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Theranostic Nanoparticles for MRI-Guided Thermochemotherapy: “Tight” Clustering of Magnetic Nanoparticles Boosts Relaxivity and Heat-Generation Power Koichiro Hayashi, Yoshitaka Sato, Wataru Sakamoto, and Toshinobu Yogo ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.6b00536 • Publication Date (Web): 17 Nov 2016 Downloaded from http://pubs.acs.org on November 21, 2016

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ACS Biomaterials Science & Engineering

Theranostic Nanoparticles for MRI-Guided Thermochemotherapy: “Tight” Clustering of Magnetic Nanoparticles Boosts Relaxivity and Heat-Generation Power Koichiro Hayashi,* Yoshitaka Sato, Wataru Sakamoto, and Toshinobu Yogo Division of Materials Research, Institute of Materials and Systems for Sustainability, Nagoya University, Furo-cho, Chikusa-ku, Nagoya 464-8603, Japan E-mail: [email protected]

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ABSTRACT. Magnetic-resonance-imaging (MRI)-guided magnetic thermochemotherapy is a

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potentially invasive technique combining diagnosis and treatment. It requires the development of

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multifunctional nanoparticles with: 1) biocompatibility; 2) high relaxivity; 3) high

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heat-generation power; 4) controlled drug release; and 5) tumor targeting. Here, we show the

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synthesis of such multifunctional nanoparticles (“Core-Shells”) and the feasibility of

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MRI-guided magnetic thermochemotherapy using the synthesized nanoparticles. “Tight”

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iron-oxide nanoparticle clustering to zero interparticle distance within the Core-Shells boosts the

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relaxivity and heat-generation power while maintaining biocompatibility. The initial Core-Shell

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drug release occurs in response to an alternating magnetic field (AMF) and continues gradually

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after removal of the AMF. Thus, a single Core-Shell dose realizes continuous chemotherapy over

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a period of days or weeks. The Core-Shells accumulate in abdomen tumors, facilitating MRI

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visualization. Subsequent AMF application induces heat generation and drug release within the

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tumors, inhibiting their growth. Core-Shell magnetic thermochemotherapy exhibits significantly 1 ACS Paragon Plus Environment


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higher therapeutic efficacy than both magnetic hyperthermia and chemotherapy alone. More

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importantly, there are minimal side effects. The findings of this study introduce new perspectives

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regarding the development of materials for MRI, magnetic hyperthermia, and drug delivery

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systems. Both conventional and novel iron-oxide-based materials may render theranostics (i.e.,

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techniques fusing diagnosis and treatment) feasible.

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Keyword: magnetic nanoparticles; MRI; hyperthermia; drug delivery; theranostics

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1. Introduction

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Magnetic nanoparticles generate heat via hysteresis loss, Néel relaxation, and Brownian

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relaxation under the influence of an alternating magnetic field (AMF).1-3 Magnetic hyperthermia

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is a new type of thermotherapy that exploits the exothermic properties of magnetic nanoparticles;

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it may facilitate cellular-level local treatment of deep tumors, because heat is generated only in

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the areas containing magnetic nanoparticles and the AMF can penetrate deep within the body.

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Biocompatible iron-oxide (Fe3O4 or γ-Fe2O3) nanoparticles are potential candidate

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materials for magnetic-hyperthermia exothermic bodies, as they are used as contrast agents for

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magnetic resonance imaging (MRI) in clinical practice. However, they exhibit insufficient

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heat-generation power to increase tumor temperatures to values lethal to cancer cells (above

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42.5°C).4 The particle concentration in tumors, which depends on the particle dosage and

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tumor-targeting ability, AMF strength (H) and frequency (f), and the tumor volume, is also an

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important factor affecting changes in the tumor temperature. Particle dosage should be

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minimized from the perspective of toxicity. Further, considering the harmful effects of AMF on 2 ACS Paragon Plus Environment

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human health, Hf should be less than 5 × 109 A m−1 s−1.5 Therefore, it is important to maximize

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the heat-generation power of iron-oxide nanoparticles within the limits of harmless Hf to

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increase the temperature of the entire tumor to values that are lethal to cancer cells, while

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employing the minimum necessary dosage. Based on the above background, a number of

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materials with high heat-generation capacity have been investigated. Fe nanoparticles and

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exchange-coupled magnetic materials containing MnFe2O4 and CoFe2O4 generate heat

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efficiently compared to their alternatives.6,7 However, these materials have problems of poor

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resistance to oxidation and the toxicity of manganese and cobalt, respectively. Therefore,

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biocompatible, high-heat-generating magnetic nanoparticles are required.

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The heat-generation power of magnetic nanoparticles depends on the particles’ magnetic

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properties, including the saturation magnetization (MS), remanent magnetization (MR), and

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coercivity (HC).8,9 Moreover, magnetic properties are closely related to the magnetic-dipole

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interaction energy (E), which is inversely proportional to the cube of the particle-particle

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separation (l).10-14 Thus, optimization of l may improve the heat-generation power of magnetic

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nanoparticles.

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MRI provides high spatial resolution (< 0.1 mm) and images that provide anatomical,

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physiological, and metabolic information non-invasively in a single imaging session.15 However,

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MRI tumor visualization is inadequate owing to the low sensitivity of this technique.15 A

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high-relaxivity MRI contrast agent is required to improve this sensitivity, which may realize an

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upgraded, powerful imaging modality. As the relaxivity depends on the agent’s magnetic

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properties and particle diameters in T2 imaging, optimization of the magnetic-dipole interaction

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(or l) may enhance both the relaxivity and heat-generation power.

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Hyperthermia

improves

the

therapeutic

efficacy

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of

chemotherapy,16

and

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anticancer-drug-containing and -releasing magnetic nanoparticles can implement both magnetic

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hyperthermia and chemotherapy for enhanced treatment. In addition, such magnetic

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nanoparticles can potentially achieve controlled drug release within tumors, leading to

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single-dose continuous chemotherapy, enhanced therapeutic efficacy, and reduced side effects.

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Multifunctional magnetic nanoparticles that serve as exothermic bodies for magnetic

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hyperthermia, MRI contrast agents, and drug-delivery-system (DDS) carriers may realize

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combined magnetic hyperthermia and chemotherapy under MRI, i.e., MRI-guided magnetic

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thermochemotherapy.

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In this paper, we present methods that enhance both the heat-generation power and

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relaxivity of magnetic nanoparticles, while maintaining their biocompatibility, by increasing the

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magnetic-dipole interactions through “tight” clustering of Fe3O4 nanoparticles. This approach is

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based on our previous paper, in which a new synthesis method for Fe3O4 nanoparticle clusters

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was reported.17 The present study reveals, for the first time, the importance of the

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primary-nanoparticle l in cluster formation for enhanced heat-generation power and relaxivity.

