Externally Triggered Heat and Drug Release from Magnetically

Jan 9, 2019 - Shanmugam, Cuthbert, Kowalewski, Boyer, and Matyjaszewski. 2018 51 (19), pp 7776–7784. Abstract: We report a novel photoiniferter ...
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Article Cite This: ACS Appl. Polym. Mater. XXXX, XXX, XXX−XXX

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Externally Triggered Heat and Drug Release from Magnetically Controlled Nanocarriers Eric G. Fuller,†,∥ Hao Sun,‡,⊥,∥ Rohan D. Dhavalikar,§ Mythreyi Unni,§ Georg M. Scheutz,‡ Brent S. Sumerlin,*,‡ and Carlos Rinaldi*,†,§ †

J. Crayton Pruitt Family Department of Biomedical Engineering, University of Florida, Gainesville, Florida 32610, United States George & Josephine Butler Polymer Research Laboratory, Center for Macromolecular Science & Engineering, Department of Chemistry, University of Florida, PO Box 117200, Gainesville, Florida 32603, United States § Department of Chemical Engineering, University of Florida, Gainesville, Florida 32603, United States

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ABSTRACT: Nanoscaled drug carriers have been developed to accumulate in tumors and release a drug cargo either passively or in response to a local stimulus. However, strategies that rely on passive release or response to a local stimulus do not allow for spatial and temporal control of drug delivery. These limitations motivate interest in drug delivery platforms that release cargo in response to an external stimulus. This contribution describes magnetically controlled nanocarriers (MCNCs) that release heat and a drug cargo in response to an applied alternating magnetic field (AMF). The MCNCs consist of a hydrophobic core of superparamagnetic iron oxide nanoparticles that release heat in response to an AMF and a thermoresponsive polymer that releases a molecular cargo via breakage of thermally labile Diels−Alder (DA) bonds. The nanocarriers are coated with polyethylene glycol-block-polylactic acid (PEG4.9kD-PLA6.0kD) block copolymer to confer colloidal stability and water solubility. The MCNCs are assembled through flash nanoprecipitation, a rapid approach to making nanoparticles that is scalable and provides control of size and composition. Release experiments show that application of an AMF results in on-demand heat and drug release. The AMF-actuated release ceases when the field is turned off, and multiple applications of AMF result in programmable release. The amount of release is tunable via the AMF field strength and can be spatially controlled using selection magnetic field gradients. These results suggest that a potent combination of magnetic hyperthermia and drug release can be actuated in a desired region. KEYWORDS: magnetic nanoparticles, magnetically controlled nanocarriers, externally triggered release, drug delivery, thermally labile bond, flash nanoprecipitation, spatially selective drug release

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INTRODUCTION

improvements in overall patient survival when compared to the free drug.4−7 A major limitation of clinically used nanoparticles is the lack of spatial and temporal control of drug release. For most nanoparticles, drug release occurs through a passive mechanism, with the rate of release modified by nanoparticle properties and composition. In such cases, once nanoparticles are administered, they begin to release their cargo in a manner that is not specific to the desired region. Nanoparticles that release their cargo too quickly show undesirable toxicities, and results are similar to the use of free drug. On the other hand, nanoparticles which release cargo slowly may have decreased toxicity but show reduced efficacy.8,9 There is also in vivo evidence that particles which release drug quickly have a

Nanoscaled drug carriers have been investigated for several decades due to their potential in the field of drug delivery.1 The overarching goal has been to develop drug delivery vehicles that carry their cargo to the desired site and release it there. This would provide the benefits of having a more potent effect in the desired region and reducing side effects from the drug in off-target areas.2,3 In addition, nanoparticles have the capability of shielding drugs from premature degradation and solubilizing hydrophobic drugs that cannot be administered intravenously. Despite these desirable properties, clinical translation of nanoparticles has been modest.4 Doxil, a liposomal form of doxorubicin, has shown decreased cardiotoxicity and other reduced side effects when compared to free doxorubicin administration in cancer patients.5 However, as a whole, nanoparticles in current clinical use still show concerning toxicities and do not show significant © XXXX American Chemical Society

Received: November 5, 2018 Accepted: December 26, 2018

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DOI: 10.1021/acsapm.8b00100 ACS Appl. Polym. Mater. XXXX, XXX, XXX−XXX

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ACS Applied Polymer Materials Scheme 1. Synthesis of a Thermally Labile Pendent Fluorescein Containing Polymer

