Biomed. Eng.-Biomed. Tech. 2015; aop

Silvio Dutz*, Robert Müller, Dietmar Eberbeck, Ingrid Hilger and Matthias Zeisberger

Magnetic nanoparticles adapted for specific biomedical applications DOI 10.1515/bmt-2015-0044 Received March 5, 2015; accepted June 1, 2015

Keywords: drug targeting; fractionation; hyperthermia; magnetic particle imaging; precipitation.

Abstract: Magnetic nanoparticles (MNPs) are used in different biomedical applications, whereby each application requires specific particle properties. To fulfill these requirements, particle properties have to be optimized by means of variation of crystal structure, particle size, and size distribution. To this aim, improved aqueous precipitation procedures for magnetic iron oxide nanoparticle synthesis were developed. One procedure focused on the cyclic growth of MNPs without nucleation of new particle cores during precipitation. The second novel particle type are magnetic multicore nanoparticles, which consist of single cores of approximately 10 nm forming dense clusters in the size range from 40 to 80 nm. Their highest potential features these multicore particles in hyperthermia application. In our in vivo experiments, therapeutically suitable temperatures were reached after 20 s of heating for a particle concentration in the tumor of 1% and field parameters of H = 24 kA/m and f = 410 kHz. This review on our recent investigations for particle optimization demonstrates that tuning magnetic properties of MNPs can be obtained by the alteration of their structure, size, and size distribution. This can be realized by means of control of particle size during synthesis or subsequent size-dependent fractionation. The here-developed particles show high potential for biomedical applications.

Introduction

*Corresponding author: Dr. Silvio Dutz, Institute of Biomedical Engineering and Informatics (BMTI), Technische Universität Ilmenau, Gustav-Kirchhoff-Straße 2, 98693 Ilmenau, Germany, E-mail: [email protected]; and Leibniz Institute of Photonic Technology (IPHT), Department of Nano Biophotonics, 07745 Jena, Germany Robert Müller: Leibniz Institute of Photonic Technology (IPHT), Department of Nano Biophotonics, 07745 Jena, Germany Dietmar Eberbeck: Physikalisch-Technische Bundesanstalt (PTB), 10587 Berlin, Germany Ingrid Hilger: Institute of Diagnostic and Interventional Radiology (IDIR), Jena University Hospital, 07747 Jena, Germany Matthias Zeisberger: Leibniz Institute of Photonic Technology (IPHT), Department of Spectroscopy and Imaging, 07745 Jena, Germany

The application of magnetic nanoparticles (MNPs) in a biomedical context is a rapidly developing field [30, 37]. MNPs suspended in aqueous liquids can be introduced into the blood circuit or the tissue of a patient, allowing for the utilization of magnetic effects such as magnetic losses, magnetic forces, and localized sources of magnetic fields. Magnetic hyperthermia [23, 36] uses the magnetization reversal losses of such particles in an alternating magnetic field (AC) to achieve a local heating of the tissue to treat tumors. An AC field with an amplitude in the order of 10 kA/m and a frequency in the order of 400 kHz has been found to be suitable for hyperthermia [26]. Magnetic drug targeting [43] uses the magnetic force acting on drug-loaded MNP in an external static field gradient to accumulate them in the tissue areas to be treated by the drugs. This method is also promising for tumor therapy, which was investigated in animal ­experiments [1, 2]. In magnetic particle imaging (MPI) [24], the local nonlinear magnetic response of MNP in the tissue to an external field is used to construct a three-dimensional image of the particle distribution, which can be related to the structural features of the tissue. The common basis of all MNP applications mentioned above is that the effects depend on the particle size d as well as on the field parameters (field strength H or AC field amplitude Hac and frequency f). Concerning the magnetic properties, we can distinguish three ranges of the particle size. There is an intermediate size where the particle magnetization consists of one stable single domain (SD) with ferrimagnetic behavior. Smaller particles show a superparamagnetic (SP) behavior (i.e. the magnetization shows a relaxation because of thermal activation). Larger particles show multiple magnetic domains (MD). There are no sharp boundaries between these three size ranges. The transition between the SP and SD particles depends on the relation between the relaxation time τ and the time scale

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2      S. Dutz et al.: Magnetic nanoparticles for biomedical applications of the field variation (e.g. the AC field period 1/f in hyperthermia). For maghemite particles under typical hyperthermia conditions, the critical diameter is between 10 and 20 nm [25]. The transition between SD and MD starts gradually with a slight inhomogeneity of the magnetization, which increases with increasing diameter and finally results in the appearance of domain walls. Theoretically, for magnetite these effects appear for particle diameters d > 70 nm [19]. MD particles are not suitable for biomedical applications because the suspensions of particles of this size are not stable against sedimentation. Within one size range, there are gradual variations of the magnetic properties. If hyperthermia is performed with SP particles, the specific heating power (SHP) depends on the particle size d via the Néel relaxation time τ = τ0 ⋅e

