at the inflection point when the sample was directly above the intersection of two loops of figure-8 coil. In Fig. 2(c), when the sample moved along the y-axis (z = 10 mm; x= 0 mm), the voltage (almost 0 V) was extremely small compared to the one when the sample moved along the x-axis around the intersection of two loops of figure-8 coil (z = 10 mm; y= 0 mm). This means that the symmetric plane of the implant can be specified with high accuracy by using figure-8 coil. As shown in Fig. 2(d), with our currently constructed experimental setup, it is possible to detect the sample up to 5 cm from the unit. Future studies are needed to investigate the optimal design of drive coil and figure-8 coil to extend the detectable distance for deeper tumor.

**References:**

[1] L. Tonthat, Y. Yamamoto, F. Aki et al, IEEE Trans. Magn., Vol. 54(7), p. 5400506 (2018)&lt;br/&gt;[2] F. Aki, L. Tonthat, H. Saito et al, Electron. Comm. Jpn., Vol. 101 (6), p. 58–66 (2018)&lt;br/&gt;[3] L. Tonthat, K. Mitobe, S. Yabukami, IEEE Trans. Magn., Vol. 57(2), doi: 10.1109/TMAG.2020.3008490 (early access).

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/eb9288a7a69c6bd175f360d34d4b58ed.jpg)
Fig. 1  The conceptual diagram of the proposed localization method for magnetic particles using a figure-8 coil. (a) The voltage changes depending on the relative position of the implant as seen from the unit when the implant moves along x-axis on xz-plane. (b) By simply rotating the unit 90° counter-clockwise on xy-plane, it is possible to specify the symmetry plane x’z’ and rotating the unit 90° around y-axis counter-clockwise for specifying the symmetry plane x’y’.

![](https://assets.underline.io/uploads/markdown_image/1/image/42c1dd14e9741c112ca8fe9f33ccc20e.png)
Fig. 2  (a) The constructed experimental setup and (b), (c), (d) the relationship between the average voltage (n=3) and the position of the sample when the sample moved manually above the fixed unit.

INTERMAG 2021

 A Simple Localization Method of Magnetic Particles for Hyperthermia Therapy using Figure-8 Coil

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Magnetic hyperthermia is a promising cancer therapy gaining great interest in recent years using heat generation ascribed to magnetic particles subjected to a high-frequency magnetic field. To make this therapy feasible in clinical settings, in addition to magnetic particles, important elements include techniques for detecting their position and temperature in a tumor region in determining the effectiveness of therapeutic heating. In previous studies, we developed thermosensitive magnetic micro/nanoparticles with high heating efficiency and considerable permeability change around the therapeutic temperature range of 42-45°C [1]. We also proposed a wireless temperature measurement method for monitoring the tumor temperature using pickup coils to detect the temperature-dependent stray magnetic field caused by the magnetic particles [2]. A challenge remains when the particles injected into the tumor region deviate from the longitudinal axis of drive coil (heating coil). As a result, the heating generation of the particles and the accuracy of the proposed thermometry are decreased. To solve this problem, by taking the advantage of the fact that the induced voltage in pickup coil depends on the position of the particles, we developed a localization system using three pickup coils symmetrically installed inside drive coil. The position of particles could be automatically localized by using a robot arm to operate the unit composed of drive coil and pickup coils in two modes of coarse adjustment (rotary scanning) and fine adjustment (linear scanning) [3]. In this study, we propose a simpler, faster and more cost-effective localization method by using only figure-8 coil instead of three pickup coils (loop type coils) and operating the unit of drive coil and figure-8 coil (hereafter the unit) only in linear scanning mode. Fig. 1 shows the conceptual diagram of the proposed localization method using figure-8 coil. Figure-8 coil consists of two reverse-winding loop coils and is adjusted so that its initial bias (the voltage without sample) close to 0 V and fixed to the center of drive coil. As a result, the induced voltage changes depending on the relative position of the implant (i.e., magnetic particles) as seen from the unit when the implant moves along the longitudinal axis of figure-8 coil (x-axis) on xz-plane. Here, xyz and x’y’z’ are the coordinate systems of the unit (known) and the implant (unknown), respectively. That is, in actual localization, by monitoring the voltage during moving the unit linearly over the existence possibility area of the implant, it is possible to specify the symmetry plane of the implant (e.g., y’z’-plane which is perpendicular to the moving x-axis when the voltage is at the inflection point). Similarly, as shown in Fig. 1(b), by simply rotating the unit 90° counter-clockwise on xy-plane, and then moving the unit along the longitudinal axis of figure-8 coil, it is possible to specify the symmetry plane x’z’. Finally, for specifying the symmetry plane x’y’, rotating the unit 90° around y-axis counter-clockwise, and the resulting three symmetry plane intersect at the position of the implant. Next, we verified the validity of the proposed method using the constructed experimental setup as shown in Fig. 2(a). The strength and frequency of magnetic field applied at the center of drive coil (pancake coil, 40 turns, internal diameter 28.1 mm, external diameter 113.5 mm) were 2.88 mT and 2 kHz, respectively. The induced voltage in figure-8 coil (each loop 1 turn, diameter 10 mm) without the sample was subtracted from the induced voltage with the sample. The sample (1 g of magnetic particles dispersed in 1mL of deionized water) used in the following experiments was our previously-developed micromagnetic particles with a Curie point of 45°C [1]. As shown in Fig. 2(b), when the sample moved manually along the x-axis (z = 10 mm; y= 0 mm or -10mm), the average voltage (all experiments were repeated three times) changed depending on the position of sample, and the voltage was at the inflection point when the sample was directly above the intersection of two loops of figure-8 coil. In Fig. 2(c), when the sample moved along the y-axis (z = 10 mm; x= 0 mm), the voltage (almost 0 V) was extremely small compared to the one when the sample moved along the x-axis around the intersection of two loops of figure-8 coil (z = 10 mm; y= 0 mm). This means that the symmetric plane of the implant can be specified with high accuracy by using figure-8 coil. As shown in Fig. 2(d), with our currently constructed experimental setup, it is possible to detect the sample up to 5 cm from the unit. Future studies are needed to investigate the optimal design of drive coil and figure-8 coil to extend the detectable distance for deeper tumor.

