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Thermoresponsive magnetic composite nanomaterials for multimodal cancer therapy
Acta Biomaterialia 6 (2010) 502–510
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Thermoresponsive magnetic composite nanomaterials for multimodal cancer therapy S. Purushotham, R.V. Ramanujan * School of Materials Science and Engineering, Nanyang Technological University, Block N4.1, Nanyang Avenue, Singapore 639798, Singapore
a r t i c l e
i n f o
Article history: Received 25 March 2009 Received in revised form 17 June 2009 Accepted 7 July 2009 Available online 23 July 2009 Keywords: Magnetic nanoparticles Thermoresponsive materials Drug delivery Hyperthermia Cancer therapy
a b s t r a c t The synthesis, characterization and property evaluation of drug-loaded polymer-coated magnetic nanoparticles (MNPs) relevant to multimodal cancer therapy has been studied. The hyperthermia and controlled drug release characteristics of these particles was examined. Magnetite (Fe3O4)–poly-n-(isopropylacrylamide) (PNIPAM) composite MNPs were synthesized in a core–shell morphology by dispersion polymerization of n-(isopropylacrylamide) chains in the presence of a magnetite ferrofluid. These core–shell composite particles, with a core diameter of 13 nm, were loaded with the anti-cancer drug doxorubicin (dox), and the resulting composite nanoparticles (CNPs) exhibit thermoresponsive properties. The magnetic properties of the composite particles are close to those of the uncoated magnetic particles. In an alternating magnetic field (AMF), composite particles loaded with 4.15 wt.% dox exhibit excellent heating properties as well as simultaneous drug release. Drug release testing confirmed that release was much higher above the lower critical solution temperature (LCST) of the CNP, with a release of up to 78.1% of bound dox in 29 h. Controlled drug release testing of the particles reveals that the thermoresponsive property can act as an on/off switch by blocking drug release below the LCST. Our work suggests that these dox-loaded polymer-coated MNPs show excellent in vitro hyperthermia and drug release behavior, with the ability to release drugs in the presence of AMF, and the potential to act as agents for combined targeting, hyperthermia and controlled drug release treatment of cancer. Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction There has been enormous interest in the exciting area of magnetic nanoparticle (MNP)-based systems for cancer treatment [1–26] as conventional cancer therapies such as radiotherapy, hyperthermia and chemotherapy have serious drawbacks. The limitations of chemotherapy techniques include adverse effects on healthy tissue due to indiscriminate distribution of cytotoxic drugs in the body, insufficient local drug concentration in the tumor and poor control over drug release. In hyperthermia therapy, the tumor is heated in the temperature range of 41–47 °C, causing death of the cancerous cells, but sparing healthy cells [20,27,28]. In conventional hyperthermia techniques—e.g. radiofrequency-based hyperthermia, isolated hepatic perfusion, water baths and heating rods—temperature control is poor, heat distribution is not optimum and there is risk of organ damage due to overheating [20]. Drug-loaded composite MNPs offer a method of simultaneously achieving drug targeting, controlled drug release and hyperthermia treatment for cancer, thus overcoming several limitations of conventional cancer therapy.
* Corresponding author. Tel.: +65 67904342; fax: +65 67909081. E-mail address: [email protected] (R.V. Ramanujan).
The synergistic therapeutic effects of simultaneous chemotherapy and hyperthermia can exceed the individual or sequential application of these techniques [28–31]. Therefore, drug-loaded composite MNPs with a magnetic core and a polymer shell capable of acting as multifunctional agents for combined drug targeting, controlled release and hyperthermia therapy are highly desirable. Such particles can be injected into the appropriate blood vessels and targeted to the tumor by means of a suitable external magnetic field gradient [1,9,24,32,33]. Chemotherapy occurs by drug release from the composite particles which have been targeted to the tumor region. An external alternating magnetic field (AMF) can be applied to generate heat in the targeted particles [13,14,20,34], and the resulting temperature rise can be used for hyperthermia treatment of cancer. If the polymer shell of the nanoparticle is responsive to stimuli such as pH and temperature, drug release can also be controlled. This paper reports a study of novel iron oxide–poly-n-isopropylacrylamide (PNIPAM) composite nanoparticles (CNPs) suitable for magnetic targeting followed by simultaneous magnetic hyperthermia and chemotherapeutic drug release. The core of the CNP consists of iron oxide; PNIPAM is chosen as the thermosensitive polymer shell. PNIPAM is an inverse thermosensitive polymer that has been studied for several biomedical applications [35–41]; it undergoes reversible volume change in water at a lower critical solution temperature (LCST) in the range 30–35 °C. The proximity
1742-7061/$ - see front matter Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.07.004
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of the LCST to the human body temperature (37 °C) and the ability to tune the LCST by addition of appropriate copolymers makes PNIPAM hydrogels useful for controlled drug delivery. Below the LCST, PNIPAM chains are soluble in water and the polymer is in a swollen state, inhibiting the transport of drug molecules through the matrix, resulting in a slow rate of release. Heat is generated in the iron oxide core upon exposure to an AMF, and this heat is conducted to the polymer shell, increasing the temperature of the PNIPAM. At and above the LCST, the shell collapses, squeezing out drug molecules, resulting in rapid drug release (Fig. 1). Both Fe3O4 (magnetite) and c-Fe2O3 (maghemite) are known to be biocompatible and non-toxic in concentrations relevant to clinical applications. In vivo experiments on animal models have shown that iron oxides are suitable for drug delivery and hyperthermia [1,17,20,22,42,43]. Human clinical trials for drug delivery [33] and hyperthermia [14,15,25] have been conducted with iron oxide-based ferrofluids, the injected dosage being well tolerated by patients. Superparamagnetic iron oxides (FDA approved) are also used as contrast-enhancing agents for clinical magnetic resonance imaging (MRI) [44–46]. In vitro drug release during AMF-induced heating has been previously studied using drug-loaded thermosensitive carriers [4,6,47] and dye-loaded PNIPAM–magnet systems [3,48]. Drug release from dox-loaded magnetic particles encapsulated by a thermosensitive polymer without an applied magnetic field has recently been studied [26]. The novelty of this work is in the synthesis, characterization and property evaluation of thermosensitive polymer-encapsulated magnetic particles relevant to combined magnetic targeting, hyperthermia and drug release applications. Previously, we have shown that drug-loaded CNPs can be targeted to hepatocellular carcinoma (HCC) in a buffalo rat model, and localization in the tumor was confirmed by MRI and histology [9]. 2. Materials and methods 2.1. Synthesis of MNPs Fe3O4 MNPs were synthesized by a reverse co-precipitation method adapted from Aono et al. [49] to obtain particles with an
Swollen PNIPAM shell Slow drug release
Drug loaded PNIPAM shell
T > LCST
T < LCST
T < LCST
Hyperthermia T > LCST MNP cluster
PNIPAM shell shrinks Faster drug release Fig. 1. Schematic diagram showing drug release from PNIPAM-coated MNPs below and above the polymer LCST.