be used to control drug release.22,43−45 Permeabilization of a liposomal membrane46−48 and breakage of a thermally labile bond, such as Diels−Alder bonds and aliphatic azo linkers,23−25 can also be utilized for triggered release. Although these approaches have shown progress in vitro, they face the challenge of keeping passive release rates low while maximizing the triggered release. Another challenge has been poor colloidal stability and low drug loading capacity of the carriers. In this study, a new platform called magnetically controlled nanocarriers (MCNCs) was developed, combining SPIONs and engineered polymers with thermally labile drug conjugates. This allows for the formation of stable carriers with the ability to control size, drug loading efficiency, drug loading content, and rates of passive and triggered release. MCNCs are assembled through flash nanoprecipitation (FNP), which allows for rapid, scalable preparation as well as tunable size and drug loadings.49−51 During FNP, the hydrophobic SPIONs and engineered polymer are combined with a poly(ethylene glycol)-block-poly(lactic acid) (PEG4.9kDPLA6.0kD) block copolymer which coats them, providing colloidal stability and water solubility to the carriers. The engineered polymers are made of a hexyl methacrylate-comaleimide methacrylate (HMA-MalMA) backbone, with maleimide moieties in side chains and furan-modified cargo loaded via a Diels−Alder mechanism (Scheme 1). The shelf life of the furan-maleimide-based Diels−Alder bond is long at body temperature, but the retro-Diels−Alder (rDA) reaction occurs at elevated temperatures, promoting cargo release.52,53 Therefore, it was hypothesized that with PEG−PLA and HMA-MalMA polymers, carriers can be made which are stable and retain cargo for long periods of time at body temperature yet release the cargo at elevated temperatures. In addition to synthesizing and characterizing the HMAMalMA copolymer, SPIONs, and MCNCs, this study demonstrates the tunable rate of release from the MCNCs at various temperatures and under passive or magnetically triggered conditions. Although fluorophore cargo is used, this article reports a platform for drug loading and release which has potential use for many actual drugs. A small molecule drug

stronger effect, whether it be a desired antitumor effect or an undesired side effect, than similar particles which release drug slowly, even when the quick-releasing particles deliver less total drug in the tumor or healthy tissues.10 Therefore, the ideal case would be to have particles which release drug very slowly while they are in off-target regions, yet very quickly when they are in the desired region. For these reasons, methods of triggering drug delivery via various stimuli that can give both spatial and temporal control have been heavily researched.1,3,11 The ability to trigger drug release on-demand in desired regions of the body could revolutionize chemotherapy and other drug treatments where dose-limiting side effects are a major concern. Some stimuli investigated for this purpose are environmental stimuli, such as temperature,12−15 pH,13,14,16,17 or the presence of certain enzymes18−20 in the region where release is desired.1 Other stimuli, such as alternating magnetic fields,21−25 ultrasound,26−28 and light,29,30 are external to the body. Recent advancements in materials science and chemistry are enabling the creation of externally triggered nanoparticles, which have the advantages of decreased variability and ondemand control of the extent of drug release.11 In the presence of an alternating magnetic field (AMF), superparamagnetic iron oxide nanoparticles (SPIONs) release heat, which can be used to cause drug release through several mechanisms.11,31 This concept is often referred to as magnetically triggered drug delivery (MTDD). AMF is a favorable external stimulus for drug release for several reasons. Negligible tissue attenuation of AMF occurs, meaning that release can be actuated in areas deep in the body.32,33 SPIONs with high heat dissipation rates have been synthesized, and SPIONs can be imaged via magnetic resonance imaging34 or magnetic particle imaging (MPI).35−38 Instrumentation already exists to generate AMFs which are suitable for MTDD.39 Furthermore, selection magnetic field gradients can be used for spatially selective heating, with cm−mm precision.33,36,40−42 The heat generated in an AMF by magnetic nanoparticles has been used to trigger drug release via several mechanisms. Conformational/solubility changes to a thermosensitive polymer possessing lower critical solution temperature can B

DOI: 10.1021/acsapm.8b00100 ACS Appl. Polym. Mater. XXXX, XXX, XXX−XXX

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ACS Applied Polymer Materials is modeled here, but large biological cargo such as proteins have been used for MTDD54 as well and could encapsulate into our platform through the use of inverse FNP, in which case the loading efficiency could be controlled by the feed ratio of protein and coating polymer.55 The platform is highly tailorable as many hydrophobic polymers could be encapsulated into stable nanocarriers. Although not a part of this study, the outer layer of PEG can be further functionalized with a targeting moiety specific to the desired application.56 Previous platforms mostly rely on noncovalent drug loading approaches, which suffer from burst release and high passive release rates. While a few recent publications23−25 report the use of thermally labile linkers, most previous approaches have focused on conjugating one drug or one dye onto the chain end of each polymer. This study reports the first example of loading drugs onto polymer side chains via Diels−Alder chemistry, resulting in tunable drug content/loading.

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Figure 1. 1H NMR spectrum of polymer−dye conjugate.

