Magnetically Induced Localized On-Demand Drug Delivery

Magnetically induced localized on-demand drug delivery

Citation for published version (APA): Rovers, S. A. (2010). Magnetically induced localized on-demand drug delivery. Technische Universiteit Eindhoven. https://doi.org/10.6100/IR674220

DOI: 10.6100/IR674220

Document status and date: Published: 01/01/2010

Document Version: Publisher’s PDF, also known as Version of Record (includes final page, issue and volume numbers)

Please check the document version of this publication:

• A submitted manuscript is the version of the article upon submission and before peer-review. There can be important differences between the submitted version and the official published version of record. People interested in the research are advised to contact the author for the final version of the publication, or visit the DOI to the publisher's website. • The final author version and the galley proof are versions of the publication after peer review. • The final published version features the final layout of the paper including the volume, issue and page numbers. Link to publication

General rights Copyright and moral rights for the publications made accessible in the public portal are retained by the authors and/or other copyright owners and it is a condition of accessing publications that users recognise and abide by the legal requirements associated with these rights.

• Users may download and print one copy of any publication from the public portal for the purpose of private study or research. • You may not further distribute the material or use it for any profit-making activity or commercial gain • You may freely distribute the URL identifying the publication in the public portal.

If the publication is distributed under the terms of Article 25fa of the Dutch Copyright Act, indicated by the “Taverne” license above, please follow below link for the End User Agreement: www.tue.nl/taverne

Take down policy If you believe that this document breaches copyright please contact us at: [email protected] providing details and we will investigate your claim.

Download date: 05. Oct. 2021

Magnetically induced localized on-demand drug delivery

PROEFSCHRIFT

ter verkrijging van de graad van doctor aan de Technische Universiteit Eindhoven, op gezag van de rector magniﬁcus, prof.dr.ir. C.J. van Duijn, voor een commissie aangewezen door het College voor Promoties in het openbaar te verdedigen op woensdag 2 juni 2010 om 16.00 uur

door

Stefan Adrianus Rovers

geboren te Helmond Dit proefschrift is goedgekeurd door de promotoren: prof.dr.ir. J.T.F. Keurentjes en prof.dr.ir. K. Kopinga

Copromotor: dr.ir. R. Hoogenboom

c 2010, Stefan Rovers

A catalogue record is available from the Eindhoven University of Technology Library PhD Thesis.– ISBN 978-90-386-2243-9

Magnetically induced localized on-demand drug delivery / by Stefan A. Rovers Eindhoven University of Technology, Eindhoven, The Netherlands, 2010.

Printed by Universiteitsdrukkerij Technische Universiteit Eindhoven Cover design: Paul Verspaget

Summary

Externally triggered on-demand drug release from an implant can significantly improve the efficiency of the drug therapy since it enables the patient or physician to control the dosing to the patient’s needs and releases the drug only at the required location in the human body. Therefore, patient compliance and efficacy will increase and toxic side effects decrease as untargeted locations are not exposed to significant drug levels as is often the case in systemic drug administration. In this work, the externally triggered drug delivery system is a thermoresponsive polymeric implant triggered using an alternating magnetic field. The thermal switch is based on a significant change in diffusivity of a solute around the glass transition temperature (Tg) of a polymer. At a temperature below the glass transition temperature of the polymer (T < Tg), the polymer is in a glassy state and the diffusion coefficient of the incorporated drug is low, limiting drug release. Increasing the temperature to above the Tg of the polymer (T > Tg), the polymer becomes rubbery. This significantly increases the flexibility and free volume of the polymer resulting in release of the active. Since the glass transition is a reversible transition, subsequent lowering of the temperature significantly decreases the drug release rate from the implant, enabling pulsatile drug administration. The temperature of the implant is increased using an externally applied alternating magnetic field. In order to increase the temperature of the implant using a magnetic field, the use of superparamagnetic iron oxide nanoparticles (SPION) is explored. These nanoparticles are used as MRI contrast agents and to locally increase the temperature in hyperthermia treatment, the destruction of tumors by elevated temperature. The particles have no remanent magnetization, are biocompatible and are able to generate thermal energy using an alternating magnetic field because of Néel and Brown relaxation. Néel relaxation is the reorientation of the magnetic moment within the particles, generating thermal energy by crossing an anisotropy barrier, and Brown relaxation the reorientation of the magnetic particles itself, generating thermal energy by viscous friction with the carrier fluid. Summary