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average size of 10–14 nm. In a typical experiment, FeCl36H2O (0.1 M, 27 g) (Merck) and FeCl24H2O (0.05 M, 10 g) (Merck) were dissolved in 200 ml of ultrapure MilliQ water and heated to 80 °C. The synthesis was performed inside a glovebox in an inert atmosphere, and nitrogen was bubbled through all solutions for 30 min before use. Under vigorous stirring, 140 ml of NaOH solution (0.14 M, 5.6 g in 140 ml) (Merck) was dripped into this solution at a rate of 6 ml min1, upon which the solution turned black, indicating the precipitation of Fe3O4 (magnetite). The mixture was maintained at 80 °C for 30 min, cooled to room temperature and centrifuged at 8000 rpm for 15 min. The supernatant was discarded and the particles were extracted. After washing several times with MilliQ water and ethanol to remove residual reactants, the particles were dried in vacuum for 24 h. 2.2. Size selection by magnetic separation The dried MNPs were ultrasonically redispersed in 500 ml of ethanol until a dark suspension was obtained. A 0.8 T block permanent magnet was placed on the outer wall of the beaker to attract the most magnetically responsive particles in the suspension. The supernatant was extracted with a peristaltic pump (Simon, Manostat). The responsive particles were redispersed in ethanol and the process was repeated three times. The magnetically separated particles were dried in vacuum and stored in a vacuum oven at room temperature. 2.3. Preparation of iron oxide–polymer CNPs CNPs were prepared, in the presence of MNPs, by free radical dispersion polymerization of NIPAM monomer in water with N,Nmethylene(bis)acrylamide (BAAm) as the cross-linker, N,N,N0 ,N0 -tetramethylene diamine (TEMED) as the accelerator and ammonium persulfate (APS) as the initiator [3,50]. Our procedure does not involve surfactants which can alter the LCST [40,51]. MilliQ water was used, through which nitrogen was bubbled for 15 min prior to use. The NIPAM monomer was purified by recrystallization from hexane and dried in vacuum for 24 h before use. In a typical experiment 35 ml of ferrofluid containing 850 mg MNP was added to a solution of 960 mg NIPAM (Aldrich, 97%) in 30 ml water and ultrasonicated for 10 min. The mixture was transferred to a beaker and stirred at 24,000 rpm by an IKA-Werke T25 Ultra Turrax high-speed dispersion unit while 2.5 ml of BAAm (Fluka, 98%) solution (8 mg in 5 ml) and 0.5 ml of APS (J.T. Baker, 98.5%) solution (180 mg in 5 ml) were added to the mixture. Finally 1 ml of TEMED (Lancaster, 99%) was added dropwise with continuous stirring. After 20 min of stirring, the mixture was left undisturbed for 2 h, after which the CNPs were magnetically separated and washed several times with MilliQ water. Some of the CNPs were dried in vacuum for 24 h, resulting in dehydration of the particles: the remainder were used as-synthesized in the swollen state for dynamic light scattering (DLS) and microcalorimetry experiments. 2.4. Drug loading The water-soluble anti-cancer drug doxorubicin (dox) was chosen as a model drug. Typically, 26.5 mg (loading target 10 wt.% of CNP) of doxorubicin hydrochloride was dissolved in 20 ml of MilliQ by ultrasonication for 10 min. Two hundred and forty-five milligrams of dehydrated CNPs was added to the dox solution and ultrasonicated for an additional 10 min. The bottle was sealed and gently shaken in a rotary shaker in the dark for 20 h at 24 °C to facilitate both reswelling of the PNIPAM shell in the aqueous dox solution as well as to facilitate dox uptake. After loading, the dox-loaded CNPs (Dox-CNPs) were magnetically separated. To determine residual dox content, the supernatant was analyzed in
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a UV–visible spectrophotometer (Shimadzu UV1700) and compared with reference plots of dox in MilliQ water at a wavelength of 482 nm. The drug loading of the Dox-CNPs was determined as the difference between the dox content of the solution before loading and the dox content of the supernatant. To investigate the effect of dehydration on the particle performance, part of the batch of prepared particles was dried in vacuum for 48 h and the remainder was used in the swollen condition for experiments. 2.5. Phase identification and crystallite size X-ray diffraction scans were performed on the dried powder using a Shimadzu 6000 X-ray diffractomer with CuKa radiation of wavelength 1.54056 Å, in the 2h range 10–90° at a scan rate of 1° min1. Phase identification was performed by matching peak positions and relative intensities to reference JCPDS files. The crystallite size was calculated using the Scherrer formula. 2.6. Particle size and hydrodynamic diameter The particle size of the bare MNPs was measured by transmission electron microscopy. A drop of MNP ferrofluid was placed on a carbon-coated copper TEM grid, dried in vacuum and imaging was carried out in a JEOL 2010F microscope operating at 200 kV. Particle size was measured from transmission electron micrographs using ImageJ image processing software. The size of the CNP particles could not be reliably examined by TEM since the electron beam caused ablation of the polymer shell. DLS was used to measure the hydrodynamic diameters of aggregates of MNP and swollen CNP samples in MilliQ water. Dilute suspensions (<0.2 mg ml1) were transferred to low-volume quartz cuvettes and investigated using a Malvern Instruments Zetasizer Nano DLS unit. The samples were equilibrated at 25 °C for 5 min before each measurement. 2.7. Thermogravimetric analysis (TGA) Twenty milligrams of dehydrated CNP powder was placed in a ceramic pan, heated to 600 °C at a rate of 20 °C min1 under nitrogen flow in a TA Instruments TGA Q500 thermogravimetric analyzer, and the weight loss was measured as a function of temperature. To determine the polymer content of the CNPs, the weight loss plot of bare MNPs measured under identical conditions was used as reference.
2.9. Magnetic properties Magnetic properties of bare and dox-loaded composite MNP powders were evaluated using a LakeShore 7404 vibrating sample magnetometer, in an applied field range of 0–10 kOe. Magnetic particle diameter was calculated by fitting the magnetization curve to the Langevin function [53]:
M lH kB T ¼ coth ; Ms kB T lH
ð1Þ
where M is the magnetization for a field strength H, Ms is the saturation magnetization, kB is the Boltzmann constant, the true magnetic moment of each particle l = (MspD3/6) and D is the magnetic particle diameter. 2.10. In vitro heating behavior in an AMF MNP, swollen CNP and swollen Dox-CNP powders were ultrasonically dispersed in PBS to prepare ferrofluids with a particle concentration of 2 mg ml1. Three milliliters of ferrofluids contained in glass bottles were placed inside a water-cooled five-loop copper induction coil energized by an AC generator (Inductelec, UK) with operating frequency of 375 kHz and field strength of 4 kA m1. Ceramic wool thermal insulation was used around the bottles to minimize heat loss, and temperature data was continuously logged using a Luxtron MD600 fiber optic thermometry unit connected to a laptop computer; at least three samples were tested for each type of ferrofluid. The heating power, or specific absorption rate (SAR), of the particles was calculated from the equation:
SAR ¼ C
DT massferrofluid ; Dt massnanoparticles
ð2Þ
where C is the mass-weighted heat capacity of the ferrofluid, and ðDT=DtÞ is the slope of the initial section of the temperature vs. time curve [34]. In order to study drug release during hyperthermia, additional experiments were carried out using dehydrated Dox-CNPs. The AMF strength was manually varied to raise the temperature in less than 5 min and maintain the temperature in the range between 41 and 48 °C for various time intervals up to 60 min. Samples of 5 ml of ferrofluid with a Dox-CNP concentration of 1 mg ml1 were used. At the end of each heating experiment, the particles were magnetically separated from the suspension, and drug release was quantified by measuring the dox content in the supernatant via UV–visible spectrophotometry.
2.8. Measurement of LCST
2.11. Temperature-dependent drug release without AMF
The determination of LCST for PNIPAM aqueous solutions is often performed by optical light transmission methods (e.g. UV–visible spectrophotometry) or microcalorimetry [51,52]. However, for composite MNPs optical methods do not yield reliable results at concentrations of interest (>1 mg ml1), and so LCST measurements were performed with a TA Instruments DSC 2920 modulated differential scanning calorimeter (MDSC). Swollen CNPs as well as Dox-CNPs, and dehydrated Dox-CNPs dispersed in PBS, were studied. The dehydrated particles were dispersed in PBS and refrigerated at 8 °C for 4 h before testing to enable reswelling of the polymer shell. Reference pans contained ferrofluid of MNPs without any polymer. A few microliters of the sample suspensions were placed in hermetically sealed aluminum pans and heated from 20 to 50 °C at a scanning rate of 0.5 °C min1 with modulation amplitude of 0.2 °C min1. The pans were equilibrated in the DSC cell at 20 °C for 20 min before a heating run. The MDSC thermograms were analyzed by TA Universal Analysis software.
Drug release for extended time periods without an AMF was studied at temperatures of 24 °C (below LCST), 37 °C (body temperature, near LCST) and 43 °C (above LCST). Samples of 5 ml of ferrofluids with concentration of 1 mg ml1 were prepared by ultrasonically dispersing Dox-CNP powders in PBS (pH 7.4) in glass test-tubes. A minimum of three samples were prepared for testing at each temperature. The sealed test-tubes were placed in preheated water baths (Fisher Isotemp 202) and the temperature stabilized (±1 °C) at 24, 37 and 43 °C for a total of 29 h. Periodically, 1 ml of the release medium, free of nanoparticles, was extracted and replaced with fresh PBS. Dox content of the extracted solution was measured by UV–visible spectrophotometry and the cumulative dox release calculated. 2.12. Controlled drug release without AMF The ability of swollen Dox-CNPs to modulate drug release by varying temperature was studied. The experimental details are as
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described in Section 2.11; the temperature was cycled between 24, 37 and 43 °C for various time periods and dox release measured via UV–visible spectrophotometry. 2.13. Modeling of power output from MNPs in an AMF The model proposed by Rosensweig [34] was used to examine the effect of polymer shell thickness on the heating power of MNPs exposed to an AMF. The following values were used for the calculations: ferrofluid concentration = 2 mg ml1 of monodisperse spherical magnetite nanoparticles; magnetite domain magnetization Md = 446 kA m1 [34]; anisotropy constant K = 32 kJ m3 [34]; temperature T = 310 K (37 °C); g = 6.92 104 N s m2 (at 37 °C); frequency f = 375 kHz; field strength H0 = 4 kA m1. The hydrodynamic diameter is taken to be the magnetic particle diameter for MNPs and the overall particle diameter for CNPs. The analysis does not incorporate polydispersity, viscous and hysteresis losses, interparticle interactions and changes in viscosity with increasing temperature. 3. Results and discussion 3.1. Particle size and structure Fig. 2 shows a transmission electron micrograph of the bare MNP. The average size of bare nanoparticles after magnetic size selection was calculated to be 11.9 nm from the XRD data, 10.1 nm from the magnetization curve, and 12.9 nm measured from transmission electron micrographs (standard deviation = 3). Fig. 3 shows the XRD pattern, which indicates that the particles are crystalline; peak matching with reference JCPDS files shows that the particles are primarily Fe3O4 (magnetite). The average hydrodynamic diameter (DH) and polydispersity index (PDI) of MNPs and CNPs in water at 25 °C were calculated from an exponential fit to the intensity correlation function of the light scattering data obtained from DLS experiments. The bare MNPs had an average DH of 203 nm with a PDI of 0.301, the corresponding figures for the CNPs were 158 nm and 0.161, respectively. From theoretical calculations of heat generation in an AMF, the optimum average size (DM) of the MNPs is 12 nm with a size dis-
Fig. 3. X-ray diffraction pattern of bare Fe3O4 nanoparticles. Numbers in brackets indicate the planes of diffraction.