RESULTS AND DISCUSSION Synthesis of a Thermally Labile Pendent Fluorescein Containing Polymer. A hydrophobic random copolymer of furan-protected maleimide methacrylate (Fur-MalMA) and hexyl methacrylate (HMA) was prepared via conventional radical polymerization in the presence of radical initiator AIBN at 60 °C (Scheme 1) and characterized by NMR spectroscopy (Figures S1 and S2). The chemical structure and composition of the resulting copolymer poly(Fur-MalMA-co-HMA) were verified by NMR spectroscopy and gel permeation chromatography (GPC). On the basis of the 1H NMR spectrum (Figure S3), the molar ratio of HMA to Fur-MalMA was 93:7 in the polymer, which is slightly lower than the feed ratio (HMA:FurMalMA = 90:10). GPC analysis indicated that the numberaverage molecular weight of polymer was 57 kDa with a molar mass dispersity index of 2.6 (Figure S4). A broad and multimodal GPC trace was observed, which is expected given that the polymer was obtained by conventional radical polymerization. Maleimide deprotection was carried out at 110 °C to furnish poly(MalMA-co-HMA) that contains pendent free maleimide functionalities. The efficiency of the deprotection was quantitative, according to 1H NMR analysis, where the complete disappearance of the Diels−Alder bridgehead protons at 5.2 ppm and the appearance of the maleimide vinyl protons signal at 6.7 ppm were observed (Figure S5). Since the hydrophobic polymer contains abundant maleimide functional groups, we reasoned that installing a furan on the small molecule drug would allow for facile covalent conjugation to the poly(MalMA-co-HMA) side chains. Furan functional fluorescein (Fur-Fluo) was synthesized via carbodiimide chemistry (Scheme S1 and Figure S6). After that, a Diels−Alder reaction between poly(MalMA-co-HMA) and Fur-Fluo resulted in the formation of the polymer@Fluo conjugate (Scheme 1). The successful synthesis was confirmed by 1H NMR spectroscopy showing both signals derived from polymer and fluorescein. More importantly, a signal ascribed to Diels−Alder linker protons at 5.1 ppm was observed, confirming that polymer and cargo were covalently bonded rather than being physically mixed (Figure 1). On the basis of the 1H NMR spectrum of polymer−dye conjugate, 50% of the maleimide groups were functionalized with Fur-Fluo, suggesting that 16.5% weight loading of model drug in the polymer@ Fluo conjugate was achieved. Although maximizing drug loading was not an aim of the current study, the drug loading of the polymer is tunable by changing the feed ratio of drug

monomer to nonfunctional monomer during the polymerization process. The drug monomer to nonfunctional monomer feed ratio can likely be increased in the feed until the steric hindrance of the abundant drug monomer becomes a limiting factor. Future studies are suggested to determine the optimal feed ratio. To show thermal responsiveness of the polymer@Fluo conjugate, the conjugate was heated in toluene at various temperatures, including 25, 60, and 90 °C. According to fluorescence spectroscopy (Figures S7 and S8), the release kinetics of Fur-Fluo were dependent on the applied temperatures. When the temperature was 25 °C, no release was observed. As the temperature was further increased, both the rate and extent of drug release were increased as a function of time, suggesting the promise of using this polymer as a heatresponsive drug delivery system. Iron Oxide Nanoparticle Synthesis and Characterization. Iron oxide nanoparticles were synthesized via the thermal decomposition approach, which has become a common route to obtain magnetic nanoparticles with narrow size distributions. An iron oleate precursor was thermally decomposed at 350 °C in a semibatch synthesis reactor.57 To obtain particles with similar physical and magnetic diameters, molecular oxygen was added to the thermal decomposition synthesis at slightly above stoichiometric ratio with respect to iron.58 Figure 2 shows the distributions for the magnetic, physical, and hydrodynamic diameters of the synthesized SPIONs. The magnetic diameter was 16.5 ± 1.05 nm, found by measuring the equilibrium magnetization of the sample in a superconducting quantum interference device (SQUID) and fitting the curve to the Langevin function weighted using a lognormal distribution, as has been previously reported.58 Figure 3 shows this equilibrium magnetization curve with a very high initial slope and no observation of hysteresis. The physical diameter plotted in Figure 2 was found by transmission electron microscopy (TEM, Figure 4) to be 18 nm, slightly larger than the magnetic diameter. The hydrodynamic diameter from dynamic light scattering (DLS) measurements, reported as the volume-weighted distribution in Figure 2, was 25 nm. The hydrodynamic diameter is expected to be larger than the physical diameter because the physical diameter only includes the iron oxide core, whereas the hydrodynamic diameter is determined by the size of the core and the oleic C

DOI: 10.1021/acsapm.8b00100 ACS Appl. Polym. Mater. XXXX, XXX, XXX−XXX

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ACS Applied Polymer Materials

nanoprecipitation (FNP), which produces nanoparticles rapidly and scalably. The SPIONs, fluorescein-pendent polymer, and PEG4.9kD-PLA6.0kD were dissolved in an organic phase of tetrahydrofuran, rapidly mixed with water, and magnetically filtered to give the final, purified MCNCs. These MCNCs were characterized for physical, hydrodynamic, and magnetic size as well. Figure 5 shows the physical and

Figure 2. Magnetic (Dm), physical (Dp), and hydrodynamic (Dh) diameters of SPIONs. The hydrodynamic diameter is expected to be larger than the physical diameter because the physical diameter only includes the iron oxide core, whereas the hydrodynamic diameter is determined by the size of the core and the oleic acid coating.

Figure 5. Physical (Dp) and hydrodynamic (Dh) diameters of MCNCs. Dh was plotted as the log-normal volume-weighted DLS intensity and has an arithmetic mean diameter of 90 nm. Dp was obtained from analysis of TEM images and is similar to Dh with an arithmetic mean of 115 nm.

hydrodynamic diameter distributions for the MCNCs. The physical diameter, obtained from TEM measurements, had an arithmetic mean of 115 nm (Figure 6). The hydrodynamic Figure 3. Equilibrium magnetization curve of the SPIONs showing a high initial slope and no hysteresis.

Figure 6. TEM images of MCNCs showing the presence of a SPION cluster inside a polymer shell. (Inset) Magnified image of a single MCNC. The high magnification MCNC example shows a coating around the outer layer of SPIONs, which we attribute to the block copolymer.