K⋅V (d) kT ⋅

with τ0 ≈ 10 -9 s,

where K is the anisotropy constant, V is volume, k is the Boltzman constant, and T is temperature [35]. The SHP shows a maximum for f≈1/τ. In a similar way for SD particles, the maximum SHP can be achieved if the anisotropy field Hk(d) is in the order of the external field amplitude Hac. Choosing the optimum parameters also requires the consideration of the outer constraints of the field para­ meters [10]. To keep the heating of the unloaded tissue due to eddy current losses sufficiently low, the product Hac·f must be within a certain limit (e.g. 4.8 × 108 A/m s), as it was given by Brezovich [3, 5]. In drug targeting, the situation is different. The force acting on an MNP increases linearly with the volume if the field is large enough to saturate the particle. The upper limit of the size is given by the sedimentation stability of the particles in the liquid and the diameter of blood vessels. MPI requires particles with low hysteresis and a steep magnetization curve (i.e. a high susceptibility). This situation is given at the upper limit of the SP range [i.e. if the diameter d fulfills τ(d)≈1/f]. In summary, we can state that for each of the applications mentioned above, there is an optimum particle size and each application needs its own particles. Real nanoparticle materials, however, show a more or less broad distribution of particle sizes (i.e. they consist of a mixture of optimal and nonoptimal particles). The aim of the investigations presented in this paper is to improve the particle properties for biomedical applications (i.e. to obtain a mean size that corresponds to the field parameters and a reduced size distribution width). These investigations comprise improved preparation methods such as cyclic precipitation and the preparation of multicore particles as well as physical after-treatments such as size-dependent fractionation to narrow the size distribution. In this paper,

we present a summary of our findings from recent studies for the tuning of MNP properties regarding their potential application in biomedicine. In addition to the physical requirements, the biocompatibility is an important issue. First, the particles should consist of nontoxic materials, excluding many materials such as barium ferrite or metallic cobalt, which are favorable from the magnetic point of view. The most common materials for biomedical applications are the magnetic iron oxides magnetite (Fe3O4) and maghemite (γ-Fe2O3), which are proven for their biocompatibility [42]. Second, the particles should allow the preparation of sedimentation stable suspensions in aqueous liquids. This requires particles with only weak magnetic interaction. SP particles meet this requirement very well. SD particles show a remnant magnetic moment and therefore a significant interaction that usually results in agglomeration. A possibility to overcome this effect is the use of multicore particles consisting of several SD primary particles with differently oriented magnetization, which reduces the total magnetic moment and therefore the interaction between the particles in absence of an external field. The coating of particles with suitable organic molecules such as carboxymethyl dextran (CMD; as used in our studies) improves the stability of the suspensions. Moreover, such molecules can be used to bind other molecules such as drugs.

Materials and methods Particle preparation The two types of particles discussed in this paper were prepared by two similar wet chemical precipitation methods [12, 32] based on precipitation in alkaline media [29, 31]. The first type of particles was prepared by a “cyclic” precipitation method [32]. A NaHCO3 solution was added to a FeCl2/FeCl3 solution (Fe2+/Fe3+ ratio = 1:1) up to pH 7, which led to the formation of a brownish precipitate. Then, a new Fe2+/Fe3+ mixture was added and the precipitation was carried out again. This procedure was repeated up to four times. After that, the solution was boiled for 10 min to form an almost black precipitate. To remove excess reaction products from the prepared particles, the particles were washed with deionized water three times. The second type are magnetic multicore nanoparticles (MCNPs) [11, 12]. In detail, a 1  M NaHCO3 solution was slowly added under permanent stirring to a FeCl2/FeCl3 solution (total Fe concentration: 0.625 M; Fe2+/Fe3+ ratio = 1/1.3) with a rate of 0.75 ml/min. This procedure was stopped when the pH value reached 8. During this routine, a brownish precipitate was formed, which was heated to 100°C for 5 min, and iron oxides with a spinel structure were formed under the release of carbon dioxide (CO2). To remove excess reaction products, the prepared particles were washed with deionized water three times.