**References:**

[1] L. Tonthat, Y. Yamamoto, F. Aki et al, IEEE Trans. Magn., Vol. 54(7), p. 5400506 (2018) [2] F. Aki, L. Tonthat, H. Saito et al, Electron. Comm. Jpn., Vol. 101 (6), p. 58–66 (2018) [3] L. Tonthat, K. Mitobe, S. Yabukami, IEEE Trans. Magn., Vol. 57(2), doi: 10.1109/TMAG.2020.3008490 (early access).

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/eb9288a7a69c6bd175f360d34d4b58ed.jpg)
Fig. 1 The conceptual diagram of the proposed localization method for magnetic particles using a figure-8 coil. (a) The voltage changes depending on the relative position of the implant as seen from the unit when the implant moves along x-axis on xz-plane. (b) By simply rotating the unit 90° counter-clockwise on xy-plane, it is possible to specify the symmetry plane x’z’ and rotating the unit 90° around y-axis counter-clockwise for specifying the symmetry plane x’y’.

![](https://assets.underline.io/uploads/markdown_image/1/image/42c1dd14e9741c112ca8fe9f33ccc20e.png)
Fig. 2 (a) The constructed experimental setup and (b), (c), (d) the relationship between the average voltage (n=3) and the position of the sample when the sample moved manually above the fixed unit.

poster

Transcranial Magnetic Stimulation (TMS) is a non-invasive technique of modulating the neurons of the brain. TMS is a tool in cognitive neuroscience for studying neuronal connectivity and brain mapping. It is also approved for the treatment of neuropsychiatric disorders [1]. It functions on the principle of electromagnetic induction in inducing electric field in the neurons of the brain. With TMS studies, accurate estimation of the induced E-Field is usually necessary. This, however, requires a lot of processes, which include the three-dimensional head model generation from magnetic resonance imaging (MRI) scans using the SimNIBS software [2] and finite element analysis to calculate the induced E-Field (Fig. 1). These processes are time-consuming and computationally expensive. Additionally, with each head model's uniqueness, outcomes cannot be generalized across a particular population as the intensity of E-Field is subject-specific [3], [4]. A few studies have explored the use of artificial neural networks to estimate the electric field from T1 MR images [5], [6], [7]; however, none has considered the effect of different coil geometry in their estimation. In this research, the authors develop a deep neural network (DNN) model to determine the induced E-Field directly from the patient’s MRI scan and across different coil types. We trained neural network models from realistically anatomic head models’ and across 16 different coil type to predict the induced electric field in the brain and scalp (E-Max brain and scalp) and the volume of stimulation of the brain and scalp (V-half brain and scalp) from T1 MR images and specific coil type. Using Deep Neural Network (DNN), the processing and estimation time for the induced E-Field is measurably reduced, which is helpful both to clinicians and researchers as the need to create subject-specific anatomical head structures is eliminated. This will also be beneficial to the TMS patients as there will be no need for additional stimulation sessions with the different coil types as the DNN network is able to predict the outcome from each coil type. The other advantage of the DNN model is that the E-Field from the different coil types can be compared simultaneously. Preliminary results for the E-Max brain report a correlation coefficient of 0.93 for the training dataset and 0.88 on the test dataset, as shown in the parity plot in Fig. 2.

**References:**

[1] L. J. Gomez, S. M. Goetz, and A. V. Peterchev, “Design of transcranial magnetic stimulation coils with optimal trade-off between depth, focality, and energy,” J. Neural Eng., vol. 15, no. 4, p. 046033, Aug. 2018, doi: 10.1088/1741-2552/aac967. [2] M. Windhoff, A. Opitz, and A. Thielscher, “Electric field calculations in brain stimulation based on finite elements: An optimized processing pipeline for the generation and usage of accurate individual head models,” Human Brain Mapping, vol. 34, no. 4, pp. 923–935, 2013, doi: 10.1002/hbm.21479. [3] F. Syeda, H. Magsood, E. G. Lee, A. A. El-Gendy, D. C. Jiles, and R. L. Hadimani, “Effect of anatomical variability in brain on transcranial magnetic stimulation treatment,” AIP Advances, vol. 7, no. 5, p. 056711, Jan. 2017, doi: 10.1063/1.4974981. [4] P. C. Miranda, L. Correia, R. Salvador, and P. J. Basser, “The Role of Tissue Heterogeneity in Neural Stimulation by Applied Electric Fields,” in 2007 29th Annual International Conference of the IEEE Engineering in Medicine and Biology Society, Aug. 2007, pp. 1715–1718, doi: 10.1109/IEMBS.2007.4352640. [5] J. A. Livezey, K. E. Bouchard, and E. F. Chang, “Deep learning as a tool for neural data analysis: Speech classification and cross-frequency coupling in human sensorimotor cortex,” PLoS Comput Biol, vol. 15, no. 9, p. e1007091, Sep. 2019, doi: 10.1371/journal.pcbi.1007091. [6] Tatsuya Yokota, Toyohiro Maki, Tatsuya Nagata, Takenobu Murakami, Yoshikazu Ugawa, Ilkka Laakso, Akimasa Hirata, Hidekata Hontani, Real-time estimation of electric fields induced by transcranial magnetic stimulation with deep neural networks, Brain Stimulation, Volume 12, Issue 6, 2019, Pages 1500-1507, ISSN 1935-861X, doi: ttps://doi.org/10.1016/j.brs.2019.06.015. [7] A. S. Lundervold and A. Lundervold, “An overview of deep learning in medical imaging focusing on MRI,” Zeitschrift für Medizinische Physik, vol. 29, no. 2, pp. 102–127, May 2019, doi: 10.1016/j.zemedi.2018.11.002.

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/10a9e1ff81dca3187fa7c5110e4343b1.jpg)
Fig. 1. Conventional Process for Induced Electric Field Estimation during TMS

![](https://assets.underline.io/uploads/markdown_image/1/image/f95e4f0342ee55c312e1ab3dc38bfa6d.jpg)
Fig. 2. Parity Plots of the E-Max Brain for the Test and Training dataset.