Figure 4. TEM image of SPIONs. This image confirms the synthesis of semispherical nanoparticles approximately 18 nm in diameter.

diameter was found through DLS measurements and showed a similar distribution as the physical diameter, with an arithmetic mean diameter of 90 nm. Although the distributions in Figure 5 are fairly similar, one potential reason that the MCNC TEM mean diameter was larger than the hydrodynamic diameter could be due to the slight spread of the MCNCs when they are dried and imaged on a TEM grid.

acid coating. The specific absorption rate (SAR) and intrinsic loss power (ILP) values found by subjecting the SPIONs to an AMF of 59.6 kA/m and 346 kHz were 558 W/gFe and 0.91 nHm2/kg, respectively, which are consistent with ranges reported in literature.58,59 Preparation and Characterization of Magnetically Controlled Nanocarriers. MCNCs were prepared by flash D

DOI: 10.1021/acsapm.8b00100 ACS Appl. Polym. Mater. XXXX, XXX, XXX−XXX

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ACS Applied Polymer Materials Figure 6 shows a TEM image of multiple MCNCs along with one MCNC at high magnification. From the TEM, it is clear that MCNCs consist of a core composed of multiple SPIONs. Due to its hydrophobic nature, the fluorophoreloaded polymer is presumed to be in the hydrophobic space around the SPIONs. The release test results in Figures 7, 8,

Figure 9. Spatial control of release using magnetic field gradients. (A) Insulated samples at position A and position B placed in an AMF producing coil reach similar temperatures on exposure to an AMF for 1 h. (B) Insulated samples located in the field free region (FFR) at position B reach a higher temperature than samples located in the saturated region at position A on exposure to an AMF for 1 h and in the presence of a magnetic field gradient. A magnetic field gradient of 1.27 T/m is generated by two opposing NdFeB permanent magnets. (C) In the absence of a field gradient, samples at position A and B show similar release (no significant difference, p > 0.3). (D) In the presence of a field gradient, samples at position B release more cargo than samples at position A since magnetic heating is suppressed at position A (significant difference, p < 0.01). Mean ± standard deviation is plotted, n = 3.

Figure 7. Release from MCNCs stops when AMF is switched off after application for 1 h. No significant release was observed over the course of the next 12 h. The “w/AMF” sample was exposed to 1 h of AMF and then kept at room temperature as fluorescence was monitored over time. The “w/o AMF” sample was kept at room temperature. Mean ± standard deviation is plotted, n = 3.

MCNC concentration. While this amount is low, achieving the maximum drug loading was not the aim of the current study and further work could improve this loading by changing the drug loading of the fluorescein pendent polymer as mentioned above and by changing the FNP conditions to include higher concentrations of fluorescein pendent polymer during MCNC formation. Magnetically Triggered Release Test Results. To confirm that the amount of release from the MCNCs is tunable with temperature, MCNCs in 25/75 v/v methanol/ PBS release media were subjected to 37 °C external heating or heating in an AMF at 45, 52.5, and 60 °C for 1 h, with the results shown in Figure S12. Release percentages were calculated by dividing the relative fluorescent units (RFU) from released fluorophore for a given sample by the RFU seen for complete release of fluorophore. Complete release was assessed by subjecting the MCNCs to 90 °C for 10 h. Heating at 37 °C for 1 h resulted in less than 1% release. There was a clear trend of greater release as temperature increased due to samples being exposed to AMF. The 60 °C AMF sample released nearly 4% of the cargo in one hour, approximately 5 times higher than the amount released from the 37 °C sample. Given a concentration of encapsulated cargo of 0.0135 mg/ mL in 10 mL of cell culture media, similar to the work in our study, a release of 4% of doxorubicin, for example, would result in a concentration above 1.0 μM in the media, which is above most reported IC50 values for doxorubicin. In addition, control MCNCs without any loaded cargo were added to MDA-MB231 cells at a relatively high concentration of 2 mg/mL, and low toxicity was seen (Figure S10). Another release experiment was designed to determine the limiting step of release from the MCNCs. Before the fluorophore cargo is released out of

Figure 8. Release from MCNCs can be triggered multiple times to provide time-specific release. Minimal passive release is observed between two field switch on time points. Mean ± standard deviation is plotted, n = 3.

and 9 also confirm that the fluorophore-loaded polymer was encapsulated. The high magnification MCNC example also shows a coating around the outer layer of the SPIONs, which we attribute to the PEG4.9kD-PLA6.0kD block copolymer. Given that the SPIONs and fluorophore-loaded polymer are insoluble in water, the water solubility of the MCNCs provides additional evidence that the block copolymer coated the outside of the MCNCs. An equilibrium magnetization curve of the MCNCs showed a high initial slope and no observation of hysteresis, similar to the equilibrium magnetization curve of the SPIONs (Figure S9). This evidence, along with the use of magnetic filtration to purify the MCNCs, is consistent with the TEM images that indicate that SPIONs were encapsulated inside of the MCNCs. Total drug loading was approximately 0.42% by mass. This was found by dividing the concentration of released drug from samples exposed to 90 °C for 10 h (total release) by the overall E