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S. Dutz et al.: Magnetic nanoparticles for biomedical applications      3 For the production of sedimentation stable suspensions of the MCNP, the particles were coated with biocompatible coating material CMD (initial material: CMD sodium salt from Fluka, Buchs, Switzerland). After washing of the particles, the pH value of the dispersed particles was adjusted to pH 2–3 by the addition of diluted HCl. Then, the suspension was homogenized by ultrasonic treatment for approximately half a minute (Sonopuls GM200, Bandelin electronic) and then kept at 45°C. An aqueous solution of CMD with a CMD/ MCNP ratio of approximately 1/3 was added to the suspension and stirred for 60 min at 45°C. Finally, the coated particles were washed with deionized water to remove the remaining salts. The magnetic particle concentration in the suspensions was adjusted to values up to 7% by mass.

Size-dependent fractionation To perform size-dependent fractionation, two different methods were applied: centrifugal force fractionation as well as asymmetric-flow field flow fractionation (AF4). For the centrifugal fractionation of MCNP into fractions of ­different mean sizes, 30  ml of the fluid were filled in a cylindrically shaped centrifugation vessel made of glass (sample height≈70 mm and diameter≈20 mm). The sample was centrifuged in a laboratory centrifuge (Cryofuge 6000, Heraeus Sepatech) at 1000 × g with a temperature of 20°C. The sediment was stored and the supernatant was recovered. A portion of the supernatant was centrifuged again at 1500 × g. This procedure was repeated twice with increasing centrifugal accelerations (2500 × g, 3000 × g). In consequence, four sediments and four supernatants were obtained representing eight fractions of the MCNP. The fractionation principle of the AF4 method is based on Stokes forces acting on dispersed particles with different strengths and directions depending on their size within a fluidic channel [22]. For this, a cross-flow was applied perpendicularly to the main flow in the separation channel to move the particles towards the semipermeable bottom wall (membrane) of the channel to fractionize the sample depending on size (Figure 1). Smaller particles diffused back into the middle of the main flow faster than the larger ones and elute first [20]. In our experimental investigations, we used an Eclipse separation system (Eclipse F, Wyatt Europe, Dernbach, Germany) connected to an isocratic pump and degasser (Agilent 1100 series, Agilent Technology, Böblingen, Germany). The separation system was connected with a multiangle laser light scattering detector (MALLS; DAWN EOS, Wyatt Europe) and data were collected in intervals of 1 s. Then, 100 μl of the sample were injected in the focus mode (focus flow 2 ml/min). After finishing the injection, the sample was eluted at fixed main flow rate (detector flow) of 1 ml/min starting with cross-flow decreasing from 2.0 to 0.2 ml/min within 5  min followed by a cross-flow

Figure 1: Sketch of the fractionation channel and the principle of AF4.

decreasing from 0.2 to 0 ml/min within 20 min. Elution proceeded over 25 min without cross-flow to ensure a complete sample elution and the samples were taken every 1 min.

Structural and magnetic characterization The dry samples of MNP were structurally characterized by field emission scanning electron microscopy (JSM 6300-F, JEOL, Japan), transmission electron microscopy (TEM; JEM 2010FEF, JEOL, Japan or DSM 960, Zeiss, Jena, Germany), and X-ray diffraction (XRD; X’pertTwin diffractometer, Philips or X’Pert PRO, PANalytical, Almelo, The Netherlands). The results of the XRD investigations gave information about magnetic phase composition [13]. The mean sizes (d) of the magnetic cores were calculated from the measurements of the XRD line width by using the Scherrer formula. The physical cluster size as well as the agglomeration behavior of the particles was derived from the TEM images. The hydrodynamic diameters (dh) of the CMD-coated particles or clusters in the ferrofluid were determined using dynamic light scattering (DLS; Zetasizer nano ZS, Malvern Instruments, Malvern, UK) and MALLS (DAWN EOS, Wyatt Europe) for the AF4 fractionated samples. The quasi-static magnetic properties of the dried particles as well as ferrofluids were investigated by a vibrating sample magnetometer (VSM). We used a MicroMag™3900 system (Princeton Measurements Corp., USA). The measurements included hysteresis loops in a wide range of maximum fields up to the saturation state and remanence curve measurements. By integrating the area of the measured minor loops, the specific hysteresis losses (SHL) per cycle depending on the field amplitude were calculated. From the SHL, a rough estimation of the SHP (a.k.a. SAR or specific absorption rate) of the particles was obtained by multiplying the SHL with the frequency of the applied AC field. For the SP samples, there is a well-established method to measure the magnetic size distribution [7]. Because this method is restricted to SP particles, we used switching field distribution for the characterization of ferrimagnetic particles used in our investigations. Remanence curves [gradual measurement of the remnant magnetization Mr(H = 0) vs. varying magnetic field strengths H] enable the investigation of the switching field distribution S(H) of particle ensembles, the immobilization state, and the particle interaction. The switching field distribution S(H) is the distribution of the amount of particles that switch their magnetization irreversibly at the field H [38, 44] with S( H ) =