 Neural Network Model for Estimation of the Induced Electric Field during Transcranial Magnetic Stimulation

In the medical world, today’s challenge is to kill and control tumors at a cellular level, allowing a more localized (confined) treatment for the eradication of malign cells. Due to this localized treatment, secondary harmful effects induced on healthy cells would be significantly reduced. In fact, different approaches have been used to apply hyperthermia in tumor regions, but with the harmful side effects it defeats the purpose. Hence, the author selected to engineer biocompatible magnetic nanoparticles, which are much smaller than the pore size of the cells, for hyperthermia and drug delivery applications. In order to achieve this goal, one has to produce biocompatible magnetic nanoparticles that are controlled regarding size and shape for hyperthermia treatment. Furthermore, FeCo nanoparticles have superior magnetic properties (230-240 emu/g) than iron oxide nanoparticles (60-80 emu/g). However, this idea is not yet fully explored because of the commercial unavailability of nanoparticles that are size and shape controlled; toxicity also plays a key role in regards to the commercial unavailability. Here, the author presents a cheap, economical, and environmentally friendly approach to prepare the surfactant coated FeCo nanoparticles in the size range of 15-25 nm with a decent amount of dispersity. The FeCo nanoparticles are prepared by using the "mixed glycol process,” with which Fe and Co metal salts are used as precursors. Mixed polyols can increase the viscosity of the solution and can reduce the migration rate of metal ions in the solution. First, the ethylene glycol, trimethylene glycol, and polyrhrtlene glycol were used in different ratios, and then the metal precursors were added as well. To increase the reaction rate, sodium hydroxide(NaOH) was used. Further, the surfactants (i) polyvinylpyrrolidone (PVP) and (ii) a mixture of oleylamine and oleic acid were used to control the particle size, shape and dispersity. The experiment was carried out at different temperatures between 70 and 110° C. Once the temperature was raised to 60° C, the solution started to change from a pink to green color. At 110° C, the solution also became a dark green color as well. After an hour, the reaction was completed, the mantle heater was switched off, and left for cooling. After cooling, the reaction flask was taken out and the viscous solution was poured into four centrifuge tubes; this is to centrifuge the samples at 2500 rotation per minute using a small lab scale centrifuge set up for 10 minutes. The following are the observations and data seen in the experiment. The particles that were prepared without surfactants had agglomerated (clustered) particles with an irregular shape. However, the PVP and oleylamine/oleic acid coated samples had a cubic shape for the particles, with average particle sizes in the range of 15 and 20-25 nm respectively. The X-ray diffraction analysis characterized that the structure was the single phase body centered cubic structure. The magnetization value is about 210 emu/g, which is higher than the iron oxide nanoparticles. The nanoparticles with surfactants show well dispersion in ethanol. Since these nanoparticle sizes are smaller than the size of the cell, it is possible to use these particles for hyperthermia and drug delivery applications.

**References:**



**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/f04444c872f5213610d3ffe6258fb41c.jpg)
Fig. 1. SEM image of the low temperature synthesized FeCo nanoparticles through mixed glycol process

 Low Temperature Synthesis of Functionalized Magnetic Nanoparticles Engineering for Cancer Theranostics

Magnetic particles (MPs) found their usefulness in different cancer treatment applications, such as magnetic hyperthermia, magnetic controlled delivery and release of antitumoral drugs at the targeted site of a tumor, or through magneto-mechanical actuation. Recently, we have introduced a new type of magnetic particles (MPs), prepared by wet milling of superferromagnetic Fe-Cr-Nb-B precursor glassy ribbons, for cancer treatment by magneto-mechanical actuation in low magnetic fields (1÷20 Oe) [1]. The rectangular shapes of newly developed MPs and the superferromagnetism of the glassy alloys of which they are made induce important magnetic shape anisotropies which, in association with a large saturation magnetization, generate an improved torque in a rotating magnetic field, producing important damages on the cellular viability of MG-63 human osteosarcoma (HOS) cells, a form of bone cancer. In this work we studied the possibility of transporting Fe-Cr-Nb-B MPs to areas with cancer cells (HOS), using adipose derived stem cells (ADSC) as carriers, considering their tumor-targeting capacity [2], and the effect of magneto-mechanical actuation of the MPs on HOS viability. Fe-Cr-Nb-B MPs have dimensions between 5 and 100 nm. The Fe-Cr-Nb-B MPs prepared by wet milling were used to obtain a ferrofluid (FF) that was added to the cell culture media before adding it in the cell cultures. Afterwards, HOS and ADSC cells were incubated for 24h with the MPs, and cellular viability assay was performed with MTT (3-(4,5-dimethylthiazolyl-2)-2,5-dyphenytetrazolium bromide). We performed the tests by measuring the cell proliferation rate for 1 mg/ml of particles on both cell types (HOS and ADSC). No cytotoxic effect was observed, the cell proliferation being improved in cell culture incubated with particles. At the same time we performed two tests to confirm the presence of MPs inside both cell types. This was accomplished using transmission electron microscopy (TEM) and colorimetric Ferrozine assay, the latter evaluating the quantity of iron inside the cell. MPs upload by HOS was confirmed using TEM, the particles can be seen both in the extracellular and intracellular medium, being incorporated as single or aggregated particles. Iron quantity per cell was evaluated 48h after adding the MPs to the culture media. We calculated the iron content per cell after subtracting the normal quantity of iron in similar culture condition. ADSC, being larger cells, incorporated larger quantities of MPs (1.12 ng per cell) than HOS (0.31 ng per cell). ADSC cellular migration, both loaded and non-loaded with MPs, was recorded by time-lapse imaging performed using an Evos optical microscope while cells were inside a self-made mini-incubator. Before evaluation, the two cell population (ADSC-MPs and tumor cells) were separated by a free space of about 0.5 mm made mechanically in the co-culture by using a pipette tip (“wound healing” model). The traveled distance of loaded ADSCs towards the zone of the tumor cells was the highest among the evaluated ones. The recorded films showed that MPs-loaded ADSC were able to easily target osteosarcoma cells, moving at twice the speed of the MPs-free ADSC (control sample) in similar cultures (Fig. 1). Magneto-mechanical actuation was performed in a system consisting of four coils placed in cross, which can produce low magnetic field (1÷20 Oe).The destruction of HOS co-incubated with MPs-loaded ADSC was performed through controlled magneto-mechanical actuation (MMA). Magneto-mechanical actuation led to the destruction of ADSC and to the release of MPs on HOS cells, the latter incorporating the released MPs (Fig. 2). By magneto-mechanic actuation, the HOS viability dropped down to 21÷22%. The rotational motion of the irregular shaped MPs produced in the rotating magnetic led to irreversible damages on the cells membranes and organelles, and, subsequently, the cells death. We have successfully showed that Fe-Cr-Nb-B magnetic particles (MPs) are highly biocompatible in vitro. MPs were easily incorporated by ADSC and HOS cells, results confirmed through TEM imaging and Ferrozine assay. MPs loaded ADSCs display increased motility towards tumor cells, compared with their unloaded counterparts. The scope of the study, the destruction of HOS through MMA of the MPs released from ADSC was confirmed through MTT assay. MMA led to the release of the MPs towards tumor cells, the latter being destroyed in high proportion (about 80%) by immediate applying of the MMA. Consequently this method could be successfully implemented in the treatment of bone cancer. Acknowledgements - Work supported by the Romanian Executive Agency for Higher Education, Research, Development and Innovation Funding (UEFISCDI) under contract no. 502PED/2020 (project PN-III-P2-2.1-PED-2019-3442/OPTIMAG).