DOI: 10.1021/acsapm.8b00100 ACS Appl. Polym. Mater. XXXX, XXX, XXX−XXX

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ACS Applied Polymer Materials MCNCs, it first must break free from the polymer via the rDA reaction and then diffuse out of the MCNC. Since the hexyl methacrylate-based hydrophobic polymer that forms the core of the MCNC has a low glass transition temperature (Tg) around −15 °C, enhanced chain mobility should allow for fast diffusion out of the core. Therefore, it was hypothesized that the rDA reaction was the limiting step in this process. To investigate this, an experiment was performed with a sample exposed to AMF at 60 °C for 1 h, and then its release was monitored over the next 13 h. The sample and the control were both kept at room temperature (23 °C). The results in Figure 7 show that the AMF did cause significant release, as the measured fluorescence emission for the AMF-treated sample was much higher than the control samples. However, during the 12 following hours after the AMF had been turned off, no increase in fluorescence was seen. These results suggest that at elevated temperatures, cargo can diffuse through the PEG−PLA layer, but this does not happen readily at lower temperatures. This agrees with the Tg of the PLA, which is near 50 °C for the PLA used in the study (a racemic mixture of the “D” and “L” lactide). The results also suggest that the limiting step of drug release is indeed the breaking of the Diels−Alder bond between the cargo and the polymer. More importantly, this shows that in the AMF, the release of cargo from the MCNCs is immediate and little release occurs in the absence of the AMF, reducing the concern of diffusion-driven passive release when AMF was removed. Due to the stability of the MCNCs, which showed negligible size change via DLS for 1 month while suspended in PBS and kept at 4 °C (Figure S11), it was plausible that the same sample of MCNCs could be subjected multiple times to AMF and show a sustained capacity for triggerable release. Nanoparticles with this property would be advantageous in a clinical setting because clinicians could determine how many sessions of triggered release a patient could undergo based on the response to the initial release. Figure 8 shows the results of a test in which one set of samples was exposed to body temperature (37 °C) via external heating for 11 h, and another set was exposed to alternating periods of 37 °C external heating and exposure to an AMF. The latter set was pulsed with an AMF at times 0, 5, and 10 h for a duration of 1 h as shown by the on/off designation in Figure 8. During these AMF pulses, the temperature of the samples rose to 50 °C. The control samples held at 37 °C had a linear release profile at a rate of 0.34% release per hour. When considered with the data in Figure 7, we suggest that passive release happened slowly at 37 °C but was negligible at room temperature of 23 °C. The AMF samples released 3.5 times more fluorophore cargo than the control samples during the first AMF pulse, as can be seen by the steeper upward slope of the AMF sample. Further AMF pulses at 5 and 10 h also showed greater release than the control and confirmed that multiple release events are feasible. With recent progress in spatially controlling the heating of SPIONs using magnetic field gradients,33,36,40−42 we evaluated if the amount of cargo released from MCNCs could be spatially controlled. The results are shown in Figure 9. When vials were placed 5.5 cm apart in positions A and B inside a coil and release was actuated via AMF as illustrated in Figure 9a, the amount of release shown in Figure 9c was similar for both positions. The temperature recorded by the fiber optic probes in vials at both positions was also similar. However, when two opposing NdFeB permanent magnets were placed outside the

coil in a configuration as illustrated in Figure 9b to create a field free region (FFR) at position B, the samples in position B showed 40% more release than those in position A (Figure 9d). A temperature difference of 5 °C was observed between the samples at position B and position A due to suppression of magnetic heating at position A because of the presence of a strong bias magnetic field. This proof-of-concept test shows that the amount of release from MCNCs can be spatially controlled. As shown in a previous theoretical study by Dhavalikar and Rinaldi,33 increasing the strength of the field gradient can tune the region of heating down to the millimeter scale. This concept could be applied to further tune the region of drug release, preventing off-target release of drugs and reducing their potential side effects.

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CONCLUSIONS We report the preparation and characterization of MCNCs loaded with cargo that is covalently conjugated through thermally labile bonds and that can be released when the temperature is elevated due to heat dissipation by SPIONs in an AMF. This AMF-triggered release is tunable with temperature increase, and little release occurs when the particles are not subjected to AMF. The release can be triggered multiple times, and magnetic field gradients can be used to spatially control the region where release is desired to prevent drug release in off-target sites. With these attributes, this platform shows potential as an externally controlled nanoparticle carrier suitable for spatiotemporally controlled drug release.