tribution in the range of 12–13 nm. With this particle size and size distribution, an increase in hydrodynamic diameter DH due to PNIPAM incorporation would not significantly affect heat generation. The influence of the particle hydrodynamic diameter (DH) on SAR for selected values of DM is shown in Fig. 4. For particles with a core diameter less than 12 nm there is a small increase in SAR as DH is increased. For 12 nm DM particles, SAR increases gradually with increasing DH. When DM is larger than 12 nm, the increase in DH is accompanied by a sharp decrease in SAR. Néel losses will be maximum for the 12–13 nm core size. At this size, although Brownian relaxation losses are low, there is no sharp reduction in SAR if DH increases either due to clustering or excessive polymer coating during synthesis. For discrete, bare MNPs, DH is usually larger than DM, so the same conditions apply. The average MNP size synthesized in our experiments, 13 nm, is in the size range predicted from our calculations to yield high SAR. DLS shows a larger hydrodynamic diameter and wider distribution for agglomerates of bare MNPs (203 nm) compared to those of the CNPs (158 nm) due to the following: bare iron oxide nanoparticles, without an appropriate surfactant or stabilizing polymer layer, have a hydrophobic surface which induces clustering in aqueous suspensions as they try to reduce their surface area in contact with water [11]. Additionally, due to the broader size distribution of these particles, strong attractive interparticle magnetic forces will encourage clustering. On the other hand, below the LCST of PNIPAM, the CNPs and clusters are partially stabilized by the hydrophilic polymer shell, and the shell also reduces the attraction between particle clusters. Since DLS depends on scattered light intensity from aggregates, it is highly sensitive to larger particle sizes (d6 dependence of Rayleigh scattering), and scattering from relatively few but larger particle clusters tends to skew the reported average size; this can explain why the CNP samples show a narrower size distribution (polydispersity index of 0.161 compared to 0.301 for MNP). 3.2. PNIPAM and dox content of CNP
Fig. 2. Bright-field transmission electron micrograph of Fe3O4 nanoparticles produced by the reverse co-precipitation method. Aggregation is due to the attractive magnetic forces between the particles and the absence of a surfactant layer.
The TGA weight loss vs. temperature plot for both MNPs and CNPs is shown in Fig. 5. Heating in the range of room temperature to 130 °C results in expulsion of water from the polymer, and PNIPAM decomposition occurs between 200 and 440 °C [54]. From the TGA data, the PNIPAM content of CNPs is estimated to be 1.3 wt.%. The MNP sample shows less than 0.5% weight loss, confirming that the weight change for the CNP sample beyond 200 °C is due to PNIPAM degradation. The UV–visible spectrophotometry results (not shown) revealed that the drug loading of the Dox-CNPs was 4.16 wt.%. Polymer encapsulation of MNPs is essential for loading drugs or other therapeutic agents onto MNPs; however, the magnetic prop-
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Fig. 4. Influence of hydrodynamic diameter DH on the SAR of Fe3O4 nanoparticles in ferrofluids for selected magnetic diameter DM values, at f = 375 kHz and H0 = 4 kA m1.
Fig. 5. TGA weight loss curves for bare MNPs (top) and PNIPAM-encapsulated MNPs (bottom).
erties of the composite particles can be inferior to those of the MNPs, hence the weight fraction of non-magnetic components must be kept to a minimum. The PNIPAM content of 1.3 wt.% in the CNPs was found to be adequate to achieve loading (4.16 wt.% Dox-CNP) and delivery of therapeutic quantities of dox with little adverse effect on magnetic properties. In dispersion polymerization, both heterogeneous and homogeneous nucleation occurs [50] with nucleation and growth of the polymeric shell around the iron oxide nanoparticle core. Bonding between the iron oxide ions and surface hydroxy groups to the free ion pair of amidic nitrogen in NIPAM aids adsorption of the monomer to the surfaces of the nanoparticles. Nucleation and growth of PNIPAM around magnetic clusters occurred rapidly and uniformly due to the combination of a high weight ratio of iron oxide to NIPAM (85:96), high stirring speed (24,000 rpm) and the presence of accelerator TEMED yielding composite particles in the nanoscale range with relatively low PDI. The volume fraction of PNIPAM in CNPs is much higher than 1.3% due to its lower density compared to Fe3O4. When drug loading is expressed in terms of PNIPAM weight, the 4.16 wt.% dox content of the CNPs is equivalent to 320 wt.% of PNIPAM; this value is
due to the large surface area to volume ratio of the polymer shell in the nanoscale range which enhances the adsorption of dox. The modulated DSC thermograms for swollen CNPs and DoxCNPs are shown in Fig. 6a and b, respectively; those for the dehydrated Dox-CNPs in Fig. 7. The LCST was measured as the onset temperature as this temperature, unlike the peak maximum temperature, is relatively insensitive to sample weight related artifacts [52]. The swollen particles exhibit two distinct endothermic peaks (Fig. 6a and b) at the LCST. For the swollen CNPs (Fig. 6a), the onset of the first endothermic event is at 35.3 °C, and the endothermic peak is at 35.8 °C. The second transition has an onset temperature of 39.4 °C and an endothermic peak at 39.9 °C. For the swollen Dox-CNPs, the onset temperatures of the events are 35 and 42 °C, with peaks at 35.5 and 43.5 °C, respectively. The thermogram of the dehydrated particles (Fig. 7) lacks the distinct peaks exhibited by those of the swollen particles, indicating the absence of a welldefined transition temperature. For the PNIPAM homopolymer, the LCST in aqueous solution is strongly influenced by additives that can bind to the polymer molecules or affect the structure of water around them [40,51]. Fig. 6a shows an LCST of 35.3 °C for the swollen CNPs in PBS, which indicates that the MNPs in PNIPAM does not significantly alter the LCST despite the hydrophobic surface of the MNP and the tendency of the ions in the buffer to interact with the PNIPAM chains. Another transition at 39.4 °C may be due to local changes in the structure of PNIPAM during dispersion polymerization. The swollen Dox-CNPs also show two endothermic events, the first of which has an onset temperature of 35 °C, close to the LCST of the CNP. Interestingly, the second and stronger transition has an onset temperature of 42 °C. The shift in the temperature for the second transition was due to the loading of the hydrophilic dox in the PNIPAM matrix, resulting in hydrogen bonding between dox and the amide group of PNIPAM. In contrast, the dehydrated Dox-CNP sample (Fig. 7) does not exhibit any distinct transition temperature, and vacuum dehydration causes a significant change in the PNIPAM structure, leading to the absence of a LCST. 3.3. Magnetic properties The MNP and Dox-CNP magnetization curves at room temperature suggest similar magnetic properties (Fig. 8). The bare MNPs
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Fig. 6. Modulated DSC thermograms for swollen (a) CNPs dispersed in PBS (b) Dox-CNPs dispersed in PBS.
3.4. In vitro heating behavior in an AMF
Fig. 7. Modulated DSC thermogram for dehydrated Dox-CNPs dispersed in PBS.
exhibit a saturation magnetization (MS) of 61 emu g1; this decreases to 56.8 emu g1 for the Dox-CNPs. The coercivity (Hc) of Dox-CNPs is 8.5 Oe, lower than the Hc of 43 Oe exhibited by the bare MNPs. The MS values for MNPs and Dox-CNPs are lower than those of bulk magnetite [55]. Assuming an anisotropy value of 32 kJ m3 [34], at 27 °C, Fe3O4 nanoparticles smaller than 15 nm are expected to be superparamagnetic and should display a sigmoidal anhysteric magnetization curve. But due to size distribution and clustering effects the experimental magnetization curves display a small amount of hysteresis (Fig. 8). The decrease in coercivity from 43 to 8.5 Oe for the Dox-CNP samples is due to the role of the polymer in reducing aggregation of MNPs (Section 3.2).
Fig. 8. Room temperature magnetization curves for bare MNPs and Dox-CNPs for field strengths up to 10 kOe.