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We also demonstrate anticancer-drug (doxorubicin; Dox) holding and release to the

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Fe3O4-nanoparticle nanoclusters. Furthermore, we report on the potential of our novel material

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for MRI-guided magnetic thermochemotherapy of intraperitoneal tumors that are externally

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invisible to the naked eye.

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2. Results and Discussion

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2.1. Synthesis and Structural Analyses


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Heretofore, nanoparticle clusters have been prepared through nanoparticle synthesis and

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aggregation. Conventional methods yield non-uniform size distributions and large l. We

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previously reported that a reversal in fabrication order yields Fe3O4-nanoparticle clusters with

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acceptable size distributions and zero l.17 Using the approach described in that previous report,

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we synthesized core-shell nanoparticles comprising tightly clustered Fe3O4 nanoparticles (core)

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and a Dox-containing polymer (shell), by creating the former within Dox-containing polymer

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nanoparticles (see Scheme 1 and Experimental Section). We also synthesized regular unclustered

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Fe3O4 nanoparticles (hereafter referred to as “Mag”) as control materials. The Dox-containing

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core-shell nanoparticles were modified with folic acid (FA) and polyethylene glycol (PEG),

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yielding “Core-Shells”.

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In the Core-Shells, primary Fe3O4 nanoparticles, with a diameter of 17 nm, were tightly

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clustered to yield 55 ± 6-nm cluster (core) diameters, as confirmed via transmission electron

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microscopy (TEM; Figures 1a, 1b and 1d). The clusters were coated with 6 ± 1-nm-thick

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polymer (Figure 1e). The Core-Shells had an overall diameter of 64 ± 6 nm (Figure 1f). Constant

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overall Core-Shell diameters were obtained for different synthesis conditions, because the

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polymer nanoparticles acted as templates. However, the primary-nanoparticle crystallite size was

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controlled by the synthesis conditions, as determined via X-ray diffraction (XRD; Figure S1a and

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S2). More importantly, the Core-Shell primary Fe3O4 nanoparticles were densely packed within

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the core with no spaces. In contrast, the Mag aggregates had spaces between the primary

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nanoparticles; these particles had a diameter of 17 nm (Figure 1c), equal to the Core-Shell

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primary particle diameters. Similarly, non-zero l values were previously reported for

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Fe3O4-nanoparticle-loaded nanocarriers such as liposome,18,19 polymersome,20,21 nanogel,22,23 and

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silica24,25, for MRI, magnetic hyperthermia, or DDS. These findings suggest that the 5 ACS Paragon Plus Environment


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magnetic-dipole interactions occurring within the Core-Shells are stronger than those in the Mag

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and other reported materials, because E ∝ 1/l3, such that

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E = ‒ (µ0m02) / (4πl3),

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where µ0 is the magnetic permeability of vacuum and m0 is the magnetic moment.10-14

(1)

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The structural properties are discussed in the Supporting Information (Figure S1). We

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evaluated the Dox stability by recording changes in the Dox absorption spectra before and after

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the Core-Shell synthetic processing treatments (Figure S3). There is little difference between the

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absorbance and peak wavelengths before and after the treatments, indicating that the Dox

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structure remains unchanged even after these treatments. This result indicates that the Dox within

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the Core-Shells maintains its anticancer effect. Furthermore, we evaluated the dispersion stability

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of the Core-Shells in physiological solutions, i.e., phosphate buffered saline (PBS; pH 7.4) and

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Good's buffer (pH 6.0), by recording the changes in the Core-Shell transmission spectra, both

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immediately and 90 days after dispersion (Figure S4). There was little difference in the

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transmittance between the fresh and 90-day-old dispersions. Thus, the Core-Shells dispersed

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stably in physiological solutions at least for 90 days regardless of the concentration and kind of

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salt and pH.

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2.2. Magnetic Properties

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Figure 2a shows the zero-field-cooling−field-cooling (ZFC-FC) curves of the

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Core-Shells and Mag in Figure 1a. The Core-Shells exhibited higher magnetization (M) than the


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Mag at 100 Oe in the 5−400-K range: M was corrected for the TG-measured Fe3O4 content. The

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magnetic-dipole interaction is described by the dimensionless dipole strength λ:

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λ = ‒ πµ0r3H2χ2 / 9kT,

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where k, T, r, χ, and H are the Boltzmann constant, temperature, the particle radius, magnetic

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susceptibility, and applied magnetic field, respectively.26 Further, M = χ × H. Thus, the

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magnetic-dipole interaction scales with M2 and r3. Recall that the Core-Shells have the same

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primary Fe3O4-nanoparticle and crystallite sizes as the Mag (Figures 1a, 1e, and 1f). Therefore,

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the difference in the magnetic-dipole interaction between the Core-Shells and Mag is reflected in

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M. Based on the above findings, tight Fe3O4-nanoparticle clustering increases the

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magnetic-dipole interactions in the Core-Shells, and consequently, the M of the Core-Shells was

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higher than that of the Mag in the ZFC-FC results. Furthermore, the Core-Shells had 50-K higher

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blocking temperature (TB) than the Mag (Core-Shell and Mag TB: 400 and 350 K, respectively),

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where

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TB = EA / k ln(τm / τ0).

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Here, EA, τm, and τ0 are the energy barrier, magnetic measurement time, and relaxation time

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constant, respectively.27 Further,

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EA = KV + nMS4V4(3cos2ψ ‒ 1)2 / 3kl6,

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where V, n, and ψ are the volume of the primary nanoparticle, the average number of nearest

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neighbors, and the first neighboring particle’s location, respectively.12 Thus, l is one of the

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important parameters affecting the change in TB. The Core-Shell l (≈ 0) was significantly shorter

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than that of the Mag (Figure 1). Therefore, the Core-Shells possessed higher TB than the Mag.