1 dM r ( H ) ⋅ M rs dH

S(H) was fitted by a lognormal distribution. The influence of the statistical orientation of the particles and the distribution of the anisotropy constant on S(H) can be seen in Ref. [44]. The switching field distribution does not correspond exactly with the particle size distribution but depends as well on the possible distribution of the effective magnetic anisotropy constant, which includes the crystal anisotropy (constant for a material) and the shape anisotropy of the particles. The latter may be distributed as well. The SHP was measured by means of magnetic field calorimetry at a field amplitude of up to 25 kA/m and a frequency of 400 kHz (investigations on biological samples) or at fields up to 30 kA/m and 210 kHz (particles with higher coercivity). These parameter combinations are suitable for hyperthermia treatment [27]. Beside

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4      S. Dutz et al.: Magnetic nanoparticles for biomedical applications measurements on mobile particles within fluids, the SHP of immobilized particles was determined for particles dispersed in a gelatin gel. It was shown in previous investigations that this method yields a strong immobilization of the particles. For the measurement of heating curves, the particle suspensions (usually 1 ml in a 2 ml Eppendorf tube) were thermally isolated in a PUR foam block and placed within the coil (Figure 2). For the temperature measurements, a fiber optic device (Fotemp, OPTOcon, Dresden, Germany) was used and the probe was placed in the center of the sample. Heating curves were recorded for temperature as function of time. The obtained curves were fitted with suitable mathematical functions (linear or third-degree polynomial, depending on measuring conditions) and a temperature drift without magnetic heating (in case the sample temperature is different from the room temperature) was corrected by a linear term. The first derivative of these functions at time point = 0 s provides an initial temperature slope (dT/dt) representing the adiabatic case, where all energy is absorbed from the sample and no heat dissipation to surrounding takes place. The SHP was calculated by SHP =

dT ms ⋅c ⋅ dt mp

where c is the specific heat capacity of the sample, ms is the total mass of the fluid sample, and mp is the mass of the iron oxide particles in the sample. Mostly, the specific heat capacity of water was used for the samples with a particle concentration below 2% in the suspension [33]. For higher concentrated samples, a more realistic c was calculated from the composition of the sample. The concentration of the samples was determined by measuring the specific saturation magnetization using the VSM and assuming the specific magnetization values, which were obtained from the powder samples. The Brownian relaxation behavior of the MCNP in the fluid was investigated by magnetorelaxometry (MRX) [18]. For these measurements, the samples were diluted by deionized water (dilution factors from 1:9 to 1:900). After applying and switching off a magnetizing field, a highly sensitive low-TC-SQUID sensor measures the magnetic induction B(t) above the sample. From these relaxation curves, the size distribution of hydrodynamic diameters of the coated particles was calculated by fitting the so-called cluster moment superposition model (CMSM) to the relaxation data [17]. The CMSM describes the

relaxation of the magnetic moment of noninteracting SD MNP in the presence of Brownian [6] and Néel [35] relaxation mechanisms. The distribution of the hydrodynamic diameters dh is assumed to be a lognormal one. Previous investigations showed a good agreement with the hydrodynamic diameter obtained by DLS [18].

Biomedical application The suitability of prepared MNP for medical application has to be confirmed by means of ex vivo and in vivo experiments. In our studies, the focus for application was on the investigation of cellular uptake of MCNP, the alteration of magnetic properties of the particles due to particle-tissue interactions after particle administration, as well as the hyperthermia heating performance of MCNP within tissue in mice. The interaction of the MCNP with human cells (cellular uptake) was studied with the breast cancer cell line MCF-7. The cells were cultivated in Dulbecco’s modified Eagle medium + 10% fetal calf serum. For incubation experiments, adherent cell culture cells were detached with a trypsin/EDTA solution and inoculated with MCNP in phosphate-buffered saline/EDTA (2 mmol) for the times indicated. After treatment, the magnetically labeled cells were separated by MACS (SuperMACS and MS columns, Miltenyi Biotec). The efflux was designed as “negative fraction” and the cells retained in the column were designed as “positive fraction.” Total cell numbers were estimated by cell counting (Coulter Z2, Beckman-Coulter). To investigate the in vivo conditions with respect to the immobilization of the particles and its influence on the heating ability and mechanism, animal experiments on female severe combined immunodeficient mice were carried out at the Jena University Hospital [14]. One experimental tumor was grown on each of four mice between the shoulder blades. The tumor volumes after 8 weeks of growth were calculated to be between 60 and 110 mm3. Before intratumoral MCNP injection (100 μl, MCNP concentration = 70 mg/ml, hydrodynamic particle diameter = 317 nm) and magnetic heating treatment (H = 25 kA/m, f = 400 kHz), the animals received an anesthetic gas. X-ray images were taken immediately before and after MCNP injection. Thirty minutes after MCNP injection, one mouse was exposed to an AC field for 150 s, reaching thermal ablative temperatures recorded via thermocouples

Temperature probe

PC

Temperature (°C)

100

Sample

80 60 40 20 0

PUR foam

0

50 100 150 200 250 300 350 Time (s)

Coil

Figure 2: Scheme of the SHP measurement setup.