**References:**

[1] H. Chiriac et al., Scientific Report, Vol. 8, Article Number 11538 (2018). [2] M.G. Scioli et al., International Journal of Molecular Sciences, Vol. 20, p. 3296 (2019).

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/31002254d1c363d3984666805bd7aadc.jpg)
Fig. 1. In vitro cell migration of ADSC loaded and non-loaded, respectively, with MPs towards tumor cells.

![](https://assets.underline.io/uploads/markdown_image/1/image/85d678c0a31d66d5a4895cd0faaa15d6.jpg)
Fig. 2. Cell viability of HOS and ADSC cells, controls and cells with MPs after MM actuation.

 Cancer cells death induced by magneto-mechanical actuation of Fe-Cr-Nb-B magnetic particles carried by Stem cells to the cancer cells area

Introduction: Red blood cell(RBC), which passes through capillaries and exchanges O2 and CO2 with surrounding tissues, plays a very important in the blood circulation in the human body. Imbalance in blood circulation could lead to various cardiovascular related diseases such as high blood pressure, arteriosclerosis, diabetes, and stroke.[1] This is closely related to RBC aggregation. Negatively charged RBC membrane arranges a compact layer of positive ions around the membrane by electrical attraction with cations in plasma. The rearranged ions layer in plasma forms zeta potential with compact layer. When the electrostatic interaction between RBCs, i.e., zeta potential increases, RBC aggregation is improved. It is known that zeta potential depends on electronegativity of the RBC membrane, the dielectric constant of the medium and ionic strength [2]. Since pulsed magnetic field (PMF) influences human physiology, many clinical studies demonstrated the effect of PMF for wound healing, reducing pain, increasing angiogenesis, as well as improving blood circulation.[3-5] In this study, we have tried to elucidate the effect of PMF on RBC aggregation by indirectly observing zeta potential using negatively-charged magnetic beads(MB). Our hypothesis is that the increase of negative charges on RBC membrane results in the increase of cations in the compact layer, thus the increase of the zeta potential. As a result, the number of MB attached to the RBC increases, as the binding force between MB and compact layer increases. Experimental Method: Blood needed for experiment was obtained from Korean Red Cross after reviewing IRB exemption in Sangji University Bioethics Committee. Whole blood of 6 ml was centrifuged at 3000 rpm for 5 min in 4 oC for separating RBCs and plasma. MB used this experiment has diameter 1 μm with OH- groups (SiMAG-Silanol, Chemicell). To clearly see the MB attached to RBCs and, in addition, to avoid blood clots caused by proteins in plasma, RBC suspension of hematocrit (Ht) of 10% was prepared using PBS(phosphate buffer saline). The amount of MB added to RBC suspension was 0.5 mg. RBCs were also oxidative stressed with tBHP(Tert-Butyl hydroperoxide) in 0.4 mM at 37 oCincubator for 30 min, in order to confirm the change of RBCs membrane charge due to oxidative stress under PMF stimulus. In oxidative stressed RBCs, hemoglobin(Fe2+) is altered to methemoglobin(Fe3+) because of loss an electron, thereby, electronegativity of RBC membrane will be changed.[6] Each sample was exposed to PMF for 3 min. Our PMF stimulator system consisted of single layered coil of 10 turns with an elliptical shape of 12.0 x 4.5 cm. The maximum intensity is 0.27 T at a transition time of 0.102 ms, with pulse intervals of 1 Hz. The numbers of RBC attached MB were compared with control group and PMF treatment group. Only RBCs with a distance of about 1 μm between the MB and the surface of RBC were counted using hemocytometer. Results and Discussion: Fig.1. shows the numbers of RBCs according to the number of attached MB in the control and PMF treatment group. The numbers of RBCs attached MB increased after PMF treatment. The more the number of MBs attached to RBCs, the thicker the compact layer surrounding RBC membranes. This phenomenon might be inferred due to an increase of zeta potential, thereby, increase of negative charges on the RBC membrane. Thus, it can be interpreted that MB having hydroxide OH- group are more attached to RBCs via binding force between MB and compact layer. Also the numbers of MB in the group treated tBHP and the group exposed to PMF after tBHP treatment (tBHP+PMF) were compared in Fig.2. Although the number of MB attached to RBCs was 1~2 in tBHP group, after PMF stimulus, i.e. in tBHP+PMF group it appears the number of MB attached to RBCs increased up to 3~4. This might be explained that PMF affects the increase of RBC membrane charge on not only normal RBCs but also oxidative stressed RBCs. Therefore, it can be deduced that high instantaneous flux change of PMF improves the aggregation of RBCs by increasing the zeta potential, due to the increase of the membrane charge.