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MATERIALS AND METHODS

Materials. 2-(2-Hydroxyethyl)-3a,4,7,7a-tetrahydro-1H-4,7-epoxyisoindole-1,3(2H)-dione (Fur-MalOH) was synthesized according to a previous procedure.60 Methacryloyl chloride (97%), furfurylamine (98%), methanol (99.8%), triethylamine (TEA, 99.5%), dichloromethane (DCM), N,N-dimethylformamide (DMF, 99.8%), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC·HCl, 98%), ethyl acetate (99.8%), and ethanolamine (99.5%) were purchased from VWR. Furan (>99%), hexyl methacrylate (98%), fluorescein, 2,2′-azobis(2-methylpropionitrile) (AIBN, 98%), dimethyl sulfoxide (DMSO, 99.8%), 2-(dimethylamino) pyridine (DMAP, 97%), 1,4-dioxane (99.8%), maleic anhydride (99%), N,N-dimethylacetamide (DMAc, 99.9%), phosphate-buffered saline (PBS), and Dulbecco’s modified eagle medium (DMEM) were purchased from Sigma-Aldrich. Toluene (99.8%, Sigma) was dried over molecular sieves. Tetrahydrofuran for polymer synthesis (THF, 99.8%, Sigma) was dried over calcium hydride and freshly distilled under reduced pressure. Monomers were passed through basic alumina columns to remove inhibitors and acidic impurities prior to polymerization. AIBN was recrystallized from methanol twice before polymerization. All other materials were used as received unless otherwise noted. PEG4.9kD-b-PLA6.0kD was purchased from Evonik Industries. Tetrahydrofuran (THF, 99.9%) used in flash nanoprecipitation, methanol (MeOH 99.9%) used in release tests, and lithium chloride (LiCl, 99%) were purchased from Acros Organics. Triton X-100 was purchased from Fisher Scientific and diluted to 1% in DIW before use. Amicon-ultra 0.5 mL 100 kDa centrifugal filters were purchased from EMD Millipore. An alternating magnetic field was generated in a custom-built coil with an Ambrell EASYHEAT 8130L1 10 kW induction heater. The temperature of the samples exposed to AMF was monitored by a fiber-optic temperature probe from Neoptix. Two NdFeB permanent magnets (BZ0Z0 × 0-N52) were purchased from K&J Magnetics to generate a field free region in the spatially controlled release experiments. Methods. TEM Analysis. Ten microliters of sample was applied onto a Formvar coated 200-mesh Cu grid. The grids were observed F

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ACS Applied Polymer Materials on a Hitachi H7000 microscope operating at 100 kV. The images were recorded with a slow-scan CCD camera (Veleta 2k × 2k). GPC Analysis. Molecular weight and polydispersity were determined by size exclusion chromatography in DMAc with 50 mM LiCl at 50 °C and a flow rate of 1.0 mL min−1 (Agilent isocratic pump, degasser, and autosampler; columns, Viscogel I-series 5 μm guard + two ViscoGel I-series G3078 mixed bed columns; molecular weight range 0−20 × 103 and 0−100 × 104 g mol−1). Detection consisted of a Wyatt Optilab T-rEX refractive index detector operating at 658 nm. Relative molecular weights were obtained through calibration with poly(methyl methacrylate) (PMMA) standards of molecular weights ranging from 9.88 × 105 to 602 g/mol. NMR Analysis. 1H NMR and 13C NMR spectra were recorded in CDCl3 or CD2Cl2 and MeOD using an Inova 500 MHz spectrometer. DLS Analysis. DLS measurements were performed using a Brookhaven Instruments 90Plus/BI-MAS, operated at room temperature. Synthesis of 2-(1,3-Dioxo-3a,4,7,7a-tetrahydro-1H-4,7-epoxyisoindol-2(3H)-yl)ethyl Methacrylate (Fur-Maleimide Methacrylate). Fur-MalOH (10.0 g, 47.8 mmol) and TEA (10.2 mL, 71.7 mmol) were suspended in dry THF (100 mL) in an ice bath. Methacryloyl chloride (7.17 mL, 71.7 mmol) was slowly added over 30 min. Following the addition, the reaction mixture was stirred at 25 °C for 12 h. The reaction was monitored by thin layer chromatography and quenched when the reaction was completed. After removal of the TEA salt, the solvent was removed by rotary evaporator, and the crude product was purified by flash chromatography on silica gel with ethyl acetate/hexane (1:1) to yield the desired product as a white solid (yield, 82%). 1H NMR (500 MHz, CDCl3, δ (ppm)): 6.50 (m, 2H), 6.11 (m, 1H), 5.60 (m, 1H), 5.25 (s, 2H), 4.29 (m, 2H), 3.69 (m, 2H), 2.94 (dd, 2H), 1.90 (s, 3H). 13C NMR (125 MHz, CDCl3, δ (ppm)): 175.9, 166.5, 136.0, 135.1, 125.9, 81.2, 60.3, 47.1, 38.0, 17.8. Preparation of Poly(Fur-MalMA-co-HMA): Radical Copolymerization of Fur-MalMA and HMA. HMA (1.8 mL, 9.0 mmol) and FurMalMA (0.27 g, 1.0 mmol) were dissolved in 15 mL of dioxane. Then, AIBN (0.015 g, 0.090 mmol) was added into the solution followed by purging with argon for 20 min. Subsequently, the reaction vessel was placed in a preheated oil bath at 60 °C for 8 h. The polymer was isolated by exposing the polymerization solution to air and then precipitating in cold methanol. The precipitate was dried under vacuum at room temperature for 24 h. Deprotection of Poly(Fur-MalMA-co-HMA). Poly(Fur-MalMA-coHMA) (1.0 g) was dissolved in dry toluene (15 mL). The reaction vessel was then placed in a preheated oil bath at 110 °C for 24 h. Upon the deprotection, the solution was concentrated to 5 mL, and the polymer was isolated by precipitating into a large amount of cold methanol. The precipitate was dried under vacuum at room temperature for 24 h. Synthesis of Furan-Modified Fluorescein. The synthesis of furanfluorescein was slightly modified according to our previous literature.61 In a typical procedure, fluorescein (1.0 g, 3.0 mmol) was dissolved in a mixed solvent of DCM (9 mL) and DMF (6 mL). Thereafter, EDC·HCl (1.3 g, 6.6 mmol) and DMAP (0.09 g, 0.84 mmol) were added into the solution. After being stirred for 3 h, furfurylamine (0.66 mL, 7.5 mmol) was added. The reaction was left to stir for 24 h. Upon the reaction, DMF and DCM were removed under vacuum. The residue was redissolved in ethyl acetate (40 mL) followed by sequential washing with HCl aqueous solution (1M) and brine. The organic layer was collected and dried over sodium sulfate. The product was collected by removing the solvent under vacuum (yield, 43%). 1H NMR (500 MHz, CD2Cl2 and MeOD, δ (ppm)): 7.91 (m, 1H), 7.52 (m, 2H), 7.02 (m, 2H), 6.60 (m, 2H), 6.35 (m, 4H), 6.02 (S, 1H), 5.76 (S, 1H), 4.31 (S, 2H). Preparation of a Polymer-Fluorescein Conjugate via Diels− Alder Reaction. In a typical procedure, poly(MalMA-co-HMA) (660 mg, 0.26 mmol of Mal groups) was mixed with furan-fluorescein (214 mg, 0.52 mmol) in dry DMSO (4 mL). The vial was placed in a 50 °C oil bath. Samples were periodically taken and analyzed by 1H NMR until the reaction was completed. The polymer was then isolated by