(2)

(3)

(4)



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The above results are consistent with previous Monte Carlo (MC) simulations, which show that

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magnetic-dipole interactions lead to increased system TB.28,29

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Figures 2b and 2c show the Core-Shell and Mag magnetization curves for TEM images

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at room temperature (RT). Both the Core-Shells and Mag exhibited ferrimagnetism at RT,

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reflecting the fact that Fe3O4 is ferrimagnetic. The magnetization curves demonstrate that the MR

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and HC of the Core-Shells were higher than those of the Mag for the same crystallite size and

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particle diameter. Figures 2d‒f show the relationships between the primary-nanoparticle

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crystallite sizes and HC, MR, and MS at RT. The Core-Shells possessed significantly higher HC

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than the Mag in the 12−26 nm range (Figure 2d). The MR of the Core-Shell trended to be higher

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than that of the Mag (Figure 2e). The difference between the MR values of the Core-Shells and

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the Mag increases with the increasing crystallite size. MC simulations show that below TB,

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magnetic-dipole interactions increase MR and HC in a strongly dipolar system.28 In contrast, in a

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weakly or moderately dipolar system, MR and HC are not significantly increased by

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magnetic-dipole interactions. Based on these MC simulations and the MR and HC results, a

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Core-Shell is considered to be a strongly dipolar system, whereas Mag is a weakly or moderately

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dipolar system. This conclusion is also supported by the TEM and ZFC-FC results. On the other

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hand, there was little difference in MS between the Core-Shells and the Mag, though the M of the

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Core-Shells at the magnetic microfield was significantly higher than that of the Mag, as shown in

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the ZFC-FC (Figure 2a) and magnetization curves (Figure 2c). These findings indicate that

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magnetic-dipole interactions affect M at the magnetic microfield, but have no effect on M at a

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high magnetic field or MS. This is because MS is an intrinsic material property. The above

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findings demonstrate that tight magnetic-nanoparticle clustering produces strong magnetic-dipole

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interactions and increases M at the magnetic microfield, HC, and MR. 8 ACS Paragon Plus Environment

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For additional comparison, we measured the MS, MR, and HC at RT of the isolated Fe3O4

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nanoparticles with a 50 nm particle diameter (designated 50 nm Fe3O4), which is almost equal to

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the Core-Shell core diameter. The MS, MR, and HC of the Core-Shells were 74.5 emu/g, 8.9

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emu/g, and 42.6 Oe, respectively, and those of the 50 nm Fe3O4 were 60.0 emu/g, 6.2 emu/g, and

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64.7 Oe, respectively (Figure S5). Although magnetization curves showed that the MS of the

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Core-Shells was higher than that of the 50 nm Fe3O4 (Figure S5), the MS of 50 nm Fe3O4 fell

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within error range of Core-Shell MS (Figure 2f). Therefore, there was little essential difference in

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MS between Core-Shells and 50 nm Fe3O4. In general, HC increases with the increasing primary

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particle diameter within the single-domain size range. Therefore, the HC of the 50 nm Fe3O4 was

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higher than that of the Core-Shells. In contrast, in general, M and MR do not necessarily depend

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on the primary particle diameter. The Core-Shells exhibit higher M at the magnetic microfield

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and MR than the 50 nm Fe3O4, suggesting that the magnetic-dipole interactions occurring within

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the Core-Shells are stronger than those in the 50 nm Fe3O4.

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2.3. Relaxivity

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To investigate the transverse relaxivity (r2) values of the Core-Shells and Mag, we

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prepared Core-Shell- or Mag-containing phantoms at 3.1-µM Fe concentration and measured the

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transverse relaxation time (T2). Here,

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1 / T2 = 1 / T2° + r2cFe,

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where 1 / T2° is the reverse proton relaxation time in the absence of Fe3O4 nanoparticles and cFe

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is the Fe concentration.30 The r2 of the Core-Shells was twice that of the Mag (357 and 182

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mM‒1 s‒1, respectively; Figure 3b). Thus, the Core-Shells significantly reduced the MRI signal 9

(6)

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compared with the Mag, as shown in the T2-weighted MR images (Figure 3a). Recent studies

2

have elucidated the contribution of magnetic-nanoparticle clustering to r2 enhancement,31‒33 i.e.,

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it increases the particle size and changes the relaxation mechanism from the motional average

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(MA) to the static dephasing (SD) mode. In the MA regime, a water molecule experiences

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different H during relaxation because of its diffusion. In contrast, in the SD regime, the field

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fluctuations are negligible. In the MA regime, r2 increases with increasing particle diameter,

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whereas r2 is independent of the particle diameter and remains unchanged in the SD regime.

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Thus, r2 reaches a limiting value in the latter. The MC simulations show that a 13.1-nm

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primary-particle cluster reaches the SD regime at a cluster diameter of 50‒80 nm.31 In addition,

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the SD-regime cluster diameter decreases with increased primary-particle diameter and

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decreased l.31 As shown in Figure 1, the primary Fe3O4 nanoparticles and the cores of the

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Core-Shells have diameters of 17 and 55 nm, respectively, and l ≈ 0 in the core. Thus, the

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Core-Shell diameter falls within the SD regime. The theoretical limit of r2 (r2*) is defined as32

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r2* = (1/1000 Nmole) (Vp / 1000) (π√15 / 9) ∆ω,

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∆ω = 8π / 3√5 γ MS η (rc / ro)3,

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where Nmole, Vp, ∆ω, γ, η, rc, and ro are the number of Fe atoms in a primary nanoparticle, particle

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volume, Larmor frequency difference, gyromagnetic ratio, packing ratio of primary nanoparticles

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in the Core-Shell core, radius of the Core-Shell core, and the overall radius of the Core-Shells,

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respectively. We assumed a packing ratio of primary nanoparticles in the Core-Shell core (η)

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corresponding to close sphere packing. The r2* of the Core-Shells was calculated to be 346 mM-1

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s-1 from equations (7) and (8). Thus, the r2 of the Core-Shells was almost identical to r2*,

(7)

(8)


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indicating that the primary Fe3O4 nanoparticles were closely packed in the core. The above

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findings demonstrate that “tight” clustering also plays a key role in r2 enhancement.

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Furthermore, we compared the r2 of the Core-Shells to that of the 50-nm Fe3O4, which

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were 357 and 315 mM‒1 s‒1, respectively (Figure S6a). Thus, the Core-Shell r2 was 13 % larger

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than that of the 50-nm Fe3O4. This is because the Core-Shells possess higher MS than the 50-nm

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Fe3O4, and r2 increases with increasing MS. The T2-weighted MR images also indicate that the

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Core-Shells reduced the MRI signal compared with the 50-nm Fe3O4 (Figure S6b). These results

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reveal that clustered Fe3O4 nanoparticles exhibit higher r2 than isolated Fe3O4 with particle sizes

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equal to the cluster size.