**References:**

[1] H. L. Goldsmith, G. R. Cokelet, and P. Gaehtgens, Am. J. Physiol. 257, H1005 (1989) [2] H. P. Fernandes, C. L. Cesar and M. D. L. Barjas-Castro, Rev. Bras. Hematol. Hemoter. 33(4), 297(2011) [3] N. M. Shupak, J. C. McKay, W. R. Nielson, G. B. Rollman, F. S. Prato and A. W. Thomas, Pain Res Manag. 11(2), 85(2006) [4] L. X. Alvarez, J. McCue, N. K. Lam, G. Askin and P. R. Fox, J. Am. Anim. Hosp. Assoc. 55, 83(2019) [5] J. W. Mok, S. H. Han and H. S. Lee, AIP Advances 10, 015005(2020) [6] A. Knishita, Y. Nakayama, T. Kitayama and M. Tomita, The FEBS Journal 274(6), 1449(2007)

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/449981197245053b780d8c1e8a7f02ec.jpg)
The numbers of RBCs according to the number of attached MB in the control and PMF treatment group

![](https://assets.underline.io/uploads/markdown_image/1/image/49f2876eafa8f3d1fcb1c73de892a6f0.jpg)
The numbers of RBCs according to the number of attached MB in the tBHP treated group and tBHP+PMF treatment group

 Observation of Red Blood Cell Membrane Charge using Magnetic Beads under Pulsed Magnetic Field

Cancer is the leading cause of death in Europe after cardiovascular disease, accounting for about 20% of deaths on that continent [1]. In recent years, hyperthermia emerged as a promising treatment approach in oncology consisting in raising the temperature of cancer cells to 40-45°C to reach apoptosis i.e., programmed cell death. Differences in the vasculature of tumors and healthy tissues leads to different heat-induced damage responses. At temperatures up to 45°, the vasculature of healthy tissues allows an effective heating dissipation, while the tumor’s lumpy and disorganized ones makes them less efficient in heat dissipation. Therefore, for the same thermal dose, the healthy tissue adjacent to the tumor can remain at lower temperatures than the tumor itself [2]. One way to reach local and controlled hyperthermia is via functionalizable nanostructures that are activated by external stimuli, such as magnetic fields or electromagnetic radiation. In this context, gold nanostructures (Au-NS) have been the subject of much attention in the academic and clinical environment, due to their biocompatibility and high absorption of electromagnetic radiation in the near-infrared range caused by its surface plasmon resonance [3]. In parallel, magnetic nanostructures based on iron (Fe-NS) have also been gaining attention, since they can combine the diagnostic (as contrast agents for magnetic resonance imaging) and therapy (magnetic hyperthermia) properties [4]. Although the Fe-oxide NS are the most reported in the literature, Fe-NS are a promising alternative. The high magnetic moment of Fe-NS can improve the heat dissipation phenomena produced by the magnetic hysteresis, due to the irreversible magnetization/ demagnetization processes induced by an applied alternating magnetic field [5]. The main goal of this work is to combine the “best of both worlds” producing multifunctional Iron-Gold nanostructures (Fe-Au-NS) with high heating performance when stimulated with radiation (500-1000nm) and with alternating magnetic fields, for applications in controlled and localized hyperthermia. Fe-Au-NS were produced by two methods: 1) through the ablation of Iron and Gold targets with a femtosecond pulsed laser in liquids (such as ethanol) and 2) by electrodeposition in self-organized alumina matrices. The first technique is particularly interesting for the development of NS-Fe-Au with complex structures such as the core-shell structure [4]. In turn, the second technique allows a controlled growth of nanowires, nanotubes, and more complex morphologies such as segmented Fe-Au-Fe nanowires [6]. The Fe-Au-NS have been fully characterized with scanning electron microscopy (SEM), X-ray diffraction (XRD) and superconducting quantum interference device (SQUID) techniques. The morphological analysis showed a narrow distribution in diameter and length of the obtained structures with improved robustness and high yield, making these techniques versatile approaches strongly compatible with large scale production. Finally, we show the possibility to tune accurately the size of the nanostructures and consequently provide an easy control over the magnetic properties of these nanostructures ultimately enabling reaching the superparamagnetic regime.

**References:**

[1] World Health Organization, Regional office for Europe: http://www.euro.who.int/en/health-topics/noncommunicable-diseases/cancer/data-and-statistics [2] A. Espinosa, et al, “Magnetic (Hyper)Thermia or Photothermia? Progressive Comparison of Iron Oxide and Gold Nanoparticles Heating in Water, in Cells, and In Vivo”, Advanced Functional Materials, 28, 37 (2018) [3] Huang, et al “Cancer Cell Imaging and Photothermal Therapy in the Near-Infrared Region by Using Gold Nanorods”, Journal of American Chemical Society, 128, 2115-2120 (2004) [4] P. Wagener, et al “Solvent-surface interactions control the phase structure in laser generated iron-gold core-shell nanoparticles”, Scientific Reports, 6, 23352 (2016) [5] L. Peixoto, et al, "Magnetic nanostructures for emerging biomedical applications", Applied Physics Reviews, 7 (2020) [6] C. Sousa, et al, “Nanoporous alumina as templates for multifunctional applications” Applied Physics Reviews, 1, 031102 (2014)

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/3c5f2bfac9400b165970e45f4430476c.jpg)
Fig 1. Fe-NPs synthetized via ablation of an Iron target by femtosecond laser in ethanol, in the presence of magnetic field applied post LAL

![](https://assets.underline.io/uploads/markdown_image/1/image/54f1dde19192a42c34f083352fd223ff.jpg)
Fig 2. Laser Ablation in Liquids Set-up