precipitating into a large amount of cold methanol. The precipitate was dried under vacuum at room temperature for 24 h. Iron Oxide Nanoparticle Synthesis. A stoichiometrically defined iron oleate was prepared according to the method developed by Vreeland et al.57 to be used in the Extended LaMer mechanism-based synthesis. Briefly, 20.0 g (56.7 mmol) of iron acetylacetonate [Fe(acac)3), > 98% pure, TCI America] and 80.0 g (284 mmol) of oleic acid (90% technical grade, Sigma-Aldrich) were charged into a 500 mL three-neck reactor flask. The reaction under 100 sccm of argon was thoroughly mixed using a Caframo compact overhead stirrer at 350 rpm. The reaction mixture was heated to about 320 °C at a ramp rate of 8 °C/min using a fabric heating mantle, and the temperature was controlled using a Digi-sense temperature controller. After 35 min at 320 °C, a dark brown waxy solid was obtained and used as the precursor for the Extended LaMer mechanism-based synthesis after 24 h. For the magnetic nanoparticle synthesis, first, 14.0 g (48.3 mmol) of docosane (90% pure, Sigma-Aldrich) was initially heated to 350 °C for 50−60 min at a ramp rate of 7−8 °C/min in a 100 mL three-neck reaction flask. The rate of addition of argon, the inert gas, was controlled at 100 sccm using mass flow controllers from Alicat Scientific. Once the reactor reached 350 °C, the controlled addition (using a syringe pump) of 30 mL of iron oleate precursor (0.63 M Fe) mixed with 55 mL of 1-octadecene (90% technical grade, SigmaAldrich) was initiated along with an oxygen feed of 20% oxygen and 80% argon (Airgas) at a rate of 9.47 sccm, controlled using a mass flow controller (Bronkhorst USA). Uniform mixing at 350 rpm was ensured, and the reaction temperature was controlled at 350 °C for 6 h using a Digisense temperature controller. The reaction mixture was allowed to cool to room temperature, and iron oxide nanoparticles obtained at the end of the reaction were purified by suspending the product in chloroform in an 1:1 volume ratio. Further, the particles were precipitated using 1:3 of chloroform/product mixture to acetone and decanted using magnets. The particles were further suspended in THF at required concentrations for preparation of MCNCs. Iron Oxide Nanoparticle Characterization. DLS measurements were performed using a Brookhaven Instruments 90Plus/BI-MAS, operated at room temperature. Equilibrium magnetization curves were found using a Quantum Design magnetic property measurement system 3 (MPMS 3) superconducting quantum inference device (SQUID) magnetometer. Magnetization curves at room temperature were obtained for liquid samples. Physical diameter was obtained with a Hitachi H 7000 TEM operating at an accelerating voltage of 100 keV. FNP for Assembly of MCNCs. A typical procedure for preparing MCNCs by FNP was as follows. First, 22.1 mg of PEG−PLA was weighed in a dry, empty 20 mL glass vial. The poly(MalMA-co-HMA) fluorescein pendent polymer at 13 mg/mL in THF (1 mL) was added to the vial. Then, 1.476 mL of SPIONs in THF at 6.3 mg/mL core (iron oxide) content was added to the vial, followed by 0.624 mL of THF, for a total amount of 3.1 mL of THF with 4.2 mg/mL poly(MalMA-co-HMA) fluorescein pendent polymer, 3.0 mg/mL SPION, and 7.2 mg/mL PEG−PLA. This solution was sonicated in a bath sonicator for 20 min. FNP was then performed by loading 3 mL of the THF solution into a 10 mL syringe and 3 mL of water into a separate 10 mL syringe, and both solutions were pushed through a confined impinging jet (CIJ) mixer at approximately 60 mL/min. The resulting product was recovered in a 20 mL vial with 12 mL of deionized water present in the vial before catching the FNP product. This product solution was then filtered magnetically by passing it through a 5 mL syringe packed with 0.04−0.05 mm in diameter steel wool, which was placed between two permanent magnets. The steel wool was packed to 2 mL with a packing density of 0.91 g/mL, and a polyether ether ketone stopcock was used to control the drip rate. The entire 18 mL of product solution was passed through the syringe between the magnets at a drip rate of approximately 2 s per drop. To remove excess starting material, 8 mL of PBS was added to the syringe and collected in 3 filtrates. A retentate of MCNC product was collected by removing the syringe from between the magnets and passing 6 mL of PBS through it. This retentate was stored at 4 °C G