The temperature rise from the ambient temperature plotted against time for MNP, CNP and Dox-CNP ferrofluids exposed to a 375 kHz, 4 kA m1 AMF is shown in Fig. 9a. The bare MNPs reaches the hyperthermia temperature (41 °C) in 335 s, while the CNPs and Dox-CNPs attain this temperature more rapidly, in 295 and 255 s, respectively. From the ferrofluid concentration and slope of the heating curve up to 120 s, the SAR was calculated to be 41.2 W g1 for MNP, 72 W g1 for CNP and 74.1 W g1 for DoxCNP samples. A typical in vitro simultaneous hyperthermia and drug release curve, in this case for 30 min of hyperthermia exposure, is shown in Fig. 9b. Compared to the heating behavior with 2 mg ml1 ferrofluids, the lower (1 mg ml1) concentration of Dox-CNP ferrofluid reaches 41 °C in 240 s due to the higher applied field strength and ambient temperature. The temperature could be readily stabilized between 41 and 47 °C for clinically relevant time periods of 60 min [14,15] by manually varying the field strength. After exposure to hyperthermia temperatures (41–47 °C) for 10 min, 26.2% of the bound dox was released from the Dox-CNPs, increasing to 32.9% for 30 min and 36.1% for 60 min of exposure. This clearly demonstrates that hyperthermia temperatures and simultaneous drug release in therapeutic quantities can be achieved in vitro by the Dox-CNP ferrofluids. The predicted SAR values for 12 and 13 nm MNPs for the conditions stated in Section 2.13, and assuming DH = DM, are 28.4 W g1 and 50.3 W g1, respectively. When the hydrodynamic diameter measured by DLS at 25 °C (203 nm) is employed for the calculations, the SAR increases to 37.5 W g1 for the 12 nm MNPs and drops to 24.2 W g1 for the 13 nm MNPs. If the DH for the CNP, derived from the DLS data is used, the SAR remains unchanged. Experimentally, the SAR was measured to be 41.2 W g1 for the MNP, 72 W g1 for the CNP and 74.1 W g1 for the Dox-CNP samples. The deviation of the experimental values from the predicted numbers for the 12.9 nm bare MNPs can be attributed to the size distribution, which puts a significant fraction of the particles outside the optimum size range, consequently lowering the heating performance; this decrease in SAR due to polydispersity has been theoretically predicted [34]. Although the high experimental SAR values of CNPs and Dox-CNPs are welcome, the reason for the large increase compared to bare MNPs and theoretical predictions is unclear. Our analysis is consistent with experimental observations that with the choice of the appropriate DM, SAR will not be strongly reduced even if DH increases. Significantly, the ferrofluid concentrations used in this work are less than the range of 10–20 mg ml1 used by some previous researchers and much lower than 120 mg ml1 (of Fe) used for clinical trials [14,15,25]. By simply increasing the concentration
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Fig. 9. Typical heating curves of (a) nanoparticles dispersed in PBS and exposed to an AMF. Field strength = 4 kA m1, frequency = 375 kHz. Ferrofluid particle concentration = 2 mg ml1 (b) Dox-CNPs dispersed in PBS and exposed to an AMF for 35 min. Field strength = up to 5 kA m1, frequency = 375 kHz. Ferrofluid particle concentration = 1 mg ml1.
of Dox-CNPs in our ferrofluid, high SAR values can be obtained at moderate field strengths and frequencies; the predicted SAR value for a concentration of 20 mg ml1 is 284 W g1. 3.5. Temperature dependence of drug release Plots for cumulative dox release in PBS without an applied magnetic field for dehydrated as well as swollen Dox-CNPs at 24, 37 and 43 °C are shown in Fig. 10a and b, respectively. At the end of 29 h, the maximum bound dox released by the dehydrated DoxCNPs was 48.1% at 43 °C followed by 39.6% at 37 °C and 31.4% at 24 °C. The corresponding values for the swollen Dox-CNPs are 78.1% at 43 °C, 63.7% at 37 °C and 42.6% at 24 °C. The dox release curves for dehydrated Dox-CNPs in PBS without an AMF (Fig. 10a) show the limited influence of temperature on the release behavior. There is an initial rapid release phase up to 2 h during which 21.9%, 31.9% and 38.4% of the loaded dox is released at 24, 37 and 43 °C, respectively. Next, a steady release phase is observed. The relatively rapid drug release during the initial 1 h at all three temperatures may be due to convection-driven migration of the drug inside the PNIPAM matrix during the water evaporation stage of the vacuum dehydration process [56]. This leads to a heterogeneous distribution of dox in the PNIPAM shell, resulting in higher dox concentrations near the surface and rapid initial release. The release behavior of the dehydrated particles suggests that in order to preserve thermosensitive properties, dehydration
should be avoided during processing of dox-loaded composite MNPs. In contrast, the swollen Dox-CNPs show pronounced thermoresponsive behavior for drug release in PBS without an AMF (Fig. 10b). The rapid release phase extends up to 4 h resulting in release of 33.7%, 50.1% and 65.7% of the loaded dox at 24, 37 and 43 °C, respectively; this is followed by steady-state release. There is a greater difference in drug release at various temperatures as well as higher cumulative release above the LCST (63.7% at 37 °C and 78.1% at 43 °C for 29 h) compared to the dehydrated DoxCNPs, e.g. dox release from the swollen particles at 43 °C in just 1 h (51.2%) is higher than that from dehydrated particles after 29 h (48.1%). The higher release from the swollen particles is due to the collapse of the PNIPAM shell above the LCST (35 °C) and the squeezing of the dox molecules into the release medium along with the expulsion of hydrophobically bound water. 3.6. Controlled drug release The dox release from swollen Dox-CNPs in PBS under temperature switching is shown in Fig. 11a; the corresponding drug release rates with respect to initial dox loading are shown in Fig. 11b. After a total exposure of 20 h at 24 °C, and 4 h each at 37 and 43 °C, 65.4% of the loaded dox was released. The maximum release rate was 35.5% h1 for the first 1 h of exposure at 37 °C, this dropped drastically for the 24 °C exposures during the 2–4 h and 6–24 h time periods.
Fig. 10. Release curves for dox from (a) dehydrated Dox-CNPs (b) swollen Dox-CNPs dispersed in PBS and tested at 24, 37 and 43 °C.
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rate at 24 °C. Only an increase in temperature above the LCST (inducing matrix shrinkage) triggers further release. 4. Conclusions The synthesis, characterization and property evaluation of drug-loaded thermosensitive polymer-coated MNPs was examined in the context of multimodal cancer therapy. The hyperthermia behavior and dox release of these CNPs was studied. Magnetite (Fe3O4) nanoparticles of average size 12.9 nm (TEM) and size range of 10–16 nm were synthesized by a reverse co-precipitation method. The magnetic particles were encapsulated with the inverse thermosensitive polymer PNIPAM using dispersion polymerization to produce CNPs and subsequently loaded with the anti-cancer drug dox. The following conclusions were reached: Good magnetic properties of both coated and uncoated MNPs were observed. Swollen dox-loaded CNPs exhibited excellent thermoresponsive properties. Drug release could be controlled by cycling the temperature above/below the LCST; thus temperature can act as an on/off switch to regulate drug release. Dehydration of drug-loaded CNPs was found to be detrimental to controlled drug delivery. Good in vitro hyperthermia performance was achieved and hyperthermia temperatures attained with relatively low concentrations of particles. Simultaneous hyperthermia and drug release in therapeutically relevant doses was achieved—hence such dox-loaded thermoresponsive polymer-coated MNPs are promising for multimodal treatment of cancer.
Fig. 11. Controlled drug release from swollen, dox-loaded PNIPAM-coated MNPs as a function of temperature: (a) cumulative dox release and (b) dox release rates.