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S. Dutz et al.: Magnetic nanoparticles for biomedical applications      5

of copper and constantan wires as described previously [28], and a further X-ray image was taken 2 h after magnetic heating. The other three mice served as controls (i.e. without magnetic heating) for the particle distribution in the tissue and the determination of the particle concentration in the tumor after injection by means of histological and VSM investigations. Mice were sacrificed through euthanasia by means of CO2 2 h after magnetic heating and MCNP injection, respectively. Tumor dissection was followed by tissue fixation and embedding, cutting, and staining for light microscopy examinations to study the spatial distribution of the particles in the tissue and to confirm the effects of magnetic thermal ablation on tumor cells. One complete tumor (diameter ∼5 mm and mass 96 mg) was measured as a whole in VSM immediately after resection, and the obtained hysteresis curves were compared to the mobile particles under fluid conditions and the particles immobilized in gelatin to evaluate the particle immobilization within the tumor tissue.

Results and discussion Particles prepared by cyclic growth The structural characterization of cyclic grown MNP by means of TEM investigation revealed particle sizes increasing from approximately 10 nm up to several 10 nm for one and four cycle particles, respectively (Figure 3). On the contrary, the sizes from XRD increased only from 10 to 19 nm for the same samples. This might be explained by a nonepitaxial growth of the outer shells on the inner cores, which results for XRD in a multicrystal structure. As found by XRD, all samples consist of spinel phase with peak positions between those of γ-Fe2O3 and Fe3O4. Smaller particles (one or two cycles) indicate mostly γ-Fe2O3. Nevertheless, single magnetic layers are in strong exchange interaction, which results in an increasing effective magnetic volume with increasing number of cycles.

Figure 3: TEM image of a three-cycle sample.

The coercivity Hc increases from 0.72 kA/m for particles from one cycle up to 6.2 kA/m for four cyclic prepared particles. The remanence curves at room temperature and corresponding switching field distributions are shown in Figure 4. It becomes clear that, with increasing number of cycles, the particles show a more distinct hard magnetic behavior. The mean value Hm of switching field distribution S(H) increases as expected with an increasing coercivity. Surprisingly, the switching field distribution becomes narrower with an increasing number of cycles. That can be explained by a partly growing of particles in cycles 2–4 starting from the particles of the previous cycle with diminished nucleation of new particles. With decreasing SP fraction in particle ensemble (means increasing cycle number), the ratio Hc /Hm increases, which might be explained by the fact that Hm includes only contributions from thermally blocked (hysteretic) particles, whereas Hc is reduced when the sample contains an SP proportion. The relative remnant magnetization Mr /Ms increases unproportionally with coercivity, which is also an indication for a certain multicrystal portion in the samples with higher numbers of cycles (Figure 4). In general, Mr /Ms increases with cycle number as expected but does not reach the values for single core samples of comparable mean size. This induces the existence of a significant fraction of SP particles within all different samples. To elucidate the effect of the SP fraction on the overall magnetic behavior of the samples, low-temperature magnetic measurements (10, …, 175 K) were carried out. In these investigations, it was found that the particles prepared in one cycle contain a larger proportion of SP cores than the particles prepared by four cycles. This absence of small SP particles in the samples prepared in more than one cycle is a clear indication for the cyclic growth without further nucleation of new small particles. Detailed results to the low-temperature measurements can be found in the study of Mueller et al. [32]. Because magnetic heating power corresponds to the hysteresis losses during the reversal of magnetization [9, 10], the areas of hysteresis curves (SHLs) were calculated as a measure of the heating performance of the samples. At low field strength (  ” sign. Due to the still low amount of particles in each fraction and thus a poor signal-to-noise ratio, a determination of the quasi-static SHLs as a measure for the heating capability of the magnetic particles in an AC field as well as the

Table 2: Summary of all determined parameters of the samples obtained from AF4. Samples unified for the measurements in VSM are indicated by a bracket ( > ). Fraction     Original   1   2   3   4   5   6   7   8   9   10   11   12   13   14   15   16  