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ACS Applied Polymer Materials when not in use for release testing. The iron oxide concentration of this retentate, determined by a previously reported o-phenanthroline spectrophotometric assay,58 was 0.8 mg/mL, indicating that 51.6% of the iron oxide that entered the FNP process was recovered. Fluorophore Release Tests. First, 1.67 mL of MeOH was added to 5 mL of the MCNC retentate in PBS to make a solution of MCNCs in 25/75 v/v MeOH/PBS release media. Then, 25% methanol was used as the release media to increase the solubility of the released furan-fluorescein and ensure sink conditions. For all release tests, the general procedure was to add 100 μL of MCNCs at 0.6 mg/mL iron oxide concentration to a small glass vial and then to expose the sample to external heating via a water bath or an alternating magnetic field to actuate release of the fluorophore. The alternating magnetic field frequency was 346 kHz with an amplitude of 59.6 kA/m. The amplitude of 59.6 kA/m was found by using a 2D magnetic field probe from AMF Life Systems and was within 5% of the theoretical value. This amplitude was decreased in most cases to keep the samples from going above the target temperature by decreasing the alternating current flowing through the coil. Released fluorophore was separated from the MCNC solution by adding 250 μL of 25/75 v/v MeOH/ PBS media to the 100 μL of sample, and then the solution was transferred to an Amicon Ultra 0.5 mL 100 kDa centrifugal filter and centrifuged at 20 000 rcf for 2.5 min. Then, 200 μL of filtrate was removed and spotted onto a black polystyrene 96-well plate, and fluorescence was measured at 494 nm excitation and 525 nm emission with a 515 nm cutoff in a SpectraMax M5 spectrometer. Calibration curves of fluorescein in the 27/75 v/v MeOH/PBS were linear when plotted as the concentration vs RFU for the applicable concentration ranges of the release tests (not shown). Figure 9 shows the “change in RFU” and was found by taking the RFU values post-AMF and subtracting from them the RFU values of samples before AMF. Figures S12 and S8 show the percentage release and relative release. These percentages were found by dividing a given sample’s change in RFU by the change in RFU from “100% release samples”. then, 100% release samples were subjected to 90 °C for 10 h, at which point no further release occurred. For the spatially controlled release test, a custom coil 12 cm in diameter and having three turns, was used to generate an AMF at 303 kHz and an amplitude of 10.8 kA/m. To produce a field gradient of 1.27 T/m, two opposing NdFeB permanent magnets (3” × 3” × 1”) were housed in a custom built water cooled magnet rig and placed on the outside of the coil, with the north ends of the magnet facing each other as illustrated in Figure 9b. Three insulated sample vials in “position A” were put on the inner side of the coil as shown in Figure 9a,b and another 3 insulated sample vials were in “position B”, which was in the center of the coil, approximately 5.5 cm away from “position A”. AMF was applied for 1 h, and the temperature in the vials and the coil was monitored using fiber optic temperature probes. The AMF coil temperature was maintained slightly above the sample temperature located in position A for experiments with field gradient to prevent cooling of the sample due to the coil. Biocompatibility of MCNCs. MDA-MB-231 cells were seeded at 3500 cells per well in 100 μL of DMEM in a 96-well plate made of 1 × 8 well strips. Two days post seeding, 2 mg/mL of control MCNCs (no cargo loaded) in media were added to 8 wells, and a negative control of media only and positive control of Triton X-100 solution were included. After 5.5 h, MCNCs were removed, and fresh media was added. Cells were then grown for 2 additional days, and a viability test was done by adding 15 μL of CellTiter-Blue to the media and measuring fluorescence at 560/590 nm.

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conjugate polymer in toluene, an additional release test with the MCNCs, magnetic and biocompatibility tests of the MCNCs, MCNC stability as seen via DLS over 1 month, and the temperature profiles for the AMF release testing (PDF)

AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]fl.edu. *E-mail: carlos.rinaldi@ufl.edu. ORCID

Hao Sun: 0000-0001-9153-4021 Brent S. Sumerlin: 0000-0001-5749-5444 Carlos Rinaldi: 0000-0001-8886-5612 Present Address ⊥

Department of Chemistry, Northwestern University, Evanston, Illinois 60201, USA. Author Contributions ∥

E.G.F. and H.S. contributed equally to this work.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported in part by NSF HRD-1345156, NSF (DMR-1606410), and NIH 1R21EB018453-01A1. We thank the University of Florida for providing support and research infrastructure to enable us to accomplish this work.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsapm.8b00100. NMR and GPC evidence of the polymer creation, a scheme describing the synthesis of furan-modified fluorescein, a fluorophore release test with just the H

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