This thermoresponsive ability of the swollen Dox-CNPs is underlined by the controlled release tests. Initial exposure to 37 °C for 1 h results in release of 35.5% of the loaded dox, followed by a 43 °C, 1 h exposure yielding 48.5% release. When the temperature is reduced below the LCST to 24 °C for 2 h, the dox release is drastically reduced. When the temperature is again increased to 37 or 43 °C the drug release resumes, and a drop to 24 °C (6–24 h) causes drug release to reduce once again, suggesting that reswelling of PNIPAM below the LCST is relatively rapid. Thus, the temperature acts like an on/off switch to regulate drug release above and below the LCST. Comparison of the release curve of the swollen Dox-CNPs at 24 °C (Fig. 10b) and the controlled release curve segments at 24 °C (Fig. 11a) shows that the release exhibited by the former is much higher (>20% initial release) than that of the latter. This is explained by the presence of dox on the outer surface of the PNIPAM shell, this dox rapidly diffuses into the release media (PBS) when the particles are dispersed at 24 °C, independent of the thermoresponsive behavior. For the controlled release test, the swollen Dox-CNPs are initially held at temperatures of 37 and 43 °C for 1 h each, and the surface release process occurs rapidly during this period. Additionally, due to PNIPAM shrinkage above the LCST, dox molecules from the interior of the polymer matrix are also released, resulting in a lower concentration of dox in PNIPAM at the end of the first 2 h. This decreases the driving force for dox release when the temperature is reduced to 24 °C; at this temperature the polymer swells, preventing further transport of dox through the matrix into the release medium, resulting in a negligible release
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Thermoresponsive magnetic composite nanomaterials for multimodal cancer therapy S. Purushotham, R.V. Ramanujan * School of Materials Science and Engineering, Nanyang Technological University, Block N4.1, Nanyang Avenue, Singapore 639798, Singapore
a r t i c l e
i n f o
Article history: Received 25 March 2009 Received in revised form 17 June 2009 Accepted 7 July 2009 Available online 23 July 2009 Keywords: Magnetic nanoparticles Thermoresponsive materials Drug delivery Hyperthermia Cancer therapy
a b s t r a c t The synthesis, characterization and property evaluation of drug-loaded polymer-coated magnetic nanoparticles (MNPs) relevant to multimodal cancer therapy has been studied. The hyperthermia and controlled drug release characteristics of these particles was examined. Magnetite (Fe3O4)–poly-n-(isopropylacrylamide) (PNIPAM) composite MNPs were synthesized in a core–shell morphology by dispersion polymerization of n-(isopropylacrylamide) chains in the presence of a magnetite ferrofluid. These core–shell composite particles, with a core diameter of 13 nm, were loaded with the anti-cancer drug doxorubicin (dox), and the resulting composite nanoparticles (CNPs) exhibit thermoresponsive properties. The magnetic properties of the composite particles are close to those of the uncoated magnetic particles. In an alternating magnetic field (AMF), composite particles loaded with 4.15 wt.% dox exhibit excellent heating properties as well as simultaneous drug release. Drug release testing confirmed that release was much higher above the lower critical solution temperature (LCST) of the CNP, with a release of up to 78.1% of bound dox in 29 h. Controlled drug release testing of the particles reveals that the thermoresponsive property can act as an on/off switch by blocking drug release below the LCST. Our work suggests that these dox-loaded polymer-coated MNPs show excellent in vitro hyperthermia and drug release behavior, with the ability to release drugs in the presence of AMF, and the potential to act as agents for combined targeting, hyperthermia and controlled drug release treatment of cancer. Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction There has been enormous interest in the exciting area of magnetic nanoparticle (MNP)-based systems for cancer treatment [1–26] as conventional cancer therapies such as radiotherapy, hyperthermia and chemotherapy have serious drawbacks. The limitations of chemotherapy techniques include adverse effects on healthy tissue due to indiscriminate distribution of cytotoxic drugs in the body, insufficient local drug concentration in the tumor and poor control over drug release. In hyperthermia therapy, the tumor is heated in the temperature range of 41–47 °C, causing death of the cancerous cells, but sparing healthy cells [20,27,28]. In conventional hyperthermia techniques—e.g. radiofrequency-based hyperthermia, isolated hepatic perfusion, water baths and heating rods—temperature control is poor, heat distribution is not optimum and there is risk of organ damage due to overheating [20]. Drug-loaded composite MNPs offer a method of simultaneously achieving drug targeting, controlled drug release and hyperthermia treatment for cancer, thus overcoming several limitations of conventional cancer therapy.
* Corresponding author. Tel.: +65 67904342; fax: +65 67909081. E-mail address: [email protected] (R.V. Ramanujan).
The synergistic therapeutic effects of simultaneous chemotherapy and hyperthermia can exceed the individual or sequential application of these techniques [28–31]. Therefore, drug-loaded composite MNPs with a magnetic core and a polymer shell capable of acting as multifunctional agents for combined drug targeting, controlled release and hyperthermia therapy are highly desirable. Such particles can be injected into the appropriate blood vessels and targeted to the tumor by means of a suitable external magnetic field gradient [1,9,24,32,33]. Chemotherapy occurs by drug release from the composite particles which have been targeted to the tumor region. An external alternating magnetic field (AMF) can be applied to generate heat in the targeted particles [13,14,20,34], and the resulting temperature rise can be used for hyperthermia treatment of cancer. If the polymer shell of the nanoparticle is responsive to stimuli such as pH and temperature, drug release can also be controlled. This paper reports a study of novel iron oxide–poly-n-isopropylacrylamide (PNIPAM) composite nanoparticles (CNPs) suitable for magnetic targeting followed by simultaneous magnetic hyperthermia and chemotherapeutic drug release. The core of the CNP consists of iron oxide; PNIPAM is chosen as the thermosensitive polymer shell. PNIPAM is an inverse thermosensitive polymer that has been studied for several biomedical applications [35–41]; it undergoes reversible volume change in water at a lower critical solution temperature (LCST) in the range 30–35 °C. The proximity
1742-7061/$ - see front matter Ó 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2009.07.004
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of the LCST to the human body temperature (37 °C) and the ability to tune the LCST by addition of appropriate copolymers makes PNIPAM hydrogels useful for controlled drug delivery. Below the LCST, PNIPAM chains are soluble in water and the polymer is in a swollen state, inhibiting the transport of drug molecules through the matrix, resulting in a slow rate of release. Heat is generated in the iron oxide core upon exposure to an AMF, and this heat is conducted to the polymer shell, increasing the temperature of the PNIPAM. At and above the LCST, the shell collapses, squeezing out drug molecules, resulting in rapid drug release (Fig. 1). Both Fe3O4 (magnetite) and c-Fe2O3 (maghemite) are known to be biocompatible and non-toxic in concentrations relevant to clinical applications. In vivo experiments on animal models have shown that iron oxides are suitable for drug delivery and hyperthermia [1,17,20,22,42,43]. Human clinical trials for drug delivery [33] and hyperthermia [14,15,25] have been conducted with iron oxide-based ferrofluids, the injected dosage being well tolerated by patients. Superparamagnetic iron oxides (FDA approved) are also used as contrast-enhancing agents for clinical magnetic resonance imaging (MRI) [44–46]. In vitro drug release during AMF-induced heating has been previously studied using drug-loaded thermosensitive carriers [4,6,47] and dye-loaded PNIPAM–magnet systems [3,48]. Drug release from dox-loaded magnetic particles encapsulated by a thermosensitive polymer without an applied magnetic field has recently been studied [26]. The novelty of this work is in the synthesis, characterization and property evaluation of thermosensitive polymer-encapsulated magnetic particles relevant to combined magnetic targeting, hyperthermia and drug release applications. Previously, we have shown that drug-loaded CNPs can be targeted to hepatocellular carcinoma (HCC) in a buffalo rat model, and localization in the tumor was confirmed by MRI and histology [9]. 2. Materials and methods 2.1. Synthesis of MNPs Fe3O4 MNPs were synthesized by a reverse co-precipitation method adapted from Aono et al. [49] to obtain particles with an
Swollen PNIPAM shell Slow drug release
Drug loaded PNIPAM shell
T > LCST
T < LCST
T < LCST
Hyperthermia T > LCST MNP cluster
PNIPAM shell shrinks Faster drug release Fig. 1. Schematic diagram showing drug release from PNIPAM-coated MNPs below and above the polymer LCST.