D (nm)  205  122  154  137  155  161  174  185  203  212  242  249  296  325  400  421  437 

DLS  MALLS  MRX        PDI D (nm) D (nm) 0.38  0.16  0.28  0.16  0.18  0.10  0.09  0.13  0.16  0.15  0.23  0.19  0.31  0.32  0.38  0.42  0.42 

215  123  –  153  155  182  –  203  –  –  273  –  322  –  –  353  372 

197  60  78  84  104  116  151  174  203  236  266  310  359  427  507  572  612 

       >     >     >     >     >     >     >     >    

calorimetrical determination of the SHP of the particles in the fractions was not possible. However, the hysteresis loop up to saturation field strength was measured and the coercivity field Hc and remanence Mr were determined. Because SHLs correlate well with the coercivity (the particles with the highest coercivity show the highest SHL [13] at ­saturation field), one can draw a connection between Hc and SHL/SHP. As a starting sample, CMD-coated MCNPs were used. The particles had a primary particle size of 14 nm (XRD), a cluster size of 65 nm (TEM), and a hydrodynamic diameter of approximately 200 nm (DLS, MALLS D 50, MRX). VSM measurements revealed a coercivity of 1.41 kA/m and a mean switching field Hm of 12.5 kA/m. Due to the soft aggregation of the larger clusters, a slightly bimodal particle size distribution in the original fluid was found in MALLS and DLS investigations. Because the magnetic measurements were carried out on dry powders, this aggregation showed no influence on the magnetic investigations. After setting up a suitable fractionation protocol by pre-measurements, the starting sample was fractio­ nated into 16 samples containing magnetic MCNPs of different size. Figure 6 shows the particle diameter of mean hydrodynamic volume calculated by fitting CMSM to MRX data (where mainly the Brownian relaxation of the particles contributed to the signal) as well as values from DLS (z-average). With increasing fraction number, the hydrodynamic diameter increases from 60 up to approximately 600 nm. Slight differences between the mean diameters determined by means of DLS and MRX were attributed to the differences in the weight of the particle sizes with

VSM Hc (kA/m)  Hm (kA/m) 1.41    0.41     1.17     1.98     2.11     2.21     2.25     2.84   3.83   

12.5 5.1 6.6 8.5 9.1 14.0 15.7 17.3 21.7

Figure 6: Hydrodynamic diameters of obtained fractions (AF4), ­calculated from MRX curves (mean volume equivalent) in ­comparison with data from DLS (z-average). From Ref. [15]. © IOP Publishing. Reproduced with permission of IOP Publishing. All rights reserved.

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S. Dutz et al.: Magnetic nanoparticles for biomedical applications      9

respect to their contribution to the signal intensity. Also, the application of different mathematical models (e.g. assuming a lognormal distribution in CMSM) for data analysis may contribute to observed differences. However, identical trends can be seen in the qualitative d ­ ependences of hydrodynamic diameter on fraction number. A summary of all important structural and magnetic parameters of the investigated fractions is given in Table 2. The first eluted particles with the smallest diameter show a narrow size distribution, whereas the larger particles eluted at the end of the fractionation have a broader distribution as can be seen from polydispersity index (PDI). The obtained values for coercivity and Hm of S(H) confirm a successful fractionation. It was found that the coercivity after size-dependent fractionation of the original sample (1.41 kA/m) increases with increasing cluster size from 0.41 kA/m for the smallest fraction up to 3.83 kA/m for the fraction with the largest clusters. A similar behavior was found for S(H): Hm of 12.5 kA/m for original sample increased in the fractions from 5.1 up to 21.7 kA/m with increasing cluster size. In summary, AF4 was found to be a very powerful method for the fractionation of MNPs, allowing the preparation of very small entities from the starting sample. Because several consecutive separation runs revealed very similar particle sizes in the fractions prepared under similar conditions, rendering AF4 fractionation is a method with a high reproducibility.

Particle-tissue interactions Another important factor for the optimization of magnetic particles is the knowledge about the influence of surrounding tissue on the magnetic properties of particles administered to biological systems. Especially for magnetic hyperthermia, it is of particular interest if particles are fixed to biological structures in the tissue or if they are mobile and able to rotate. For this, the magnetic behavior of MNP suspended in water, immobilized in a gelatin gel, and applied to living tumors in animal experiment was determined by means of VSM hysteresis curves. A visual investigation of the resected tumors showed particle accumulation in certain regions of the tumor. Within the tumor, a mean particle concentration of approximately 0.9% by mass was determined by magnetic measurements. From the shape of hysteresis curves (Figure 7) for a maximum field strength of 25 kA/m (which is comparable to applied fields during hyperthermia treatment), it becomes obvious that the particles in the tumor show similar magnetic behavior like particles immobilized in