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2.4. Heat-Generation Power

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We dispersed the Core-Shells and Mag in water (see Experimental Section) and exposed

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them to the AMF in order to investigate their heat-generation power (W), which was estimated in

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terms of the specific absorption rate (SAR). The thermography results demonstrated that the

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Core-Shells generated heat in response to the AMF (Figure 3c). The Core-Shells and Mag

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increased the water T by 20.6 and 5.7°C, respectively, at a concentration of 1 mg/mL under AMF

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exposure for 20 min (Figure 3d). Here,

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SAR = (C/m) (dT/dt) = W/m,

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W = ∫H dM,

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where C is the specific heat capacity of water, m is the weight of the magnetic nanoparticles, and

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dT/dt is the initial slope of T versus the AMF application time (t), or the increase in T for the first

(9)

(10)



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1 min.2 Figure 3e shows that the SAR increased with the increasing crystallite size for the

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Core-Shells (194‒353 W/g) and Mag (57‒128 W/g). The SAR depends on the magnetic

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hysteresis area, as shown in equation (10).9 Tight Fe3O4-nanoparticle clustering increased M at

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the magnetic microfield, HC, and MR (i.e., the magnetic hysteresis area), as shown in Figure 2;

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consequently, the SAR of the Core-Shells was several times larger than that of the Mag. To date,

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conflicting results have been reported regarding the SAR dependence on magnetic-dipole

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interactions,9,34‒39 which is probably due to differences in l values. Previously reported materials

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have l values of at least a few nanometers, whereas l was almost zero in the Core-Shells. The

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present study provides new evidence that the tight clustering of Fe3O4 nanoparticles enhances not

10

only r2 but also the SAR, although previous studies have reported clustering effects on r2 only.

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This finding provides a useful solution for one of the essential problems of magnetic

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hyperthermia treatment, that is, the difficulty in sustaining both the biocompatibility and

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heat-generation power of exothermic bodies.

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2.5. Drug Holding Ability

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The Dox content of the Core-Shells was estimated from the fluorescence intensity due

17

to the presence of Dox in the supernatant collected from the as-prepared Core-Shell-containing

18

solution via centrifugation (see Experimental Section). The Dox-induced supernatant

19

fluorescence intensity was 1.5 at a wavelength of 540 nm (Figure S7). Through substitution of

20

this fluorescence intensity value in the calibration curve (see Experimental Section), the

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Core-Shell Dox content was estimated to be 7.5 nmol. Because 2.8-µmol Dox was added for the


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Core-Shell synthesis, 99.7% of the Dox additive amount was incorporated in the Core-Shells.

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This either equals or surpasses that of the FDA-approved liposomal doxorubicin, Doxil®.40

3 4

2.6. Drug Release

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We remotely controlled the release of Dox from the Core-Shells using the AMF as a

6

trigger. AMF-induced core heat generation softened the shells, as shown in Figure 3f.

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Consequently, the aqueous solution was yellow tinged (Dox colored); thus, the Core-Shells

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released Dox in response to the AMF. To elucidate the drug releasing mechanism, we measured

9

the thermal behavior of Core-Shells using differential scanning calorimetry (DSC). The results

10

indicate that the Shell had a glass transition temperature (Tg) of 46°C and a melting point (Tm) of

11

177°C (Figure S8). As shown in Figure 3d, the Core-Shells increase the T of the entire aqueous

12

solution by 20°C. Therefore, the T of the shell at least reaches Tg. Furthermore, the local

13

temperature surrounding the core potentially reaches the Tm of the polymer shell. In either case,

14

Dox is released from the polymer shell, because Dox is a small molecule that can diffuse from

15

the shell into the media upon limited softening of the shell. If the T of the shell reaches the Tm,

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the Dox will most likely resist the heat generated from the core, because Dox does not

17

decompose at temperatures below 230°C.41

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To simulate in vivo release, we exposed the Core-Shell solution (pH 7.4) to the AMF

19

for 20 min and then observed its behavior without the AMF for 10 days. The supernatant was

20

collected via centrifugation at specific times. The released Dox content (16% after 20 min;

21

Figure 3h) was estimated from the supernatant’s Dox-induced fluorescence intensity. After the

22

AMF removal, the Core-Shells continued to gradually release the remaining Dox, because the 13 ACS Paragon Plus Environment


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shells had already been softened by the AMF-induced core heat generation. Almost all the Dox

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was released over 10 days (Figure 3g). Presumably, the π-π interaction between the melted

3

polymer and Dox inhibits the rapid diffusion of Dox, resulting in the slow diffusion of Dox over

4

10 days. In contrast, in the absence of the AMF application, the Core-Shells released very little

5

Dox (0.8% Dox content leakage; Figure S9). Thus, a single Core-Shell dose may realize

6

continuous in vivo chemotherapy over a period of days to weeks.

7

The pH in tumors (pH 5.8‒7.5) is lower than that in normal tissues (pH 7.3‒7.8). To

8

investigate the Dox release behavior of the Core-Shells in the tumors, we used a solution at pH

9

6.0. The Core-Shell solution was exposed to an AMF for 20 min and then placed out of the

10

influence of the AMF for 10 days. The Core-Shells released 49% of the Dox content due to the

11

20 min AMF application (Figure S10b). This release rate was faster than that at pH 7.4 (16%;

12

Figure 3h). These results suggest that the Core-Shells release Dox efficiently in tumors in

13

response to the AMF compared to in normal tissues.

14 15

2.7. Cellular Uptake

16

To confirm that Core-Shells bind selectively to tumor cells, we cultured human ovary

17

cancer cell line (HAC-2) tumor cells, in which folate receptors were over-expressed,43 in the

18

presence of Core-Shells for 24 h. Bright-field and fluorescent images show that the Core-Shells

19

and Dox were at the same positions (Figure 4a, black and red, respectively); thus, no Dox

20

leakage from the Core-Shells was found in the culture medium. Furthermore, the Core-Shells

21

were localized in the HAC-2 cells. In contrast, unmodified Core-Shells were not located in the

22

HAC-2 cells, as shown in Figure S11. These results suggest that the FA on the Core-Shell 14 ACS Paragon Plus Environment

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1

surfaces binds selectively to the folate receptors of the HAC-2 cells. Furthermore, in order to

2

establish the surface FA of the Core-Shell-induced HAC-2 cell targeting, we investigated

3

whether or not the Core-Shells target folate-receptor negative cells (MCF-7). As shown in Figure

4

S12, fluorescence derived from Dox within the Core-Shells was not observed at the MCF-7 cell

5

position at 24 h after adding the Core-Shells to the culture medium, indicating that the

6

Core-Shells were not localized in the MCF-7 cells. These findings demonstrate that the

7

Core-Shells target HAC-2 cells via binding of the FA on the Core-Shell surfaces to the

8

folate-receptors of the HAC-2 cells.