Since the nanoparticles are used for heating a polymer implant, different preparation methods for an iron oxide - polymer nanocomposite have been investigated. Freeze drying a mixture of a ferrofluid with a poly(methyl methacrylate) (p(MMA)) latex and subsequent compounding, results in an optimal distribution of the particles. It is expected that the particles do not agglomerate because of the combination of stabilization of both the iron oxide particles and polymer latex by surfactants, and the lack of mobility during freeze drying. Other methods used, e.g. solvent casting and direct injection of the ferrofluid into the compounder, result in significant agglomeration of the particles. Subsequently, the particle distribution has been shown to have a significant effect on the heating of the particle. An optimal distribution of the particles results in the highest specific absorption rate (SAR), the amount of thermal energy generated per gram of iron oxide, because of a minimum in interparticle interactions. Since the nanoparticles incorporated in a polymer are immobilized, the particles are not able to generate thermal energy by Brown relaxation. By a direct comparison of the specific absorption rate of particles suspended in liquid and incorporated in p(MMA) using the optimal freeze drying method, the contribution of both Néel and Brown relaxation to the heating of SPION has been investigated. Since the observed SAR is identical in both situations, it is concluded that at the frequency used (745 kHz), Néel relaxation is the only relaxation process that contributes to the heating of the particles in ferrofluid, because of the significantly shorter relaxation time for Néel relaxation. Using a cylindrical core of iron oxide - p(MMA) nanocomposite, coated with a thermoresponsive poly(butyl methacrylate-stat-methyl methacrylate) (p(BMA- MMA)) layer, externally triggered on-demand drug release has been investigated. A model drug, ibuprofen, has been incorporated in the thermoresponsive p(BMA- MMA) coating. Upon exposure of the sample to an alternating magnetic field (on situation), the drug release rate is significantly increased compared to the release rate without the magnetic field (off situation). After the magnetic field is removed, the release rate decreases back to the rate prior to the exposure, demonstrating the reversibility of the system. Multiple consecutive exposures to the external trigger result in similar increases of the release rate. Increasing the iron oxide concentration in the core of the device increases the release rate upon exposure, whereas the release rate without exposure is not influenced, therefore increasing the on/off ratio, because of a higher temperature increase upon exposure. Even though externally triggered pulsatile drug release has been shown, the maximum on/off ratio obtained is only 16.5. This relatively low ratio is primarily due to the suboptimal nature of

ii Summary the used commercially available iron oxide and the relative high off release rate of ibuprofen from p(BMA-MMA). In order to increase the on/off ratio, a cylindrical iron core has been used, coated with an ibuprofen incorporated poly(styrene-stat-butyl methacrylate) (p(S-BMA)) layer. The heat generated in the iron core upon exposure to the magnetic field is due to induction heating. Externally triggered pulsatile drug release has been shown using this concept with on/off ratios exceeding 2000, where both the on/off ratio and the release rate are affected by the concentration of ibuprofen. Generally, decreasing the base temperature of the release experiments from 37 to 25 ◦C significantly increases the on/off ratio. The effect of the orientation of the cylindrical iron rod with respect to the direction of the magnetic field on the heating of the device has been investigated using a Comsol model. Even though the effect of orientation is limited at small angles, a significantly lower surface temperature has been shown for larger angles, up to 20 ◦C. This can result in a several orders of ∼ magnitude difference for the diffusion coefficient of ibuprofen in the polymer. Subsequently, the requirement of alignment between the sample and the magnetic field has been circumvented by the use of a macroscopic spherical iron core, coated with ibuprofen incorporated p(S-BMA). The absence of an alignment effect has been shown using 1 sample and 2 samples in line with the magnetic field, as this does not influence the release rate and on/off ratio, normalized to the surface area available for release in on-demand release experiments. Therefore, it is possible to use multiple samples to increase the attainable drug dose. Increasing the size of the spherical iron core and, therefore, decreasing the polymer thickness, only increases the release rate upon exposure, resulting in higher on/off ratios. In the case of a thinner polymer layer, the distance between the heating core and the outer surface of the polymer is smaller, resulting in a higher temperature of the outer layer. The solubility of a solute in a polymer is predominantly important for the release characteristics of that solute from the polymer. Therefore, the solubility of ibuprofen in p(S-BMA) has been investigated. Even though samples of p(S-BMA) with an ibuprofen concentration above 31 wt% show a clear phase separation, indicating maximum solubility at 31 wt% ibuprofen, measurement of the glass transition temperature of composites show that the system of p(S-BMA) with ibuprofen concentrations below 31 wt% is in a meta-stable state. In conclusion, repetitive on-demand drug release from a polymeric implant can be externally triggered using an alternating magnetic field. Due to their biocompatibility and the absence of an alignment effect, superparamagnetic iron oxide nanoparticles