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average size of 10–14 nm. In a typical experiment, FeCl36H2O (0.1 M, 27 g) (Merck) and FeCl24H2O (0.05 M, 10 g) (Merck) were dissolved in 200 ml of ultrapure MilliQ water and heated to 80 °C. The synthesis was performed inside a glovebox in an inert atmosphere, and nitrogen was bubbled through all solutions for 30 min before use. Under vigorous stirring, 140 ml of NaOH solution (0.14 M, 5.6 g in 140 ml) (Merck) was dripped into this solution at a rate of 6 ml min1, upon which the solution turned black, indicating the precipitation of Fe3O4 (magnetite). The mixture was maintained at 80 °C for 30 min, cooled to room temperature and centrifuged at 8000 rpm for 15 min. The supernatant was discarded and the particles were extracted. After washing several times with MilliQ water and ethanol to remove residual reactants, the particles were dried in vacuum for 24 h. 2.2. Size selection by magnetic separation The dried MNPs were ultrasonically redispersed in 500 ml of ethanol until a dark suspension was obtained. A 0.8 T block permanent magnet was placed on the outer wall of the beaker to attract the most magnetically responsive particles in the suspension. The supernatant was extracted with a peristaltic pump (Simon, Manostat). The responsive particles were redispersed in ethanol and the process was repeated three times. The magnetically separated particles were dried in vacuum and stored in a vacuum oven at room temperature. 2.3. Preparation of iron oxide–polymer CNPs CNPs were prepared, in the presence of MNPs, by free radical dispersion polymerization of NIPAM monomer in water with N,Nmethylene(bis)acrylamide (BAAm) as the cross-linker, N,N,N0 ,N0 -tetramethylene diamine (TEMED) as the accelerator and ammonium persulfate (APS) as the initiator [3,50]. Our procedure does not involve surfactants which can alter the LCST [40,51]. MilliQ water was used, through which nitrogen was bubbled for 15 min prior to use. The NIPAM monomer was purified by recrystallization from hexane and dried in vacuum for 24 h before use. In a typical experiment 35 ml of ferrofluid containing 850 mg MNP was added to a solution of 960 mg NIPAM (Aldrich, 97%) in 30 ml water and ultrasonicated for 10 min. The mixture was transferred to a beaker and stirred at 24,000 rpm by an IKA-Werke T25 Ultra Turrax high-speed dispersion unit while 2.5 ml of BAAm (Fluka, 98%) solution (8 mg in 5 ml) and 0.5 ml of APS (J.T. Baker, 98.5%) solution (180 mg in 5 ml) were added to the mixture. Finally 1 ml of TEMED (Lancaster, 99%) was added dropwise with continuous stirring. After 20 min of stirring, the mixture was left undisturbed for 2 h, after which the CNPs were magnetically separated and washed several times with MilliQ water. Some of the CNPs were dried in vacuum for 24 h, resulting in dehydration of the particles: the remainder were used as-synthesized in the swollen state for dynamic light scattering (DLS) and microcalorimetry experiments. 2.4. Drug loading The water-soluble anti-cancer drug doxorubicin (dox) was chosen as a model drug. Typically, 26.5 mg (loading target 10 wt.% of CNP) of doxorubicin hydrochloride was dissolved in 20 ml of MilliQ by ultrasonication for 10 min. Two hundred and forty-five milligrams of dehydrated CNPs was added to the dox solution and ultrasonicated for an additional 10 min. The bottle was sealed and gently shaken in a rotary shaker in the dark for 20 h at 24 °C to facilitate both reswelling of the PNIPAM shell in the aqueous dox solution as well as to facilitate dox uptake. After loading, the dox-loaded CNPs (Dox-CNPs) were magnetically separated. To determine residual dox content, the supernatant was analyzed in
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a UV–visible spectrophotometer (Shimadzu UV1700) and compared with reference plots of dox in MilliQ water at a wavelength of 482 nm. The drug loading of the Dox-CNPs was determined as the difference between the dox content of the solution before loading and the dox content of the supernatant. To investigate the effect of dehydration on the particle performance, part of the batch of prepared particles was dried in vacuum for 48 h and the remainder was used in the swollen condition for experiments. 2.5. Phase identification and crystallite size X-ray diffraction scans were performed on the dried powder using a Shimadzu 6000 X-ray diffractomer with CuKa radiation of wavelength 1.54056 Å, in the 2h range 10–90° at a scan rate of 1° min1. Phase identification was performed by matching peak positions and relative intensities to reference JCPDS files. The crystallite size was calculated using the Scherrer formula. 2.6. Particle size and hydrodynamic diameter The particle size of the bare MNPs was measured by transmission electron microscopy. A drop of MNP ferrofluid was placed on a carbon-coated copper TEM grid, dried in vacuum and imaging was carried out in a JEOL 2010F microscope operating at 200 kV. Particle size was measured from transmission electron micrographs using ImageJ image processing software. The size of the CNP particles could not be reliably examined by TEM since the electron beam caused ablation of the polymer shell. DLS was used to measure the hydrodynamic diameters of aggregates of MNP and swollen CNP samples in MilliQ water. Dilute suspensions (<0.2 mg ml1) were transferred to low-volume quartz cuvettes and investigated using a Malvern Instruments Zetasizer Nano DLS unit. The samples were equilibrated at 25 °C for 5 min before each measurement. 2.7. Thermogravimetric analysis (TGA) Twenty milligrams of dehydrated CNP powder was placed in a ceramic pan, heated to 600 °C at a rate of 20 °C min1 under nitrogen flow in a TA Instruments TGA Q500 thermogravimetric analyzer, and the weight loss was measured as a function of temperature. To determine the polymer content of the CNPs, the weight loss plot of bare MNPs measured under identical conditions was used as reference.
2.9. Magnetic properties Magnetic properties of bare and dox-loaded composite MNP powders were evaluated using a LakeShore 7404 vibrating sample magnetometer, in an applied field range of 0–10 kOe. Magnetic particle diameter was calculated by fitting the magnetization curve to the Langevin function [53]:
M lH kB T ¼ coth ; Ms kB T lH
ð1Þ
where M is the magnetization for a field strength H, Ms is the saturation magnetization, kB is the Boltzmann constant, the true magnetic moment of each particle l = (MspD3/6) and D is the magnetic particle diameter. 2.10. In vitro heating behavior in an AMF MNP, swollen CNP and swollen Dox-CNP powders were ultrasonically dispersed in PBS to prepare ferrofluids with a particle concentration of 2 mg ml1. Three milliliters of ferrofluids contained in glass bottles were placed inside a water-cooled five-loop copper induction coil energized by an AC generator (Inductelec, UK) with operating frequency of 375 kHz and field strength of 4 kA m1. Ceramic wool thermal insulation was used around the bottles to minimize heat loss, and temperature data was continuously logged using a Luxtron MD600 fiber optic thermometry unit connected to a laptop computer; at least three samples were tested for each type of ferrofluid. The heating power, or specific absorption rate (SAR), of the particles was calculated from the equation:
SAR ¼ C
DT massferrofluid ; Dt massnanoparticles
ð2Þ
where C is the mass-weighted heat capacity of the ferrofluid, and ðDT=DtÞ is the slope of the initial section of the temperature vs. time curve [34]. In order to study drug release during hyperthermia, additional experiments were carried out using dehydrated Dox-CNPs. The AMF strength was manually varied to raise the temperature in less than 5 min and maintain the temperature in the range between 41 and 48 °C for various time intervals up to 60 min. Samples of 5 ml of ferrofluid with a Dox-CNP concentration of 1 mg ml1 were used. At the end of each heating experiment, the particles were magnetically separated from the suspension, and drug release was quantified by measuring the dox content in the supernatant via UV–visible spectrophotometry.
2.8. Measurement of LCST
2.11. Temperature-dependent drug release without AMF
The determination of LCST for PNIPAM aqueous solutions is often performed by optical light transmission methods (e.g. UV–visible spectrophotometry) or microcalorimetry [51,52]. However, for composite MNPs optical methods do not yield reliable results at concentrations of interest (>1 mg ml1), and so LCST measurements were performed with a TA Instruments DSC 2920 modulated differential scanning calorimeter (MDSC). Swollen CNPs as well as Dox-CNPs, and dehydrated Dox-CNPs dispersed in PBS, were studied. The dehydrated particles were dispersed in PBS and refrigerated at 8 °C for 4 h before testing to enable reswelling of the polymer shell. Reference pans contained ferrofluid of MNPs without any polymer. A few microliters of the sample suspensions were placed in hermetically sealed aluminum pans and heated from 20 to 50 °C at a scanning rate of 0.5 °C min1 with modulation amplitude of 0.2 °C min1. The pans were equilibrated in the DSC cell at 20 °C for 20 min before a heating run. The MDSC thermograms were analyzed by TA Universal Analysis software.
Drug release for extended time periods without an AMF was studied at temperatures of 24 °C (below LCST), 37 °C (body temperature, near LCST) and 43 °C (above LCST). Samples of 5 ml of ferrofluids with concentration of 1 mg ml1 were prepared by ultrasonically dispersing Dox-CNP powders in PBS (pH 7.4) in glass test-tubes. A minimum of three samples were prepared for testing at each temperature. The sealed test-tubes were placed in preheated water baths (Fisher Isotemp 202) and the temperature stabilized (±1 °C) at 24, 37 and 43 °C for a total of 29 h. Periodically, 1 ml of the release medium, free of nanoparticles, was extracted and replaced with fresh PBS. Dox content of the extracted solution was measured by UV–visible spectrophotometry and the cumulative dox release calculated. 2.12. Controlled drug release without AMF The ability of swollen Dox-CNPs to modulate drug release by varying temperature was studied. The experimental details are as
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described in Section 2.11; the temperature was cycled between 24, 37 and 43 °C for various time periods and dox release measured via UV–visible spectrophotometry. 2.13. Modeling of power output from MNPs in an AMF The model proposed by Rosensweig [34] was used to examine the effect of polymer shell thickness on the heating power of MNPs exposed to an AMF. The following values were used for the calculations: ferrofluid concentration = 2 mg ml1 of monodisperse spherical magnetite nanoparticles; magnetite domain magnetization Md = 446 kA m1 [34]; anisotropy constant K = 32 kJ m3 [34]; temperature T = 310 K (37 °C); g = 6.92 104 N s m2 (at 37 °C); frequency f = 375 kHz; field strength H0 = 4 kA m1. The hydrodynamic diameter is taken to be the magnetic particle diameter for MNPs and the overall particle diameter for CNPs. The analysis does not incorporate polydispersity, viscous and hysteresis losses, interparticle interactions and changes in viscosity with increasing temperature. 3. Results and discussion 3.1. Particle size and structure Fig. 2 shows a transmission electron micrograph of the bare MNP. The average size of bare nanoparticles after magnetic size selection was calculated to be 11.9 nm from the XRD data, 10.1 nm from the magnetization curve, and 12.9 nm measured from transmission electron micrographs (standard deviation = 3). Fig. 3 shows the XRD pattern, which indicates that the particles are crystalline; peak matching with reference JCPDS files shows that the particles are primarily Fe3O4 (magnetite). The average hydrodynamic diameter (DH) and polydispersity index (PDI) of MNPs and CNPs in water at 25 °C were calculated from an exponential fit to the intensity correlation function of the light scattering data obtained from DLS experiments. The bare MNPs had an average DH of 203 nm with a PDI of 0.301, the corresponding figures for the CNPs were 158 nm and 0.161, respectively. From theoretical calculations of heat generation in an AMF, the optimum average size (DM) of the MNPs is 12 nm with a size dis-
Fig. 3. X-ray diffraction pattern of bare Fe3O4 nanoparticles. Numbers in brackets indicate the planes of diffraction.