9 10

2.8. In Vivo MRI-Guided Magnetic Thermochemotherapy

11

Further, we investigated the feasibility of MRI-guided magnetic thermochemotherapy

12

using Core-Shells for mice with intraperitoneal tumors, which were not clearly visible to the

13

naked eye outside the body. We scanned the abdomens using MRI before and after Core-Shell

14

injection (Figure 4b; red dotted square). The Core-Shells were injected intraperitoneally into the

15

mice because in ovarian-cancer treatment, intraperitoneal injection of anticancer drugs often

16

yields a more favorable outcome than intravenous injection.44 The Core-Shell dosage was 4 µg/g,

17

which is also the optimum magnetic hyperthermia dosage, as described below. Twenty-four

18

hours after Core-Shell injection, the MRI signal in the intraperitoneal tumor (Figure 4b; yellow

19

dotted circle) was locally decreased, whereas that for the neighboring tissues remained

20

unchanged. Moreover, time-course MRI revealed that the MRI signal in the intraperitoneal tumor

21

gradually decreased after injection of the Core-Shells (Figure S13). In contrast, in mice injected

22

with unmodified Core-Shells, there was no significant change in the MRI signal of the 15 ACS Paragon Plus Environment


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1

intraperitoneal tumor (Figure S13). The above results demonstrate that FA modification of the

2

Core-Shells is effective at boosting tumor accumulation.

Page 16 of 41

3

The amount of iron in the tumors at various time points was measured by inductively

4

coupled plasma mass spectrometry (ICP-MS). The average iron content in the tumors of

5

non-treated mice was 48.5 ± 4.0 µgFe/g, and those of mice injected with the Core-Shells were

6

53.7 ± 5.2, 54.6 ± 5.5, and 55.2 ± 6.6 µgFe/g at 3, 7, and 24 h after injection, respectively. Thus,

7

the Core-Shells accumulated in the tumors at the concentration of 5.2, 6.1, and 6.7 µgFe/g at 3, 7,

8

and 24 h after injection, respectively.

9

Immediately after MRI, we excised the intraperitoneal tumors and prepared tumor

10

sections. Hematoxylin and eosin (HE) and iron (Berlin blue) staining of the tumor sections

11

revealed that the Core-Shells were abundantly present in the tumor (Figure 4c). Thus, MRI and

12

histological analyses demonstrated that the Core-Shells locally accumulated in the intraperitoneal

13

tumors.

14

Next, to assay the therapeutic efficacy of magnetic thermochemotherapy using the

15

Core-Shells, we applied the AMF to the mice with intraperitoneal tumors for 20 min, 24 h after

16

the Core-Shell intraperitoneal injection. For comparison, we prepared a number of control groups

17

(see Experimental Section). Non-Dox-containing Core-Shells (control material) are designated as

18

Core-Shell (Dox‒). Therefore, for the sake of clarity, in the section on therapeutic efficacy only,

19

the Core-Shell is designated as Core-Shell (Dox+). Applying an AMF after the Core-Shell

20

injection heated the mice abdomens locally by approximately 5.5°C in 20 min (Figures 4d and

21

4e). This temperature increase corresponds to the abdominal skin temperature rather than the

22

intraperitoneal temperature. Therefore, the intraperitoneal temperature is considered to be higher. 16 ACS Paragon Plus Environment

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1

In contrast, AMF application alone and AMF application following Mag injection did not

2

increase the abdominal temperature; in fact, decreases in this temperature were observed because

3

of the influence of anesthesia. The above results demonstrate that the Core-Shells can visualize

4

and heat the tumors at less than half the dose of a clinical iron-oxide-based contrast agent for

5

MRI, Resovist (8.7 µg/g).45

6

We weighed the mice at 7-day intervals post treatment. Furthermore, 42 days post

7

treatment, the tumors were excised and weighed in order to assay the therapeutic effects. AMF

8

application after Mag injection was ineffective for tumor growth inhibition (Figure 4f). Dox

9

administration (chemotherapy alone) and AMF application after Core-Shell (Dox‒) injection

10

(i.e., magnetic hyperthermia alone) inhibited tumor growth to some degree. However, in the

11

chemotherapy group, the body weight decreased significantly in the first 2 weeks and for 42 days

12

after Dox administration, indicating that the Dox caused serious side effects (Figure 4g). In mice

13

exposed to the AMF after Core-Shell (Dox+) injection (magnetic thermochemotherapy), the

14

tumor growth was inhibited significantly compared to chemotherapy alone and magnetic

15

hyperthermia alone (Figure 4f). As noted above, the Core-Shells (Dox+) probably released Dox

16

both during and after AMF application. Thus, the magnetic thermochemotherapy with

17

Core-Shells yielded a favorable outcome compared to chemotherapy alone. Importantly, the

18

mouse body weight increased at the same rate as that of the non-treated mice (Figure 4g). Thus,

19

magnetic thermochemotherapy using Core-Shells yields minimal serious side effects.

20

To investigate the mechanism of magnetic thermochemotherapy using the Core-Shells,

21

we prepared stained tumor sections with HE, TdT-mediated dUTP nick end labeling (TUNEL),

22

and iron as well as undyed tumor sections of both untreated mice and mice 24 h after magnetic

23

thermochemotherapy (Figure 4h and 4i). The Core-Shells were present in the tumors at 24 h after 17 ACS Paragon Plus Environment


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1

treatment (Figure 4i). Dox spread to an area distant from the area in which the Core-Shells were

2

present. Tumor tissues underwent necrosis in the area in which the Dox was present, even if the

3

Core-Shells were not present. Therefore, the necrotic area widened centering around the area in

4

which the Core-Shells were present. However, tumor tissues were surviving in the area which

5

Dox did not reach. Thus, the Core-Shells generated heat in response to the AMF and released

6

Dox, resulting in the necrosis of tumor tissues over a wide area. However, tumor cells were not

7

completely destroyed and the surviving tumor cells regrew. Therefore, the tumors were not cured

8

completely (Figure 4f).

Page 18 of 41

9

Kossatz et al. reported the therapeutic efficacy of two magnetic hyperthermia

10

treatments, one at the onset of evaluation and one at 7 days after the first treatment. The two

11

treatments reduced the tumor volume to about 40% of the initial tumor volume (about 20% of

12

untreated control) at 21 days after treatment. However, afterward, the tumor volume increased by

13

5‒10% in 1 week. In the present study, we evaluated therapeutic efficacy of a one-time

14

treatment. The tumor volume at 42 days after treatment was 39% of the untreated control. As

15

mentioned above, the tumor cells survived slightly at 1 day after treatment (Figure 4i). The

16

regrowth of the surviving tumor cells hindered complete recovery. In the future, we will attempt

17

to establish the optimum frequency and timing of the therapy and enhance the therapeutic

18

efficacy by multiple courses of magnetic thermochemotherapy using the Core-Shells.

19 20

2.9. Toxicity

21

Finally, we evaluated the Core-Shell in vitro and in vivo toxicity through water-soluble

22

tetrazolium salt (WST-1) assays, biochemical assays, and histological analyses. The WST-1 18 ACS Paragon Plus Environment

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1

assays revealed that the Core-Shells show no significant cellular toxicity (Figure S14), indicating

2

that there was no Dox leakage from the Core-Shells. This is supported by the fluorescence image

3

in Figure 4a.