iii Summary are preferable for the required heat generation. However, more optimal nanoparticles are required for high on/oﬀ ratios, as has been shown using another material for heat generation.

iv Samenvatting

Extern aangestuurde medicijnafgifte uit een implantaat kan de efficiëntie van de behandeling sterk verbeteren omdat de patiënt of arts daarmee direct de medicijntoediening aan kan passen aan de behoefte van de patiënt en het medicijn alleen vrijkomt op de gewenste locatie in het lichaam. Daardoor zal de patiënt zich beter houden aan de behandeling en zal de biologische beschikbaarheid groter zijn. Daarnaast zullen de schadelijke bijwerkingen worden verminderd doordat andere delen van het lichaam niet worden blootgesteld aan hoge concentraties van het medicijn, zoals vaak het geval is bij systemische medicijntoediening. Dit project richt zich op een medicijnafgifte systeem dat bestaat uit een temperatuurgevoelig polymeer dat van buiten het lichaam aan en uit geschakeld kan worden door een wisselend magnetisch veld. Het temperatuurgevoelige aspect berust op een grote verandering van de diffusiecoëfficiënt van een opgeloste stof in een polymeer rond de glasovergangstemperatuur (Tg) van het polymeer. Bij een temperatuur onder de Tg van het polymeer (T < Tg) is het polymeer glasachtig en is de diffusiecoëfficiënt van het medicijn in het polymeer erg laag, waardoor weinig medicijn vrijkomt. Bij het verhogen van de temperatuur tot boven de Tg van het polymeer (T > Tg) wordt het polymeer rubberachtig. Daardoor neemt de flexibiliteit van de polymeerketens en het vrije volume enorm toe, met het resultaat dat de medicijnafgiftesnelheid toeneemt. Aangezien de glasovergang een omkeerbare overgang is, neemt de medicijnafgiftesnelheid weer af als de temperatuur weer daalt. Zodoende is gepulseerde medicijntoediening mogelijk. In deze studie wordt temperatuur van het systeem verhoogd door een extern wisselend magnetisch veld. Om de temperatuur van het implantaat te verhogen met een magnetisch veld is gebruik gemaakt van superparamagnetische ijzeroxide nanodeeltjes. Deze nanodeeltjes worden als contrastmiddel in MRI gebruikt, en om de temperatuur lokaal te verhogen bij de behandeling van tumoren waarbij deze door de verhoogde temperatuur worden vernietigd. De deeltjes zijn biocompatibel en kunnen warmte genereren in een wisselend magnetisch veld door Néel en Brown relaxatie. Bij Néel relaxatie wordt Samenvatting het magnetisch moment van het deeltje gericht naar het magnetisch veld waardoor warmte wordt gegenereerd door het overschrijden van een anisotropie-barrière. Bij Brown relaxatie richt het totale deeltje zich met het magnetisch veld, waarbij warmte wordt gegenereerd door frictie tussen het deeltje en de vloeistof. Aangezien de nanodeeltjes worden gebruikt voor het opwarmen van een polymeren implantaat zijn verschillende methodes bekeken om de deeltjes in het polymeer te verdelen. Vriesdrogen van een mengsel van gesuspendeerde deeltjes en een poly(methyl methacrylaat) (p(MMA)) latex en het vervolgens compounderen resulteert in een optimale verdeling van de deeltjes. Dit komt waarschijnlijk doordat zowel de ijzeroxide deeltjes als de polymeer deeltjes worden gestabiliseerd door een surfactant en doordat de deeltjes tijdens het vriesdrogen nagenoeg niet kunnen migreren. Andere methodes voor de verwerking in het polymeer, zoals het direct injecteren van de ijzeroxide suspensie in de compounder, resulteren in het agglomereren en een slechte verdeling van de deeltjes. De verdeling heeft een duidelijk effect op de opwarming van de deeltjes. De optimale verdeling van de deeltjes resulteert in de hoogste waarde voor de specifieke absorptie snelheid, de hoeveelheid gegenereerde warmte per gram ijzeroxide, door de minste interacties tussen deeltjes onderling. Aangezien de nanodeeltjes in het polymeer niet mobiel zijn kan er geen warmte worden gegenereerd door middel van Brown relaxatie. Hierdoor kan met een directe vergelijking tussen specifieke absorptie snelheid van de deeltjes, gesuspendeerd en in het polymeer, bepaald worden wat de bijdrage van Néel en Brown relaxatie is aan het opwarmen van de deeltjes in suspensie. Doordat in beide situaties de specifieke absorptie snelheid identiek is kan worden geconcludeerd dat ook in suspensie Néel relaxatie het enige proces is dat bijdraagt aan de opwarming van de deeltjes bij de frequentie (745 kHz) die in dit project is gebruikt. Dit is te verklaren door de veel kortere relaxatietijd van Néel relaxatie. Gebruikmakend van een cilindrische kern van p(MMA) met ijzeroxide deeltjes, gecoat met een temperatuurgevoelig polymeer, poly(butyl methacrylaat-stat-methyl methacrylaat) (p(BMA-MMA)) is extern aangestuurde medicijnafgifte onderzocht. Een modelstof voor de afgifte, ibuprofen, is verdeeld in de temperatuurgevoelige p(BMA-MMA) coating. Tijdens het blootstellen van het implantaat aan het wisselend magnetisch veld (aan situatie) is de medicijnafgiftesnelheid significant hoger dan zonder het magnetisch veld (uit situatie). Na het verwijderen van het magnetisch veld daalt de medicijnafgiftesnelheid weer naar de snelheid voor het blootstellen aan het magnetisch veld, wat de reversibiliteit van het systeem laat zien. Meerdere achtereenvolgende blootstellingen aan het magnetisch veld resulteren