tribution in the range of 12–13 nm. With this particle size and size distribution, an increase in hydrodynamic diameter DH due to PNIPAM incorporation would not significantly affect heat generation. The influence of the particle hydrodynamic diameter (DH) on SAR for selected values of DM is shown in Fig. 4. For particles with a core diameter less than 12 nm there is a small increase in SAR as DH is increased. For 12 nm DM particles, SAR increases gradually with increasing DH. When DM is larger than 12 nm, the increase in DH is accompanied by a sharp decrease in SAR. Néel losses will be maximum for the 12–13 nm core size. At this size, although Brownian relaxation losses are low, there is no sharp reduction in SAR if DH increases either due to clustering or excessive polymer coating during synthesis. For discrete, bare MNPs, DH is usually larger than DM, so the same conditions apply. The average MNP size synthesized in our experiments, 13 nm, is in the size range predicted from our calculations to yield high SAR. DLS shows a larger hydrodynamic diameter and wider distribution for agglomerates of bare MNPs (203 nm) compared to those of the CNPs (158 nm) due to the following: bare iron oxide nanoparticles, without an appropriate surfactant or stabilizing polymer layer, have a hydrophobic surface which induces clustering in aqueous suspensions as they try to reduce their surface area in contact with water [11]. Additionally, due to the broader size distribution of these particles, strong attractive interparticle magnetic forces will encourage clustering. On the other hand, below the LCST of PNIPAM, the CNPs and clusters are partially stabilized by the hydrophilic polymer shell, and the shell also reduces the attraction between particle clusters. Since DLS depends on scattered light intensity from aggregates, it is highly sensitive to larger particle sizes (d6 dependence of Rayleigh scattering), and scattering from relatively few but larger particle clusters tends to skew the reported average size; this can explain why the CNP samples show a narrower size distribution (polydispersity index of 0.161 compared to 0.301 for MNP). 3.2. PNIPAM and dox content of CNP
Fig. 2. Bright-field transmission electron micrograph of Fe3O4 nanoparticles produced by the reverse co-precipitation method. Aggregation is due to the attractive magnetic forces between the particles and the absence of a surfactant layer.
The TGA weight loss vs. temperature plot for both MNPs and CNPs is shown in Fig. 5. Heating in the range of room temperature to 130 °C results in expulsion of water from the polymer, and PNIPAM decomposition occurs between 200 and 440 °C [54]. From the TGA data, the PNIPAM content of CNPs is estimated to be 1.3 wt.%. The MNP sample shows less than 0.5% weight loss, confirming that the weight change for the CNP sample beyond 200 °C is due to PNIPAM degradation. The UV–visible spectrophotometry results (not shown) revealed that the drug loading of the Dox-CNPs was 4.16 wt.%. Polymer encapsulation of MNPs is essential for loading drugs or other therapeutic agents onto MNPs; however, the magnetic prop-
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Fig. 4. Influence of hydrodynamic diameter DH on the SAR of Fe3O4 nanoparticles in ferrofluids for selected magnetic diameter DM values, at f = 375 kHz and H0 = 4 kA m1.
Fig. 5. TGA weight loss curves for bare MNPs (top) and PNIPAM-encapsulated MNPs (bottom).
erties of the composite particles can be inferior to those of the MNPs, hence the weight fraction of non-magnetic components must be kept to a minimum. The PNIPAM content of 1.3 wt.% in the CNPs was found to be adequate to achieve loading (4.16 wt.% Dox-CNP) and delivery of therapeutic quantities of dox with little adverse effect on magnetic properties. In dispersion polymerization, both heterogeneous and homogeneous nucleation occurs [50] with nucleation and growth of the polymeric shell around the iron oxide nanoparticle core. Bonding between the iron oxide ions and surface hydroxy groups to the free ion pair of amidic nitrogen in NIPAM aids adsorption of the monomer to the surfaces of the nanoparticles. Nucleation and growth of PNIPAM around magnetic clusters occurred rapidly and uniformly due to the combination of a high weight ratio of iron oxide to NIPAM (85:96), high stirring speed (24,000 rpm) and the presence of accelerator TEMED yielding composite particles in the nanoscale range with relatively low PDI. The volume fraction of PNIPAM in CNPs is much higher than 1.3% due to its lower density compared to Fe3O4. When drug loading is expressed in terms of PNIPAM weight, the 4.16 wt.% dox content of the CNPs is equivalent to 320 wt.% of PNIPAM; this value is
due to the large surface area to volume ratio of the polymer shell in the nanoscale range which enhances the adsorption of dox. The modulated DSC thermograms for swollen CNPs and DoxCNPs are shown in Fig. 6a and b, respectively; those for the dehydrated Dox-CNPs in Fig. 7. The LCST was measured as the onset temperature as this temperature, unlike the peak maximum temperature, is relatively insensitive to sample weight related artifacts [52]. The swollen particles exhibit two distinct endothermic peaks (Fig. 6a and b) at the LCST. For the swollen CNPs (Fig. 6a), the onset of the first endothermic event is at 35.3 °C, and the endothermic peak is at 35.8 °C. The second transition has an onset temperature of 39.4 °C and an endothermic peak at 39.9 °C. For the swollen Dox-CNPs, the onset temperatures of the events are 35 and 42 °C, with peaks at 35.5 and 43.5 °C, respectively. The thermogram of the dehydrated particles (Fig. 7) lacks the distinct peaks exhibited by those of the swollen particles, indicating the absence of a welldefined transition temperature. For the PNIPAM homopolymer, the LCST in aqueous solution is strongly influenced by additives that can bind to the polymer molecules or affect the structure of water around them [40,51]. Fig. 6a shows an LCST of 35.3 °C for the swollen CNPs in PBS, which indicates that the MNPs in PNIPAM does not significantly alter the LCST despite the hydrophobic surface of the MNP and the tendency of the ions in the buffer to interact with the PNIPAM chains. Another transition at 39.4 °C may be due to local changes in the structure of PNIPAM during dispersion polymerization. The swollen Dox-CNPs also show two endothermic events, the first of which has an onset temperature of 35 °C, close to the LCST of the CNP. Interestingly, the second and stronger transition has an onset temperature of 42 °C. The shift in the temperature for the second transition was due to the loading of the hydrophilic dox in the PNIPAM matrix, resulting in hydrogen bonding between dox and the amide group of PNIPAM. In contrast, the dehydrated Dox-CNP sample (Fig. 7) does not exhibit any distinct transition temperature, and vacuum dehydration causes a significant change in the PNIPAM structure, leading to the absence of a LCST. 3.3. Magnetic properties The MNP and Dox-CNP magnetization curves at room temperature suggest similar magnetic properties (Fig. 8). The bare MNPs
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Fig. 6. Modulated DSC thermograms for swollen (a) CNPs dispersed in PBS (b) Dox-CNPs dispersed in PBS.
3.4. In vitro heating behavior in an AMF
Fig. 7. Modulated DSC thermogram for dehydrated Dox-CNPs dispersed in PBS.
exhibit a saturation magnetization (MS) of 61 emu g1; this decreases to 56.8 emu g1 for the Dox-CNPs. The coercivity (Hc) of Dox-CNPs is 8.5 Oe, lower than the Hc of 43 Oe exhibited by the bare MNPs. The MS values for MNPs and Dox-CNPs are lower than those of bulk magnetite [55]. Assuming an anisotropy value of 32 kJ m3 [34], at 27 °C, Fe3O4 nanoparticles smaller than 15 nm are expected to be superparamagnetic and should display a sigmoidal anhysteric magnetization curve. But due to size distribution and clustering effects the experimental magnetization curves display a small amount of hysteresis (Fig. 8). The decrease in coercivity from 43 to 8.5 Oe for the Dox-CNP samples is due to the role of the polymer in reducing aggregation of MNPs (Section 3.2).
Fig. 8. Room temperature magnetization curves for bare MNPs and Dox-CNPs for field strengths up to 10 kOe.