4

In biochemical assays, we used the sera of normal mice and tumor-bearing mice without

5

or with injection of Core-Shells, Dox, or Mag (Figure S15). The sera of tumor-bearing mice

6

subjected to injection of the above materials were collected at 1 hour, 1 day, or 1 week after

7

injection. No significant differences were observed between the total protein (TP), albumin

8

(ALB), blood urea nitrogen (BUN), creatinine (CRE), sodium, potassium, calcium, chlorine,

9

inorganic phosphate (IP), aspartic aminotransferase (AST), alanine aminotransferase (ALT),

10

lactic dehydrogenase (LDH), amylase (AMY), total cholesterol (T-CHO), triglyceride (TG),

11

high-density lipoprotein cholesterol (HDL-C), total bilirubin (T-BIL), and glucose (GLU) values

12

in normal mice, tumor-bearing mice without injection, and tumor-bearing mice injected with the

13

Core-Shells or Mag at any time point. In contrast, the tumor-bearing mice injected with Dox

14

exhibited significantly higher CRE and calcium and lower HDL-C values within 1 day after

15

injection of Dox than normal mice and tumor-bearing mice without injection. These findings

16

indicate that the Dox caused acute kidney toxicity.47 Within 1 day after injection of Dox, the

17

AST, ALT, LDH, and T-BIL values in the tumor-bearing mice injected with Dox were

18

significantly higher than those in the normal mice or tumor-bearing mice without injection,

19

indicating that Dox caused acute liver toxicity.47 Furthermore, the tumor-bearing mice injected

20

with Dox exhibited significantly low GLU values compared to the normal mice or tumor-bearing

21

mice without injection, resulting from adrenal cortex disorders.47 The above results demonstrate

22

that the Core-Shells had no significant toxicity in the mice, and there was little Dox leakage from

23

the Core-Shells in vivo, whereas Dox spreads throughout the body immediately after injection 19 ACS Paragon Plus Environment


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Page 20 of 41

1

and causes acute toxicities for various organs. These Dox side effects caused decreased body

2

weight (Figure 4g). No histological abnormalities such as inflammation or necrosis were

3

observed in the heart, lung, liver, spleen, or kidney 24 h after injection of the Core-Shells (Figure

4

S16).

5 6

3. Conclusions

7

This study demonstrates that reducing the value of l to zero is important for both r2 and

8

SAR enhancement, thus introducing new perspectives for the development of biocompatible

9

materials with a high r2 and SAR. Furthermore, the developed Core-Shells are potentially

10

applicable to MRI-guided magnetic thermochemotherapy, which is a minimally invasive

11

technique. The findings of this study introduce new perspectives regarding the development of

12

materials for MRI, magnetic hyperthermia treatment, and drug-delivery systems. Both

13

conventional and novel iron-oxide-based materials may render theranostics (i.e., techniques

14

fusing diagnosis and treatment) feasible.

15 16

4. Experimental Section

17

Materials. Iron (III) chloride hexahydrate, polyvinyl alcohol (PVA, Mw = 40,000), and 50 nm

18

Fe3O4 were purchased from Sigma Aldrich (MO, USA). Hydrazine monohydrate,

19

pyrrole-3-carboxylic acid, Dox, 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride

20

(EDAC), and N-hydroxysuccinimide (NHS) were purchased from Tokyo Chemical Industry

21

(Tokyo, Japan). Amine- and FA-heterobifunctionalized polyethylene glycol (NH2-PEG-FA, MW

22

= 10,000) was purchased from Nanocs (NY, USA).

23 20 ACS Paragon Plus Environment

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1

Core-Shell synthesis. The Core-Shells were synthesized as follows (Scheme 1). Iron (III)

2

chloride hexahydrate (647 µmol) was added to a PVA solution (0.94 mM), which was then

3

stirred at RT for 1 h. Dox (2.8 µmol) and pyrrole-3-carboxylic acid (0.56 mmol) were added to

4

the solution, which was then stirred at RT for 72 h in the dark to yield polymer nanoparticles.17

5

The pyrrole-3-carboxylic acid was added as a monomer; this monomer was selected because it

6

contains both a pyrrole skeleton that achieved π-πinteraction with the Dox and carboxylic acid,

7

which was useful for surface modification. The polymer nanoparticles encapsulated Dox through

8

π-π interactions. Further, the polymer also contained Fe3+, which was no longer required. Thus,

9

aqueous hydrazine was added to the polymer-nanoparticle-containing solution to reduce a

10

portion of the Fe3+ to Fe2+. Specifically, 5.3‒42 mmol of hydrazine monohydrate was added to

11

the polymer-nanoparticle solution, which was then stirred for 24‒72 h in the dark at 90°C to

12

yield clustered Fe3O4 nanoparticles within the polymer nanoparticles. Thus, unmodified

13

Core-Shells were obtained. Note that the Fe3O4-nanoparticle crystallite size was controlled by

14

adjusting the additive amount of hydrazine and the reaction time (Figure S2). The unmodified

15

Core-Shells were collected from the dispersion via centrifugation (20,000 g, 30 min) and

16

re-dispersed in distillated water. This sequence was repeated three times so as to wash the

17

unmodified Core-Shells. Then, the unmodified Core-Shells were modified with FA and PEG via

18

amidation

19

FA-heterobifunctionalized PEG (NH2-PEG-FA) amino group. For the FA and PEG modification,

20

NHS (20 nmol), EDAC (20 nmol), and FA-PEG-NH2 (20 nmol) were added to the unmodified

21

Core-Shell aqueous dispersion, which was then stirred at RT for 24 h in the dark. (Note that FA

22

binds specifically to the folate receptors that are over-expressed in certain tumor cells, whereas

23

PEG prevents phagocytosis by macrophages.) The Core-Shells were collected from the

between

the

polymer-shell

carboxylic

acid


and

the

amine-

and


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1

dispersion by centrifugation (20,000 g, 30 min) and re-dispersed in distillated water. This

2

sequence was repeated three times so as to wash the Core-Shells.

Page 22 of 41

3 4 5

Core-Shell (Dox‒) synthesis. Core-Shell (Dox‒) was obtained in the same manner as the Core-Shells, with the omission of the Dox addition step.

6 7

Mag synthesis. Mag was obtained using methods similar to the Core-Shell fabrication process,

8

except that the polymer nanoparticle synthesis processes were omitted. Iron (III) chloride

9

hexahydrate (647 µmol) was added to a PVA solution (0.94 mM), which was then stirred at RT

10

for 1 h. Hydrazine monohydrate (5.3‒42 mmol) was added to the solution, which was then

11

stirred for 24‒72 h to yield the Mag. The Mag was collected from the dispersion by

12

centrifugation (20,000 g, 30 min) and then re-dispersed in distillated water. This sequence was

13

repeated three times in order to wash the Mag sample.