 Multifunctional Fe-Au nanostructures for biomedical applications

Ferrite nanoparticles (NPs) have demonstrated considerable potential for use in biomedical applications such as magnetic resonance imaging, magnetic particle imaging, and magnetic hyperthermia treatment (MHT)[1,2]. In the medical application of ferrite NPs, biocompatibility must be ensured. To enhance the biocompatibility, we modified the NPs by using polyethylene glycol (PEG), which is highly hydrophilic and biocompatible. Furthermore, in the heat dissipation of the magnetic NPs, caused by the relaxation mechanism, the imaginary part of the magnetic susceptibility χ’’ at the room temperature is a key parameter. Moreover, M–Zn ferrite NPs (M = Ni, Co, Mn, Fe) are multielement ferrites, whose magnetization can be increased by doping with nonmagnetic Zn ions . The purpose of this study was to develop a sample with improved biocompatibility by PEGylation, to clarify the change in magnetization characteristics due to the Zn doping effect, and to clarify the heat generation mechanism of the sample for MHT application. This study was focused on Ni-Zn ferrite NPs , and NPs with different contents of Zn ions were prepared. Ni1-xZnxFe2O4 (x = 0, 0.2, 0.5, 0.8) NPs coated with PEG were synthesized using our original method[2]. The crystal structures of these particles were examined through X-ray diffraction and X-ray absorption spectroscopy. The particle diameters were controlled to be between 5–20 nm by controlling the amount of water, and the samples with different Zn doping amounts were adjusted to have a particle size of 10 nm. PEGylation in the particles was confirmed through the Fourier-transform infrared spectrometry technique(Fig. 1) and by using a thermogravimeter-differential thermal analyzer. The magnetization curves of all the samples were obtained at 300 K in a ±10 kOe field. All the samples exhibited superparamagnetic behavior with no coercive force, and the magnetization values ranged between 39 and 43 emu g-1 with only a slight afference. The field-cooled/zero-field-cooled measurement showed that, in terms of the temperature dependence of the magnetization, the blocking temperature TB ranged between 105 and 150 K(Fig. 2). χ’’ exhibited a peak value in the x = 0.5 sample at room temperature, and the most effective temperature rise was expected in this composition because the heat dissipation of superparamagnetic particles depends on the value of χ’’ . Moreover, we measured the increase in the temperature in the AC magnetic field of 15 kHz, 130 Oe for all the samples. The temperature of the samples was measured using an optical fiber thermometer. The value of heat dissipation could be attributed to the Néel relaxation system. To clarify if the heat generation mechanism pertained to the Néel relaxation or Brownian relaxation, the heating effect in silicone oil and water was evaluated at a frequency of 210 kHz. It was noted that the Néel relaxation and Brownian relaxation were dominant for particle sizes of approximately 10 nm and more than 20 nm, respectively.Thus, we successfully developed highly biocompatible magnetic NPs and experimentally clarified the contribution of the Néel and Brownian relaxation depending on the particle size.

**References:**

[1]Shigeoka, D., Yamazaki, T., Ishikawa, T., Miike, K., Fujiwara, K., Ide, T., Oshima, A., Hashimoto, T., Aihara, D., Kanda, K., Usui, A., Hosokai, Y., Saito, H. and Ichiyanagi, Y., 2018. Functionalization and magnetic relaxation of ferrite nanoparticles for theranostics. IEEE Transactions on Magnetics, 54(11), p.6100707. [2]Oshima, A., Kanda, K., Fujiwara, K., Ide, T., Takano-Kasuya, M., and Ichiyanagi, Y., 2020. PEGylation of Co–Zn Ferrite Nanoparticles for Theranostics. Journal of Nanoscience and Nanotechnology, 20(12). Pp.7255–7262.

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/fc45e4b9b7cfba5f388926fbeb91745c.jpg)
FT-IR spectra for PEG400 reagent and NiFe2O4 NPs coating with PEG400.

![](https://assets.underline.io/uploads/markdown_image/1/image/10f142a824d277e48e34d286e45fb494.jpg)
Temperature dependence of magnetization of both field-cooled (FC) and zero-field-cooled (ZFC) magnetization under 100 Oe field for x =0, 0.5, 0.8 samples.

 PEGylation of Ni1-xZnxFe2O4 Nanoparticles with Heat Dissipation Based on Néel and Brownian Relaxation

Background Over a long period of time, the Mexican axolotl (Ambystoma mexicanum) has been used as an invaluable model for regenerative capacity [1]–[3]. On the other hand, there are no reported studies of the observation of the capacity under exposure to extremely low frequency (ELF) magnetic fields. Our previous report already provided a precious finding about the influences of a gradient/ELF magnetic field on axolotl metamorphosis induced forcibly with a thyroid hormone [4]. In this research, we investigated the ELF magnetic effects on axolotl-tailfin regeneration, as well as the survival rates of the axolotls influenced by the presence of thyroxine (T4). Experimental Twenty axolotls (about 120 mm) were bred under the same condition as group feeding. Before performing this experiment, axolotls were individually kept in 0.85-L square boxes containing a de-chlorinated water (0.7 L) without aeration under illumination of 250 μEm-2s-1on a 12 : 12 h L : D photocycle. Employing an original water-renewing system [4], the water temperature was strictly controlled at 24°C. Exposures of axolotls to ELF fields (10 mT at 10–100 Hz) were performed using an air-cored coil system. The regenerations of the axolotl-tailfinswere monitored every day, and the survival rates were evaluated minutely. Results As it turned out, we found that the initiation timings of ELF field exposure did affect the rapidity of the axolotl-tailfin regeneration, as well as the metamorphic rapidity of the T4-administrated axolotls (Fig. 1). In addition, there were a few distinct patterns that revealed a serious indigestion in each of the salamanded bodies (Fig. 2). Our data could prove to be helpful for future reference with respect to a magnetic control study of tissue-regeneration.

**References:**

1) T. Demircan et al., Proteomics 17, 1600338 (2017). 2) L. Alibardi, J. Exp. Zool. A, Ecol. Genet. Physiol. 323, 757 (2015). 3) K. Mitogawa et al., PLoS ONE 10, e0133375 (2015). 4) H. Nakagawa and M. Ohuchi, IEEE Trans. Magn. 54, 5000405 (2018).

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/4a942854fcb78e8733ae4f38379fd913.jpg)
Fig. 1. Timeframe of morphological changes in T4-administrated axolotls. (A) Day 0. (B) Day 7. (C) Day 14.

![](https://assets.underline.io/uploads/markdown_image/1/image/2f3a8a5b941fe4281747daae9fb7c49c.jpg)
Fig. 2. Mexican axolotl with serious indigestion.