vi Samenvatting steeds in een gelijke toename van de medicijnafgiftesnelheid. Verhoging van de concentratie van ijzeroxide in de kern van het systeem verhoogt de afgiftesnelheid in het veld, terwijl deze in de uit situatie niet wordt be¨ınvloed. Hierdoor wordt de aan/uit verhouding verhoogd. Ondanks dat extern aangestuurde gepulseerde medicijnafgifte mogelijk is is de maximaal gehaalde aan/uit verhouding slechts 16.5. Dit komt hoofdzakelijk door het gebruik van niet optimale commerciële ijzeroxide deeltjes en de relatief hoge afgiftesnelheid van ibuprofen uit p(BMA-MMA) in de uit situatie. Om de aan/uit verhouding te verhogen is gebruik gemaakt van een cilindrische ijzeren kern, gecoat met het temperatuurgevoelige poly(styreen-stat-butyl methacrylaat) (p(S-BMA)) waarin ibuprofen is verdeeld. De warmte wordt in de ijzeren kern gegenereerd door het magnetisch veld door middel van inductie. Ook gebruikmakend van dit systeem is extern geschakelde medicijnafgifte mogelijk, waar de aan/uit verhoudingen oplopen tot boven 2000. Zowel de afgiftesnelheid als de aan/uit verhouding zijn afhankelijk van de concentratie ibuprofen in de p(S-BMA) laag en in het algemeen wordt de aan/uit verhouding verhoogd indien de basistemperatuur wordt verlaagd van 37 naar 25 ◦C doordat de afgiftesnelheid in de uit situatie wordt verlaagd. Bovendien zijn berekeningen uitgevoerd aan het effect van de hoek tussen de kern en de richting van het magnetisch veld op de opwarming van het systeem. Hoewel dit effect relatief klein is bij een kleine hoek tussen de kern en het veld, kan bij grote hoeken de temperatuur aan het oppervlak aanzienlijk lager liggen, tot 20 ◦C. ∼ Dit kan resulteren in een verschil van een aantal ordergrootten in diffusiecoëfficiënt en dus medicijnafgifte. Vervolgens is gebruik gemaakt van macroscopische ijzeren bolletjes, gecoat met p(S- BMA) met verdeelde ibuprofen. Door de ronde geometrie van deze samples kan de noodzaak voor een bepaalde orientatie met het veld teniet worden gedaan. Dit is aangetoond doordat de aan/uit verhouding en afgiftesnelheid, beide genormeerd naar het oppervlak beschikbaar voor afgifte, van 1 implantaat en 2 implantaten op een rij in de richting van het magnetisch veld gelijk zijn. Daardoor is het mogelijk om de maximaal haalbare dosering te verhogen door gebruik te maken van meerdere samples. Het vergroten van de bolvormigen ijzere kernen, en daardoor het verkleinen van de dikte van de polymeerlaag, resulteert in een hogere aan/uit verhouding. Door de dunnere polymeerlaag is de afstand tussen de opgewarmde kern en het buitenoppervlak van het polymeer kleiner, waardoor de temperatuur op dit oppervlak hoger wordt. De oplosbaarheid van een stof in een polymeer is uitermate belangrijk voor