14 15

Structural analyses and magnetic properties. The Core-Shell and Mag crystalline phases were

16

analyzed using XRD (Rigaku SmartLab; Rigaku, Tokyo, Japan). The crystallite size was

17

estimated using the (311) reflection of Fe3O4 based on the Scherrer equation. The inorganic and

18

organic phase fractions of the Core-Shell and Mag were measured using differential thermal

19

analysis thermogravimetry (DTA-TG; TG 8120, Rigaku). The sizes and shapes of the

20

Core-Shells and Mag were observed using TEM (H-800; Hitachi, Tokyo, Japan). The

21

hydrodynamic diameters and zeta potentials were measured using dynamic light scattering (DLS;

22

DelsaMax PRO equipped with DelsaMax ASSIST; Beckman Coulter, CA, USA). The magnetic

23

properties were measured using a vibrating sample magnetometer (VSM; VSM-C7-10A; Toei 22 ACS Paragon Plus Environment

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1

Kogyo, Tokyo, Japan) at RT, whereas the ZFC-FC curves were measured using a

2

superconducting quantum interference device (SQUID; MPMS-5P; Quantum Design, CA, USA)

3

at 100 Oe and from 5 to 400 K. The absorption and Fourier-transform infrared (FT-IR) spectra

4

were measured using an ultraviolet-visible (UV-Vis) spectrophotometer (V-570; JASCO, Tokyo,

5

Japan) and an FT-IR spectrometer (Nexus 470; Nicolet, Madison, WI, USA), respectively.

6 7

Relaxivity and in vitro MRI. To prepare phantoms with 3.1 µM iron concentration, we

8

dispersed the Core-Shell or the Mag uniformly in 1 wt% agar aqueous solutions using sonication,

9

which were cooled to 4°C. The phantom T2 values were measured using time-domain nuclear

10

magnetic resonance (NMR) spectroscopy (Minispec mq 20; Bruker, MA, USA). The

11

T2-weighted MR images of the phantoms were acquired using a compact 1.5 T MRI system

12

(MRmini SA; DS Pharma Biomedical, Osaka, Japan) with a 2D-multislice sequence. The

13

parameters were as follows: repetition time (TR) = 2000; echo time (TE) = 69 ms; thickness =

14

1.5 mm; number of averages = 2.

15 16

Core-Shell and Mag heat generation. The Core-Shell and Mag aqueous solutions (1 mg/mL)

17

were placed in the center of a 12-cm-diameter coil and then exposed to an AMF with a frequency

18

(f) of 217 kHz and an H of 8 kA/m (Hf = 1.7 × 109 A m‒1 s‒1); note that an AMF with an Hf

19

value of less than 5 × 109 A m‒1 s‒1 is considered to be harmless. The Core-Shell and Mag

20

temperatures were monitored using an infrared thermography camera (Thermo Gear G100EX;

21

NEC Avio Infrared Technologies, Tokyo, Japan).

22



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Page 24 of 41

1

Drug content. To estimate the Core-Shell Dox content, we collected the supernatant from the

2

as-synthesized Core-Shell solution via centrifugation at 15,000 rpm for 20 min, and subsequently

3

measured the Dox-induced supernatant fluorescent intensity using a fluorescence spectrometer

4

(FP-6600; Jasco, Tokyo, Japan). The Dox content was calculated using the fluorescent intensity

5

(y) vs. concentration (x) curve (y = 10.285x, R2 = 0.9947).

6 7

Drug release. Core-Shells were dispersed in PBS (pH 7.4) or Good's buffer (pH 6.0) at a

8

concentration of 1.75 mg/mL. The Core-Shell solution was placed in the center of the

9

12-cm-diameter coil and exposed to an AMF with f = 217 kHz and H = 8 kA/m for 20 min.

10

Then, the Core-Shell solution was left for 10 days with no AMF interference. The supernatant

11

was collected by centrifugation at set times. To estimate the Dox fraction released from the

12

Core-Shells with and without applied AMF, the supernatant fluorescence intensity was measured

13

over time using a fluorescence spectrometer.

14 15

Cell lines and animals. HAC-2 human ovarian cancer and MCF-7 human breast cancer cell

16

lines were obtained from the RIKEN Cell Bank. The cells were maintained in DMEM

17

(Sigma-Aldrich, St. Louis, MO, USA) containing fetal bovine serum (Gibco) in a humidified

18

atmosphere containing 5% CO2 at 37°C. BALB/c-nu/nu mice (female; 4 weeks of age) were

19

purchased from Japan SLC (Shizuoka, Japan) and maintained in a specific pathogen-free facility

20

at the Center for Animal Research and Education at Nagoya University. All animal experiments

21

were conducted with the approval of the Animal Care Committee of Nagoya University and were

22

in accordance with the guidelines of the Fundamental Academic Research Institution in Japan.

23

The inoculation of HAC-2 cells (1 × 106 cells/animal) was performed through intraperitoneal 24 ACS Paragon Plus Environment

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1

injection into the mice. Eight days after inoculation, the mice were subjected to MRI and

2

therapeutic efficacy evaluation.

3 4

In vivo MRI. T2-weighted MR images of mice with intraperitoneal tumors were acquired before

5

and 3, 7, and 24 h after intraperitoneal injection of the Core-Shells or unmodified Core-Shells (4

6

µg/g) using a compact 1.5 T MRI system with a 2D-multislice sequence. The parameters were as

7

follows: TR = 2000; TE = 69 ms; thickness = 1.5 mm; number of averages = 2.

8 9

Histological analyses. The tumors were excised after the MR images were obtained. The heart,

10

lung, liver, spleen, and kidney were excised 24 h after injection of the Core-Shells. The tumors

11

and organs were immediately immersed in 4% paraformaldehyde at 4°C for 48 h. Fixed tissues

12

were encased in paraffin blocks and sectioned to 2-µm thickness. HE and iron staining of the

13

tissue sections were performed using standard methods.