 Influence of ELF magnetic Field on Thyroxine-Inducing Forced Metamorphosis of Axolotl (Ambystoma mexicanum)

Transcranial Magnetic Stimulation (TMS) is a non-invasive technique that is approved by the U.S. Food and Drug Administration (FDA) for the treatment of certain neurological disorders [1]. When comparing TMS to techniques with similar desired outcomes such as conventional brain surgery, TMS has far fewer risks due to its non-invasive nature. TMS works by inducing an electric field that polarizes or depolarizes brain tissues depending on treatment mode. TMS researchers have proposed different coil designs to achieve higher focality and increased depth of penetration of the induced electric field [2]. The Quadruple Butterfly Coil (QBC) is a novel coil design with increased focality when compared to the commercially available Figure-of-eight (FOE) coil [3]. The human brain reaches 90% of its adult volume at age six and does not reach full maturity until approximately the age of 25, and throughout adolescence, the brain is still maturing [4]. The differences in the brain from adolescents to adults could affect the effectiveness of TMS, mainly because the coils have been designed for use on adult patients [5], [6]. Comparing the efficacy of TMS coils across age groups to determine if the difference in brain volume is enough to warrant specific TMS treatment protocols for adolescents is of the essence in advancing TMS research [7]. In this research, the authors compare the use of the Quadruple Butterfly Coil (QBC) across different age groups to test the effectiveness of the coil on various stages of brain development. Simulations were run using a finite element analysis software, Sim4Life. The models used in this study were sourced from the Computable Virtual Population Models from the IT’IS Foundation. The models have well defined anatomy and were selected to each represent a different age group. The adult models used as the control are 34 years old male (34M) and 26 years old female (26F), as they both fall in the typical age range of models (25 to 35 years) used for TMS studies. The models used to represent the pediatric population are 14 years old male (14M), 11 years old female (11F), 8 years old female (8F), 8 years old male (8M), 6 years old male (6M), and 5 years old female (5F) [8]. Fig. 1 presents the distribution of the induced electric field intensity with depth across the different models. Besides age and brain volume, other factors such as the brain – scalp distance (BSD), cerebrospinal fluid (CSF) thickness, scalp thickness, and skull thickness of the models were also considered in the analysis. This study provides an understanding of the various factors contributing to the response from TMS. The maximum electric field intensity on the scalp (E-Max scalp), maximum electric field intensity on the brain (E-Max brain) and the proportion of the E-Max scalp to E-Max brain are parameters considered in this analysis.

**References:**

[1] M. Hallett, "Transcranial magnetic stimulation and the human brain", Nature, vol. 406, no. 6792, pp. 147-150, 2000. Available: https://www.nature.com/articles/35018000. [2] M. Hallett, "Transcranial Magnetic Stimulation: A Primer", Neuron, vol. 55, no. 2, pp. 187-199, 2007. Available: https://www.sciencedirect.com/science/article/pii/S0896627307004606. [3] P. Rastogi, E. G. Lee, R. L. Hadimani, and D. C. Jiles, “Transcranial Magnetic Stimulation-coil design with improved focality,” AIP Adv., vol. 7, no. 5, p. 056705, May 2017, doi: 10.1063/1.4973604. [4] J. Stiles and T. Jernigan, "The Basics of Brain Development", Neuropsychology Review, vol. 20, no. 4, pp. 327-348, 2010. Available: https://www.ncbi.nlm.nih.gov/pmc/articles/PMC2989000/. [5] L. Crowther, R. Hadimani and D. Jiles, "A numerical dosimetry study for pediatric transcranial magnetic stimulation", 2013 6th International IEEE/EMBS Conference on Neural Engineering (NER), 2013. Available: https://ieeexplore.ieee.org/abstract/document/6695916/authors#authors. [Accessed 10 September 2020]. [6] J. Weissman, C. Epstein and K. Davey, "Magnetic brain stimulation and brain size: relevance to animal studies", Electroencephalography and Clinical Neurophysiology/Evoked Potentials Section, vol. 85, no. 3, pp. 215-219, 1992. Available: https://www.sciencedirect.com/science/article/pii/016855979290135X. [7] M. Garvey and V. Mall, "Transcranial magnetic stimulation in children", Clinical Neurophysiology, vol. 119, no. 5, pp. 973-984, 2008. Available: https://www.sciencedirect.com/science/article/pii/S1388245707007353. [8] M. Gosselin et al., "Development of a new generation of high-resolution anatomical models for medical device evaluation: The Virtual Population 3.0", Physics in Medicine and Biology, vol. 59, no. 18, pp. 5287-5303, 2014. Available: 10.1088/0031-9155/59/18/5287.

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/27f067807ad62659e1f06d16031a63f5.jpg)
Fig. 1. Variation of the Induced Electric Field (V/m) with depth across the different subjects.

 Transcranial Magnetic Stimulation: The Effect of Age and Other Factors on the Intensity of the Quadruple Butterfly Coil

Nuclear magnetic resonance imaging is the most used technique for investigating human tissue. This is a noninvasive technique for the patient, based on the use of an intense magnetic field and radiofrequency pulses applied at certain angles. The application of the radiofrequency pulse has the effect disruption of the equilibrium state of nuclear atomic moments of atoms, induced signals being described by longitudinal relaxation time (T1) and transverse relaxation time (T2). For a better view of tumor tissues, there are used contrast agents; a dose of contrast agent distributed throughout the body can be used to quickly view extensive areas of anatomy. Most of the NMR contrast agents are either complex paramagnetic ions or superparamagnetic loadstone particles (iron oxide). However, the most used agents in NMR investigations are based on Gadolinium chelates. These are less toxic and present positive contrast in the T1 sequence. Superparamagnetic loadstone particles (iron oxide) are at an experimental stage; the sensitivity of these techniques is currently not suitable for clinical applications, new significant relaxation agents targeted to specific tissues or disease processes being required. In this study, the CoZnFe2O4Ag nanoparticle series was dispersed in distilled water, thus forming dilutions, and the same amounts of ferrite were stabilized in agar gels. The obtained dilutions and agar fantoms were scanned with the MRI medical facilities with the intensity of the magnetic field of 1.5 T and 0.4 T respectively. For these, the basic sequences, T1 and T2, were obtained. The contribution of CoZnFe2O4Ag nanoparticles on the contrast of nuclear magnetic resonance images obtained from scanning of two medical MRI installations as compared to images obtained by scanning several commercial contrast agents used in medical imaging. The resulting images were analyzed using a commercial graphics program to highlight the influence of particles on the contrast of the images. The analyzed images indicate that the CoZnFe2O4Ag ferries show positive contrast in the T1 sequence. Small ferrite concentrations (0.05, 0.08, 0.1) have better contrast compared to higher concentrations (0.5, 1). Large concentrations of ferrite have the effect of decreasing magnetization and, implicitly, of lowering the signal (contrast). The CoZnFe2O4Ag series could be used as a possible contrast agent for MRI investigations, only if they are compatible by coating with polymers compatible with the human body. In this respect, severe toxicity tests should be carried out.