vii Samenvatting de afgifte kenmerken van de opgeloste stof uit het polymeer. Om die reden is de oplosbaarheid van ibuprofen in p(S-BMA) onderzocht. Hoewel bij ibuprofen concentraties van meer dan 31 gew% een duidelijke fasescheiding optreedt, wat wijst op een maximale oplosbaarheid van 31 gew% ibuprofen, laten metingen van de glasovergangstemperatuur zien dat het systeem van p(S-BMA) met een ibuprofen concentratie onder 31 gew% zich in een meta-stabiele toestand bevindt. Samenvattend, gepulseerde medicijnafgifte uit een polymeren implantaat kan extern aangestuurd worden door een wisselend magnetisch veld. Vanwege hun biocompat- ibiliteit en het ontbreken van een uitlijningseﬀect, hebben superparamagnetische ijzeroxide nanodeeltjes de voorkeur om gebruikt te worden voor de benodigde warmte generatie. Er zijn echter geoptimaliseerde nanodeeltjes nodig om hoge aan/uit verhoudingen te behalen, zoals geobserveerd bij het gebruik van andere materialen voor de warmte generatie.

viii Table of Contents

Summary i

Samenvatting v

1 Introduction 1 1.1 Introduction to Controlled Drug Delivery ...... 1 1.1.1 Conventional & Sustained Drug Delivery ...... 1 1.1.2 PreprogrammedDrugDelivery ...... 3 1.1.3 Self Responsive Drug Delivery ...... 3 1.1.4 Externally Controlled Drug Delivery ...... 3 1.1.5 Concept for Magnetically Induced Repetitive Drug Release using the Glass Transition as a Thermoresponsive Switch . . . 4 1.2 MagnetisminMedicine...... 7 1.2.1 MagneticImaging&Spectroscopy...... 7 1.2.2 Hyperthermia ...... 11 1.2.3 MagneticSeparation ...... 14 1.2.4 MagneticDrugTargeting...... 16 1.2.5 MagneticDrugRelease ...... 18 1.2.6 Conclusion...... 21 1.3 ThesisOutline...... 21 References...... 23

2 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles 35 2.1 Introduction...... 36 2.2 Materials&Methods ...... 36 2.2.1 Materials ...... 36 2.2.2 MagneticField ...... 36 2.2.3 CharacterizationoftheParticles...... 38 Table of Contents

2.2.4 TemperatureMeasurements ...... 39 2.3 Results&Discussion ...... 39 2.3.1 Characterization of the Magnetic Field ...... 39 2.3.2 CharacterizationoftheParticles...... 40 2.3.3 Inﬂuence of Field Strength on the Speciﬁc Absorption Rate. . 45 2.4 Conclusion...... 47 References...... 49