14 15

In vivo therapeutic efficacy. The therapeutic efficacy of magnetic thermochemotherapy was

16

investigated using mice with intraperitoneal tumors. The mice were assigned randomly to the

17

following seven groups (mice per group; n = 5): No treatment; AMF application; Core-Shell

18

(Dox+) injection; Dox injection (chemotherapy alone); AMF application 24 h after Mag

19

injection; AMF application 24 h after Core-Shell (Dox‒) injection (magnetic hyperthermia

20

alone);

21

thermochemotherapy). The Core-Shells (Dox+), Core-Shell (Dox‒), and Mag were

22

intraperitoneally injected at a dose of 4 µg/g. The Core-Shells (Dox+) contained 0.15 mg Dox

and

AMF

application

24

h

after

Core-Shell


(Dox+)

injection

(magnetic


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1

per gram. Thus, 0.6 µg Dox was contained in 4 µg of the Core-Shells (Dox+). The Dox dose (0.6

2

µg/g) was equivalent to the Dox clinical dose used for patients with ovarian cancer and the Dox

3

content in the Core-Shells (Dox+).48 The AMF application time was 20 min. Before treatment

4

initiation, the mice were anesthetized with isoflurane (DS Pharma Animal Health, Osaka, Japan)

5

using an anesthesia inhalation apparatus (NARCOBIT; Natsume Seisakusho, Tokyo). The mice

6

were placed in the 12-cm-diameter coil, through which the AMF (217 kHz, 8 kA/m) was applied.

7

Each treatment was performed just once at the onset of the evaluation. Forty-two days after

8

treatment, the intraperitoneal tumors were excised from the mouse abdomens and weighed.

9

Furthermore, the body weights were measured at 7-day intervals post treatment. The tumor and

10

body weights for the different treatment groups were statistically compared. P < 0.05 was

11

considered to be statistically significant.

Page 26 of 41

12 13

Cellular toxicity. The Core-Shells were added to the culture media in the presence of HAC-2

14

or MCF-7 cells. The survival rate of the cells at 24 h after adding the Core-Shells to the culture

15

media was evaluated via a WST-1 assay. HAC-2 or MCF-7 cells cultured in the absence of the

16

Core-Shells were used as controls.

17 18

Biochemical assays. The Core-Shell, Mag and Dox toxicities were evaluated through

19

biochemical assays using serum samples from mice. The sera were collected at 24 h after

20

intraperitoneal injection of the Core-Shells (4 µg/g), Mag (4 µg/g), or Dox (0.6 µg/g) (n = 5 mice

21

per group). Normal mice and tumor-bearing mice without injection were used as controls. The

22

biochemical components (TP, ALB, BUN, CRE, sodium, potassium, calcium, chlorine, IP, AST, 26 ACS Paragon Plus Environment

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1

ALT, LDH, AMY, T-CHO, TG, HDL-C, T-BIL, and GLU) of the serum were determined using

2

an automated biochemical analyzer (Hitachi 7180; Hitachi).

3 4

Quantity of iron in the tumors. The tumors were excised from mice at 3, 7, and 24 h after

5

injection of the Core-Shells. The quantity of iron in the tumors was measured by ICP-MS

6

(ICPM-8500; Shimadzu, Kyoto, Japan). The pretreatment of tumors for ICP-MS measurement

7

was performed according to the literature.49 In brief, tumors were digested in a 1:3 volumetric

8

mix of trace metal-grade hydrochloric acid and nitric acid. Digestions were incubated at 70°C for

9

up to 12 hours until it appeared all tumors had gone into solution. Tubes were weighed

10

post-digestion to determine the mass of the final digested sample.

11 12

Supporting Information Available. Additional figures. This material is available free of charge

13

via the Internet at http://pubs.acs.org.

14 15

AUTHOR INFORMATION

16

Corresponding Author

17

* Address correspondence to [email protected].

18 19

ACKNOWLEDGMENTS

20

We are grateful to the Center for Animal Research and Education (CARE) and the Technical

21

Center at Nagoya University. This work was supported by a Grant-in-Aid for Young Scientists

22

(A) (26709050) and a Grant-in-Aid for Exploratory Research (15K14146) from the Japan 27 ACS Paragon Plus Environment


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1

Society for the Promotion of Science (JSPS). This work was also supported by the Kato

2

Foundation for the Promotion of Science. Additionally, the work was partly supported by a

3

Health Labor Sciences Research Grant from the Ministry of Health Labor and Welfare and a

4

Long-range Research Initiative (LRI) grant from the Japan Chemical Industry Association

5

(JCIA).

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Scheme 1. Core-Shell synthesis scheme. 1



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1 Figure 1. Core-Shell and Mag characterization. (a) TEM images of Core-Shells. (b) Magnified view of Figure 1a. (c) TEM image of Mag. (d) Core diameter, (e) shell thickness, and (f) overall diameter distributions of Core-Shells. 2


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1 Figure 2. Core-Shell and Mag magnetic properties. (a) ZFC-FC and (b) magnetization curves. (c) Magnified view of the origin in Figure 2b. Relationship between crystallite size and (d) HC, (e) MR, and (f) MS in Core-Shells and Mag. 2



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Figure 3. Relaxivity and heat-generation power of Core-Shells and Mag. (a) T2-weighted MR images of phantoms containing Core-Shells or Mag. (b) r2 of Core-Shells and Mag. (c) Thermal images of a Core-Shell aqueous solution before and after AMF application. (d) Changes in the temperature of an aqueous solution containing Core-Shell or Mag with AMF application time. (e) Relationships between SAR values of Core-Shells and Mag and crystallite size. (f) Photographs of Core-Shell aqueous solution before and after AMF application. (g) Dox-releasing behavior of Core-Shells at pH 7.4. (h) Magnified view of the first 60 min of behavior shown in Figure 3g. 1


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1 Figure 4. Cellular uptake and in vivo MRI-guided magnetic thermochemotherapy. (a) Core-Shell cellular uptake: (left to right) bright field, fluorescence, and overlay images. (b) 39 ACS Paragon Plus Environment


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Photograph and T2-weighted MR images of mice subjected to Core-Shell injection. The red square in the photograph indicates the MRI scanning area, whereas the yellow circles in the MRI images indicate tumor sites. (c) HE- and iron-stained tumor sections for mice with and without Core-Shell injection: The area stained blue in the iron-stained sections indicates the area in which the iron ions, i.e., the Core-Shells, are present. (d) Thermal images of mice with and without Mag or Core-Shell injection under influence of AMF: The white circles indicate abdominal tumor sites. (e) Change in mouse abdominal temperature with and without Mag or Core-Shell injection under influence of AMF. (f) Weights of intraperitoneal tumors excised from mouse abdomens 42 days after treatment: *p < 0.05, **p < 0.01, ***p < 0.005. (g) Changes in body weight post treatment: *p < 0.05, **p < 0.01, ***p < 0.005. HE-, TUNEL-, and iron-stained as well as undyed tumor sections of untreated mice (h) and mice 24 h after magnetic thermochemotherapy (i). The surviving tumor tissues are enclosed in yellow dotted line. In the TUNEL-stained sections, the areas stained brown and blue indicate dead and surviving tissues, respectively. In the iron-stained sections, the area stained blue indicates the area in which the iron ions, i.e., the Core-Shells, are present. In the undyed sections, the areas stained red indicate the areas in which Dox is present. Scale bar is 800 µm. 1


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