Figure 7: Comparison of the normalized minor loops for mobile MCNPs in fluid, immobilized MCNPs (gelatin), and MCNPs inside the tumor (H = 25 kA/m). From Ref. [14]. © IOP Publishing. Reproduced with permission of IOP Publishing. All rights reserved.

gelatin. There are only slight differences in coercivity and remanence, which are probably caused by the aggregation of the clusters during immobilization. This immobilization of particles within tissue might prevent a movement or rotation of particles inside the tissue and the reversal of magnetization can take place by intrinsic rotation only. For this case, ferrimagnetic particles show higher heating than SP ones if the field is sufficiently high. To confirm this finding, the coercivity, remnant magnetization, and S(H) were analyzed for all samples (see Table 3). The immobilized particles show a coercivity of 3.91 kA/m, which is a clear indication for hysteretic behavior. Thus, these particles are suitable for the investigation of the immobilization status by means of the S(H). The immobilized particles show an Hm of approximately 13 kA/m and the mobile particles of 0.7 kA/m. Ideally, the Hm for the mobile particles should be 0. However, particle agglomeration or the formation of chains leads to a small hysteresis (Hc, Hm > 0) caused by particle interactions within the agglomerates. Assuming the immobilization degree to be 100% for particles immobilized in gelatin and 0% for fluid particles, the immobilization rate of the particles in the tumor 2 h after particle injection is approximately 89%. This means that the main proportion of the particles is fixed to the tumor tissue. The immobile particles in gelatin and the particles in the tumor after 2 h show nearly the same behavior regarding agglomeration as can be seen from the coercivity and the relative remanence. With increasing time (4, 6, and 24 h) after injection, the immobilization degree decreased down to 85%. Simultaneously, the coercivity and the relative remanence

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10      S. Dutz et al.: Magnetic nanoparticles for biomedical applications Table 3: Magnetic properties of all immobilization experiment samples measured by VSM. Sample



Mobile   Immobile   Tumor, after 2 h   Tumor, after 4 h   Tumor, after 6 h   Tumor, after 24 h 

Hc (kA/m) 

1.31  3.91  3.85  4.11  4.14  4.20 

Mr/Ms    0.1078  0.1787  0.1754  0.1829  0.1855  0.1869 

increased with decreasing immobilization degree. A possible reason for this contrasting behavior could be the decomposition of the tissue due to the decay of the cells of the tumor. This would lead to a more fluidic consistence of the tissue and a lower immobilization rate of the particles. Additionally, the particles became more mobile and magnetic attraction forces between the single clusters could have led to agglomeration of the clusters, which lead to an increased coercivity and relative remanence. In summary, the quasi-static magnetic characterization of the loaded tumor tissue ex vivo showed a strong immobilization of the used particles in the investigated type of tumor tissue. This means that the particles are bound very strongly to the tissue and a reversal of magnetization can only take place by Néel relaxation (for SP particles) or by overcoming the coercivity (for ferrimagnetic particles). A contribution to the heat generation by Brownian relaxation can be neglected. Only a small proportion due to oscillation (the Brownian rotation of very limited torsion due to the elasticity of the tissue) is conceivable. This finding is of major importance for the correct choice of the parameters of the AC field for magnetic hyperthermia depending on the size of the particles.

Biomedical applications The interaction of the MCNP with human cells was analyzed with the breast cancer cell line MCF-7 in cellular uptake studies. For this, MCNPs and cells were incubated for the defined duration and then separated magnetically into MNP-loaded (positive fraction) and MNP unloaded (negative fraction) cells. It was found that the tumor cells are labeled rapidly with the nanoparticles. Within 4 min, more than 50% of the cells are loaded with MNP and are detected in the positive fraction (Figure 8) in the MACS system. Prolonged incubation leads to an increase of cell content in the positive fraction up to 85%. These data are in good correlation

SFD  Hm (kA/m) 

σ

0.72  13.29  11.86  11.62  11.54  11.38 

0.67  0.67  0.68  0.68  0.68  0.68 

Immobilization (%)

0 100 89 87 86 85

Figure 8: Uptake of MCNP in MCF-7 tumor cells as a function of the incubation time. For comparison data of single core MNP obtained in an earlier investigation, shown as hatched bars. From Ref. [12]. With permission of Elsevier.