3 Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate) 51 3.1 Introduction...... 52 3.2 Materials&Methods ...... 53 3.2.1 Materials ...... 53 3.2.2 Distribution of EMG705 Particles in Polymer Matrix . . ... 53 3.2.3 Distribution of EMG1200 Particles in Polymer Matrix . .... 54 3.2.4 Characterization ...... 54 3.2.5 TemperatureMeasurements ...... 55 3.3 Results&Discussion ...... 55 3.3.1 Characterization of Iron Oxide Nanoparticles ...... 55 3.3.2 Characterization of EMG Nanoparticles Incorporated in p(MMA)...... 58 3.4 Conclusion...... 65 References...... 67

4 Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate) 69 4.1 Introduction...... 70 4.2 Materials&Methods ...... 71 4.2.1 Materials ...... 71 4.2.2 Distribution of EMG705 Particles in Polymer ...... 71 4.2.3 Characterization ...... 71 4.2.4 TemperatureMeasurements ...... 72 4.3 Results&Discussion ...... 73 4.3.1 Characterization ...... 73 4.3.2 TemperatureMeasurements ...... 74 4.4 Conclusion...... 76 References...... 78

x Table of Contents

5 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices 79 5.1 Introduction...... 80 5.2 Materials&Methods ...... 81 5.2.1 Materials ...... 81 5.2.2 Preparation of Iron Oxide Containing Heatable Core . . ... 82 5.2.3 Preparation of Thermoresponsive Release Coating ...... 82 5.2.4 ReleaseMeasurements ...... 82 5.3 Results&Discussion ...... 83 5.3.1 EﬀectofIronOxideConcentration ...... 83 5.3.2 EﬀectofIbuprofenConcentration ...... 85 5.3.3 Release from Coatings with High Drug Loading ...... 85 5.4 Conclusion...... 89 Appendix ...... 91 References...... 94

6 Repetitive on-demand drug release from polymeric matrices containing a cylindrical iron core 97 6.1 Introduction...... 98 6.2 Materials&Methods ...... 99 6.2.1 Materials ...... 99 6.2.2 DiffusionMeasurements ...... 100 6.2.3 PreparationofReleaseCoating ...... 101 6.2.4 On-demand Release Measurement ...... 101 6.3 Results&Discussion ...... 102 6.3.1 Diffusion of Ibuprofen in Poly(styrene-stat-butyl methacrylate) 102 6.3.2 On-demand Release from Poly(styrene-stat-butyl methacrylate)104 6.3.3 Effect of Alignment with Magnetic Field ...... 108 6.4 Conclusion...... 111 Appendix ...... 113 References...... 114

7 Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core 117 7.1 Introduction...... 118 7.2 Materials&Methods ...... 119 7.2.1 Materials ...... 119

xi Table of Contents

7.2.2 PreparationofReleaseCoating ...... 119 7.2.3 ReleaseMeasurement...... 121 7.3 Results&Discussion ...... 121 7.3.1 Magnetically Triggered Drug Release from Macroscopic Core- shellParticles ...... 122 7.3.2 Eﬀect of Alignment: 1 Sample vs 2 Samples ...... 123 7.3.3 Eﬀect of Ibuprofen Concentration & Core Size ...... 124 7.4 Conclusion...... 127 References...... 129

8 Additional aspects of magnetically induced drug delivery implants131 8.1 Synthesis of Superparamagnetic Iron Oxide Nanoparticles...... 132 8.1.1 Introduction...... 132 8.1.2 Materials&Methods ...... 132 8.1.3 Results&Discussion ...... 134 8.1.4 Conclusion...... 139 8.2 Solubility of Ibuprofen in p(S-BMA) ...... 139 8.2.1 Experimental ...... 139 8.2.2 Results&Discussion ...... 140 8.3 Perspectives of an AC Magnetic Field as External Trigger for RepetitiveOn-demandDrugRelease ...... 143 8.3.1 Temperature Control in Magnetically Triggered Thermo- responsive Drug Delivery Systems ...... 143 8.3.2 DesignCriteria ...... 153 8.3.3 In Vivo Application ...... 156 References...... 160