**References:**

Rinck, Peter A MAGNETIC RESONANCE IN MEDICINE - A Critical Introduction. 12th edition (2018) Chrysikopoulos, Haris S, Clinical MR Imaging and Physics., (2009) J. Lohrke, T. Frenzel 25 Years of Contrast-Enhanced MRI: Developments, Current Challenges, and Future Perspectives, (2016) Yu-Dong Xiao, Ramchandra Paudel, MRI contrast agents: Classification and application., Vol. 38 (2016) Perry Sprawls, Magnetic Resonance Imaging- Principles, Methods, and Techniques. (2000) Alshowiman, Sami S, MRI Contrast Agents: Developments, Challenges, and Future Trends., Vol. 8 (2018) C. P. Constantin, A. Iordana, O. F. Caltun, CoxZnx-1Fe2O4 nanoparticles ferrite series as magnetic resonance imaging contrast agents. (2012)

**Images:**

 The series study of the CoZnFe2O4Ag in the MRI contrast

Over the past decade, superparamagnetic iron oxide (Fe3O4) nanoparticles have drawn significant attention for different applications such as contrast agents for magnetic resonance imaging (MRI), drug delivery or magnetic hyperthermia(MRI) [1-3]. Due to their negative contrast enhancement effect on T2 and T2* weighted images, they have played a role as contrast agents in MRI [4]. They provide the most change in signal per unit of metal particularly on T2* weighted images as each particle is composed of thousands of iron atoms and defeats the inherent low contrast agent sensitivity of MRI [5]. However, the contrast enhancing effectiveness of a superparamagnetic nanoparticle depends on size of the iron oxide crystal, the composition, the charge, the nature of the coating, the hydrodynamic size of the coated nanoparticle, magnetization and aggregation state in the biological environment as these parameters influence the longitudinal and transverse proton relaxivities [6, 7]. The surface properties of the magnetic NPs play an important role in their efficiency as MRI contrast agents. Therefore, they require a surface coating in order to improve their dispersion and stability in aqueous media. This work was related to the effect of silica shell thickness on proton relaxivities and magnetic properties of superparamagnetic iron oxide nanoparticles. It is possible to predict and finely tune the silica shell thickness by varying the amount of TEOS used to adjust the magnetic properties and proton relaxometric properties. Core-shell structures of the form Fe3O4@mSiO2 with different silica shell thicknesses i.e 45, 63 nm and 73 nm were synthesized by a surfactant template sol-gel method. The silica-coated magnetic iron oxide nanoparticles were found to be spherical in TEM images with a core shell structure with average diameters of 136, 155, and 175 nm for 45, 63 nm and 73 nm shell thickness coatings, respectively. A decrease in Ms value was found with decreasing silica shell thickness. Both r1 and r2 proton relaxivities inversely correlated with the increasing silica shell thickness and size of the particles. The r2 relaxivities were found to be much higher than the r1 relaxivities, which reflects the fact that the T2 proton relaxation is mainly influenced by outer sphere processes. All the Fe3O4@mSiO2 particles are relatively weak T1 contrast agents. While, as T2 contrast agents, they showed higher relaxivity with decreasing silica coating of the magnetic core. It is interesting to note that the changes in r1 and r2 due to silica shell thickness indicate that it is an important factor to take into consideration when designing magnetic nanoparticle probes for molecular imaging applications using MRI. It is crucial to develop an understanding of the structure and properties of the nanoparticle since the different silica shell thicknesses around magnetite have an effect on the proton transverse relaxivity.

**References:**

[1] R. Qiao, C. Yang, M. Gao, Superparamagnetic iron oxide nanoparticles: from preparations to in vivo MRI applications, Journal of Materials Chemistry, , 19 (2009) 6274-6293. [2] M. Mahmoudi, S. Sant, B. Wang, S. Laurent, T. Sen, Superparamagnetic iron oxide nanoparticles (SPIONs): development, surface modification and applications in chemotherapy, Adv Drug Deliv Rev, 63 (2011) 24-46. [3] Q.A. Pankhurst, J. Connolly, S.K. Jones, J. Dobson, Applications of magnetic nanoparticles in biomedicine, Journal of Physics D: Applied Physics, 36 (2003). [4] D.-X. Chen, N. Sun, Z.-J. Huang, C.-M. Cheng, H. Xu, H.-C. Gu, Experimental study on T2 relaxation time of protons in water suspensions of iron-oxide nanoparticles: Effects of polymer coating thickness and over-low 1/T2, Journal of Magnetism and Magnetic Materials, 322 (2010) 548-556. [5] D.W. McRobbie, E.A. Moore, M.J. Graves, M.R. Prince, MRI From picture to proton, Second ed., Cambridge University Press, UK, 2003. [6] S.H. Koenig, K.E. Kellar, Theory of 1/T1 and 1/T2 NMRD profiles of solutions of magnetic nanoparticles, Magn Reson Med, 34 (1995) 227-233. [7] P.H. Qiu, C. Jensen, N. Charity, R. Towner, C.B. Mao, Oil Phase Evaporation-Induced Self-Assembly of Hydrophobic Nanoparticles into Spherical Clusters with Controlled Surface Chemistry in an Oil-in-Water Dispersion and Comparison of Behaviors of Individual and Clustered Iron Oxide Nanoparticles, Journal of the American Chemical Society, 132 (2010) 17724-17732.

**Images:**

![](https://assets.underline.io/uploads/markdown_image/1/image/1d51332af5f385219245ab0c3ce7dfcb.jpg)
Fig. 1. Plot of transverse relaxation rates (1/T2), at different iron oxide concentrations for Fe3O4@mSiO2 with different silica shell thickness, measured at 25 °C and 60 MHz; number in the brackets indicate the thickness of the silica shell.

![](https://assets.underline.io/uploads/markdown_image/1/image/efe5f149c36e526b04ec9386da5543ea.jpg)
Fig. 2. Plot of longitudinal relaxation rates (1/T1) at different iron oxide concentrations for Fe3O4@mSiO2 with different silica shell thickness, measured at 25 °C and 60 MHz; number in the brackets indicate the thickness of the silica shell.

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