The temperature rise from the ambient temperature plotted against time for MNP, CNP and Dox-CNP ferrofluids exposed to a 375 kHz, 4 kA m1 AMF is shown in Fig. 9a. The bare MNPs reaches the hyperthermia temperature (41 °C) in 335 s, while the CNPs and Dox-CNPs attain this temperature more rapidly, in 295 and 255 s, respectively. From the ferrofluid concentration and slope of the heating curve up to 120 s, the SAR was calculated to be 41.2 W g1 for MNP, 72 W g1 for CNP and 74.1 W g1 for DoxCNP samples. A typical in vitro simultaneous hyperthermia and drug release curve, in this case for 30 min of hyperthermia exposure, is shown in Fig. 9b. Compared to the heating behavior with 2 mg ml1 ferrofluids, the lower (1 mg ml1) concentration of Dox-CNP ferrofluid reaches 41 °C in 240 s due to the higher applied field strength and ambient temperature. The temperature could be readily stabilized between 41 and 47 °C for clinically relevant time periods of 60 min [14,15] by manually varying the field strength. After exposure to hyperthermia temperatures (41–47 °C) for 10 min, 26.2% of the bound dox was released from the Dox-CNPs, increasing to 32.9% for 30 min and 36.1% for 60 min of exposure. This clearly demonstrates that hyperthermia temperatures and simultaneous drug release in therapeutic quantities can be achieved in vitro by the Dox-CNP ferrofluids. The predicted SAR values for 12 and 13 nm MNPs for the conditions stated in Section 2.13, and assuming DH = DM, are 28.4 W g1 and 50.3 W g1, respectively. When the hydrodynamic diameter measured by DLS at 25 °C (203 nm) is employed for the calculations, the SAR increases to 37.5 W g1 for the 12 nm MNPs and drops to 24.2 W g1 for the 13 nm MNPs. If the DH for the CNP, derived from the DLS data is used, the SAR remains unchanged. Experimentally, the SAR was measured to be 41.2 W g1 for the MNP, 72 W g1 for the CNP and 74.1 W g1 for the Dox-CNP samples. The deviation of the experimental values from the predicted numbers for the 12.9 nm bare MNPs can be attributed to the size distribution, which puts a significant fraction of the particles outside the optimum size range, consequently lowering the heating performance; this decrease in SAR due to polydispersity has been theoretically predicted [34]. Although the high experimental SAR values of CNPs and Dox-CNPs are welcome, the reason for the large increase compared to bare MNPs and theoretical predictions is unclear. Our analysis is consistent with experimental observations that with the choice of the appropriate DM, SAR will not be strongly reduced even if DH increases. Significantly, the ferrofluid concentrations used in this work are less than the range of 10–20 mg ml1 used by some previous researchers and much lower than 120 mg ml1 (of Fe) used for clinical trials [14,15,25]. By simply increasing the concentration
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Fig. 9. Typical heating curves of (a) nanoparticles dispersed in PBS and exposed to an AMF. Field strength = 4 kA m1, frequency = 375 kHz. Ferrofluid particle concentration = 2 mg ml1 (b) Dox-CNPs dispersed in PBS and exposed to an AMF for 35 min. Field strength = up to 5 kA m1, frequency = 375 kHz. Ferrofluid particle concentration = 1 mg ml1.
of Dox-CNPs in our ferrofluid, high SAR values can be obtained at moderate field strengths and frequencies; the predicted SAR value for a concentration of 20 mg ml1 is 284 W g1. 3.5. Temperature dependence of drug release Plots for cumulative dox release in PBS without an applied magnetic field for dehydrated as well as swollen Dox-CNPs at 24, 37 and 43 °C are shown in Fig. 10a and b, respectively. At the end of 29 h, the maximum bound dox released by the dehydrated DoxCNPs was 48.1% at 43 °C followed by 39.6% at 37 °C and 31.4% at 24 °C. The corresponding values for the swollen Dox-CNPs are 78.1% at 43 °C, 63.7% at 37 °C and 42.6% at 24 °C. The dox release curves for dehydrated Dox-CNPs in PBS without an AMF (Fig. 10a) show the limited influence of temperature on the release behavior. There is an initial rapid release phase up to 2 h during which 21.9%, 31.9% and 38.4% of the loaded dox is released at 24, 37 and 43 °C, respectively. Next, a steady release phase is observed. The relatively rapid drug release during the initial 1 h at all three temperatures may be due to convection-driven migration of the drug inside the PNIPAM matrix during the water evaporation stage of the vacuum dehydration process [56]. This leads to a heterogeneous distribution of dox in the PNIPAM shell, resulting in higher dox concentrations near the surface and rapid initial release. The release behavior of the dehydrated particles suggests that in order to preserve thermosensitive properties, dehydration
should be avoided during processing of dox-loaded composite MNPs. In contrast, the swollen Dox-CNPs show pronounced thermoresponsive behavior for drug release in PBS without an AMF (Fig. 10b). The rapid release phase extends up to 4 h resulting in release of 33.7%, 50.1% and 65.7% of the loaded dox at 24, 37 and 43 °C, respectively; this is followed by steady-state release. There is a greater difference in drug release at various temperatures as well as higher cumulative release above the LCST (63.7% at 37 °C and 78.1% at 43 °C for 29 h) compared to the dehydrated DoxCNPs, e.g. dox release from the swollen particles at 43 °C in just 1 h (51.2%) is higher than that from dehydrated particles after 29 h (48.1%). The higher release from the swollen particles is due to the collapse of the PNIPAM shell above the LCST (35 °C) and the squeezing of the dox molecules into the release medium along with the expulsion of hydrophobically bound water. 3.6. Controlled drug release The dox release from swollen Dox-CNPs in PBS under temperature switching is shown in Fig. 11a; the corresponding drug release rates with respect to initial dox loading are shown in Fig. 11b. After a total exposure of 20 h at 24 °C, and 4 h each at 37 and 43 °C, 65.4% of the loaded dox was released. The maximum release rate was 35.5% h1 for the first 1 h of exposure at 37 °C, this dropped drastically for the 24 °C exposures during the 2–4 h and 6–24 h time periods.
Fig. 10. Release curves for dox from (a) dehydrated Dox-CNPs (b) swollen Dox-CNPs dispersed in PBS and tested at 24, 37 and 43 °C.
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rate at 24 °C. Only an increase in temperature above the LCST (inducing matrix shrinkage) triggers further release. 4. Conclusions The synthesis, characterization and property evaluation of drug-loaded thermosensitive polymer-coated MNPs was examined in the context of multimodal cancer therapy. The hyperthermia behavior and dox release of these CNPs was studied. Magnetite (Fe3O4) nanoparticles of average size 12.9 nm (TEM) and size range of 10–16 nm were synthesized by a reverse co-precipitation method. The magnetic particles were encapsulated with the inverse thermosensitive polymer PNIPAM using dispersion polymerization to produce CNPs and subsequently loaded with the anti-cancer drug dox. The following conclusions were reached: Good magnetic properties of both coated and uncoated MNPs were observed. Swollen dox-loaded CNPs exhibited excellent thermoresponsive properties. Drug release could be controlled by cycling the temperature above/below the LCST; thus temperature can act as an on/off switch to regulate drug release. Dehydration of drug-loaded CNPs was found to be detrimental to controlled drug delivery. Good in vitro hyperthermia performance was achieved and hyperthermia temperatures attained with relatively low concentrations of particles. Simultaneous hyperthermia and drug release in therapeutically relevant doses was achieved—hence such dox-loaded thermoresponsive polymer-coated MNPs are promising for multimodal treatment of cancer.
Fig. 11. Controlled drug release from swollen, dox-loaded PNIPAM-coated MNPs as a function of temperature: (a) cumulative dox release and (b) dox release rates.
This thermoresponsive ability of the swollen Dox-CNPs is underlined by the controlled release tests. Initial exposure to 37 °C for 1 h results in release of 35.5% of the loaded dox, followed by a 43 °C, 1 h exposure yielding 48.5% release. When the temperature is reduced below the LCST to 24 °C for 2 h, the dox release is drastically reduced. When the temperature is again increased to 37 or 43 °C the drug release resumes, and a drop to 24 °C (6–24 h) causes drug release to reduce once again, suggesting that reswelling of PNIPAM below the LCST is relatively rapid. Thus, the temperature acts like an on/off switch to regulate drug release above and below the LCST. Comparison of the release curve of the swollen Dox-CNPs at 24 °C (Fig. 10b) and the controlled release curve segments at 24 °C (Fig. 11a) shows that the release exhibited by the former is much higher (>20% initial release) than that of the latter. This is explained by the presence of dox on the outer surface of the PNIPAM shell, this dox rapidly diffuses into the release media (PBS) when the particles are dispersed at 24 °C, independent of the thermoresponsive behavior. For the controlled release test, the swollen Dox-CNPs are initially held at temperatures of 37 and 43 °C for 1 h each, and the surface release process occurs rapidly during this period. Additionally, due to PNIPAM shrinkage above the LCST, dox molecules from the interior of the polymer matrix are also released, resulting in a lower concentration of dox in PNIPAM at the end of the first 2 h. This decreases the driving force for dox release when the temperature is reduced to 24 °C; at this temperature the polymer swells, preventing further transport of dox through the matrix into the release medium, resulting in a negligible release
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