Dankwoord 166

List of Publications 169

About the author 171

xii Chapter 1

Introduction

1.1 Introduction to Controlled Drug Delivery

1.1.1 Conventional & Sustained Drug Delivery

In conventional drug delivery, the main component of a new medical treatment has been the development of the active substance. Depending on the active substance and medical requirements, a simple dosage form was chosen which would result in an optimum combination of eﬃcacy and patient comfort and compliance. These conventional dosage forms are often tablets and suspensions for oral administration and injections (e.g. intravenous, intramuscular and subcutaneous). Oral administration is a very patient friendly, easy, relatively inexpensive administration route, and therefore, the most used administration form. For drugs administered orally, absorption may begin in the mouth (e.g. sublingual dosage form). However, the majority of orally administered drugs are absorbed in the small intestine, and therefore, have to pass the acidic environment of the stomach. Consequently, tablets for these drugs have an enteric coating to protect the drugs from this harsh environment and therefore, increase the bioavailability. For drugs that have low bioavailability, due to low absorption, ﬁrst-pass metabolism, partial degradation or alteration of the drug before absorption, drugs are often administered by injection (e.g. bioavailability for intravenous administration is 100 %). However, this is obviously less convenient for the patient and it is preferable to develop formulations that are non-invasive. Chapter 1

Conventional dosing Sustained delivery dosing Toxic level Drug plasma level

Minimum effective level

Dose of Dose Dose Dose sustained delivery Time

Figure 1.1: Drug plasma level resulting from conventional and sustained drug release.

Even though these conventional administration techniques have the advantage that the patient or physician can directly control when the drug is administrated, multiple dosing is required to maintain a drug plasma level within the therapeutic window, see Figure 1.1. The therapeutic window is the drug plasma level between the minimum effective level and the toxic level. In particular for drugs which have short half- life, systems have been developed releasing the drug in a sustained manner, see Figure 1.1. As the drug is released over a prolonged period of time, less frequent dosing is required, resulting in an enhanced safety of the system and better patient compliance. Over the last decades, several sustained release systems have been developed of which most are based on biocompatible polymers, because of their tuneable release properties.1–5 However, the majority of sustained release systems are parenteral systems used for systemic release, e.g. for contraception or chronic pain, whereas polymeric drug delivery systems can also be used for local drug delivery.6,7 In the latter case, the targeted tissue can be exposed to significant drug levels, while the drug concentration is significantly reduced at non-targeted tissue, resulting in a decrease or elimination of unwanted side effects due to systemic release. Furthermore, a lower total dose is required to expose the targeted tissue to sufficient drugs, potentially decreasing the cost of the treatment. The majority of the sustained drug delivery systems start releasing the drug directly after administration, e.g. by subcutaneous injection (e.g. Implanon R and Lupron Depot R ), small insertion (e.g. Duros R ) or surgery (e.g. Gliadel R ), however, a non-penetrable coating can be applied which has to erode before the drug is released.

2 Introduction

1.1.2 Preprogrammed Drug Delivery

Since it may be desirable for the treatment to release after a certain lag time or in timed pulses, devices have been developed with preprogrammed release characteristics. For single drug doses, systems include carriers bursting after a predetermined lag time8,9 or erosion of a seal.10,11 Moreover, multiple pulses of drug release at preprogrammed times can be achieved by a sequence of drug loaded and empty layers.12 As the carrier erodes within a predetermined time, the drug is released in case the drug loaded layer is eroding. However, drug release is temporarily stopped when an empty layer erodes. Furthermore, microchips have been developed with many separate reservoirs coated with different molecular weight degradable polymer membranes.13,14 The different molecular weights result in different degradation times, and therefore, the reservoirs open sequentially, each at their predetermined times.

1.1.3 Self Responsive Drug Delivery

Recently, there is much interest in the development of systems that respond to internal or external stimuli. For example, the pH-dependent swelling and deswelling of hydrogels15 results in different diffusion rates of drug molecules incorporated in the gel.16,17 By incorporation of glucose oxidase in a pH-responsive polymer, the local pH is lowered as glucose is converted in glucose acid.18 Exposing such a polymer, additionally loaded with trilysyl insulin, to glucose results in an increased diffusion of insulin out of the gel.19 Another possibility to induced drug release by an internal stimulus based on competitive binding. As both glycosylated insulin and glucose bind to concanavalin A, the insulin is released from the concanavalin A in the presence of sufficient glucose.20 Furthermore, polymer carriers are able to release the drug loaded in the presence of a specific ion21–23 or under mechanical stress.24