to previous results with CMD-coated MNP with a mean diameter of 10 nm [40]. These data show that the cellular uptake of MCNP is similar to that of small (10 nm) particles and thus the here-described CMD-coated MCNPs are a suitable tool for cell separation. MCNPs exhibit their highest potential in hyperthermia application. In our animal investigations, a particle concentration of 0.9% by mass was achieved by the intratumoral injection of 100 μl of a ferrofluid into a tumor with a mass of approximately 100 mg. In vivo magnetic heating experiments showed that this concentration in combination with an SHP of 400 W/g for mobile particles or 262 W/g for the same particles under immobilized conditions was sufficient to reach thermal ablative temperatures within a 5 mm diameter tumor. The temperature increase dT reached 20 K in the first 60 s of magnetic heating treatment and 25 K after 140 s. As a reference, a thermocouple was placed into the rectum showing only a slight temperature rise of 2  K during magnetic heating for 140 s. This

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S. Dutz et al.: Magnetic nanoparticles for biomedical applications      11

indicates that the temperature generation from MCNP is very suitable for magnetic particle hyperthermia and the heating is limited to the region of the tumor. For diagnostic purposes, it was found that the heredeveloped MCNPs show promising behavior as tracers for MPI [16]. It could be demonstrated that MPI performance depends mainly on the cluster size of MCNP. For best performance, these particles need further development regarding the size of primary particles and clusters.

Conclusions In our studies, novel aqueous precipitation procedures for magnetic iron oxide nanoparticle synthesis were developed. One procedure focused on the cyclic growth of MNPs without the nucleation of new particle cores during precipitation. This led to larger ferrimagnetic particles with a relatively narrow size distribution, which is favorable for most medical applications. As a second novel particle type, we developed so-called magnetic multicore particles, which consist of SD subgrains of approximately 10 nm forming dense clusters in the size range from 40 to 80 nm. Due to the statistical orientation of magnetic easy axis of subgrains within clusters, the resulting magnetization without an external magnetic field is comparable low. This allows a sedimentation stable dispersion of these relatively large magnetic structures. A further promising approach for the optimization of magnetic particle properties is the size-dependent fractionation. A fivefold SHP of best fraction compared to original sample after centrifugation fractionation confirmed the suitability of size-dependent fractionation for particle optimization. AF4 was found to be a very reproducible fractionation method for MNPs, allowing the preparation of very small entities from the starting sample. After the optimization of magnetic multicore particles and the transfer of particles into ferrofluids, these particles were tested for their performance in biomedical applications. First, human tumor cells were incubated with these particles and a good cellular uptake of particles resulted. The magnetically labeled tumor cells can be removed from healthy cells by a magnetic separation process now. For diagnostic purposes, it was found that the here-developed ferrimagnetic particles show a promising behavior as tracers for MPI. For best performance, these particles need further development regarding the size of subgrains and clusters. Magnetic multicore particles feature their highest potential in hyperthermia application. During in vivo experiments in mice, it was found that, immediately after

administration to living tumor, the particles were fixed to cellular structures. This prevents a movement or rotation of particles inside the tissue and the reversal of magnetization can take place by intrinsic rotation only. For this case, ferrimagnetic particles show a higher heating power at sufficient field amplitudes than the SP ones. In our experiments, therapeutically suitable temperatures ( > 55°C) were reached after 20  s of heating for a particle concentration in the tumor of 1% by mass and field parameters of H = 24 kA/m and f = 410 kHz, whereby a tolerable body temperature increase of approximately 2 K was observed. In our investigations, we could demonstrate that the magnetic properties of MNPs can be purposefully tuned by the modification of their structure, size, and size distribution and the here-developed particles show a high potential for various biomedical applications. Acknowledgments: The authors gratefully acknowledge financial support by the Deutsche Forschungsgemeinschaft (PAK 151: ZE 825/1-1, HI 689/7-1, TR 408/4-1). The authors thank Dr. Joachim Clement (University Hospital Jena) for the investigation of cellular uptake, Prof. Judith Kuntsche (Syddansk Universitet Odense/University of Halle) for AF4 experiments, Dr. Melanie Kettering und Susann Burgold (IDIR, University Hospital Jena) for hyperthermia animal experiments, as well as their colleagues from IPHT Jena: Christa Schmidt (XRD), Franka Jahn (TEM), Hanna Steinmetz, and Dr. Rudolf Hergt for their support. Furthermore, the authors thank Dr. Lutz Trahms (PTB Berlin), Prof. Stefan Odenbach (Technische Universität Dresden), Prof. C. Alexiou (University Hospital Erlangen), and their coworkers for the fruitful collaboration in the framework of PAK 151.

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Magnetic nanoparticles adapted for specific biomedical applications.

Magnetic nanoparticles (MNPs) are used in different biomedical applications, whereby each application requires specific particle properties. To fulfil...
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