1.1.4 Externally Controlled Drug Delivery

In certain cases, it is preferred that the drug is released upon external stimuli, as the exact time and dosing can be adjusted to match the patient’s needs. Externally triggered pulse-wise release has been reported using ultrasound,25–28 electrical,29–32 light33–36 and magnetic triggers.37–43 Release upon exposure to the external trigger can be based on destructive, irreversible changes or on a reversible transition of the polymer carrier. Destructive changes include degradation, carrier rupture and dissolution. For example, drug release by polymer degradation can be induced by

3 Chapter 1 ultrasound25 or the polymer carrier can be ruptured by increasing the internal pressure by exposing incorporated azobisisobutyronitrile to light.33 The polymer coating, applied as a membrane covering drug reservoirs on a microchip or as a film incorporating the drug, can be dissolved using electrical currents.29,30,32 However, repetitive on-demand drug release induced by an external trigger would be beneficial, e.g. for pain control or treatment of infections. Repetitive on- demand release from mesoporous composites has been shown using cavitation in ultrasound.28 However, reversible temperature sensitive transitions in polymers are most often used to switch the drug release on and off. Using the lower critical solution temperature (LCST) phase transition, the drug can be released or retained, depending on the application. Polymers showing an LCST phase transition are strongly swollen at a temperature below the LCST, whereas entropy driven collapse occurs due to hydrophobic interactions above the LCST.44 Therefore, drug release can be increased by increasing the temperature from below to above the LCST resulting in squeezing out the incorporated drug40,45–49 or can be increased by decreasing temperature when the drug is entrapped in the collapsed state and released in the swollen state.50,51. In the later case, the drug diffuses out of the system when swollen. Furthermore, a thermoresponsive barrier can be created using the LCST by grafting such a polymer on drug transporting membranes.52 Even though the LCST is a reversible transition, the majority of the incorporated drug is often squeezed out during one exposure to the temperature increasing trigger.

1.1.5 Concept for Magnetically Induced Repetitive Drug Release using the Glass Transition as a Thermo- responsive Switch

In addition to the lower critical solution temperature, the glass transition temperature, Tg, can be used as a reversible thermoresponsive switch. Below the glass transition temperature, the polymer is in a glassy state, where polymer chain movement and hole free volume are low, see Figure 1.2. Therefore, the diffusion coefficient of a solute, incorporated in the polymer matrix, is low. Increasing the temperature above the Tg changes the polymer from the glassy to the rubbery state. In the rubbery state, polymer chain movement and hole free volume are significantly higher than in the glassy state. Therefore, the diffusion coefficient of the incorporated solute is several orders of magnitude higher.53 Consequently, the solute is released from the matrix. Since the glass transition temperature is

4 Introduction

V Hole free volume g

Specific Volume V c Interstitial volume

Core volume

T g Temperature

Slow diffusion Fast diffusion Figure 1.2: Specific volume of an amorphous polymer around the glass transition temperature and the subsequent reversible change in diffusion coefficient of an incorporated solute.

a reversible transition, subsequent decrease of the temperature to below the Tg decreases the solute release rate. Incorporation of drugs in the polymer matrix results in a reversible thermoresponsive drug delivery device. The incorporated drug is not squeezed from the polymer matrix, i.e. emptying the matrix, and, consequently increasing the temperature of the system repetitively results in multiple doses of released drug. In order to externally induce repetitive on-demand drug release using the glass transition temperature as a thermoresponsive switch, the external trigger has to increase the temperature of the device remotely. Ultrasound and near-infrared radiation have been investigated previously as external triggers.54,55 The advantage of ultrasound is that the absorption coeﬃcients of polymers are larger than that of tissue, e.g. fat, liver and soft tissue. Therefore, no additives are required to selectively increase the temperature of the polymer. However, the attenuation of muscle, bone, and lungs is relatively high, reducing the applicability of ultrasound induced triggering to depths ranging from subcutaneous to a few centimeters. Use

5 Chapter 1

ON Magnetic field OFF

Time

Toxic level Drug level

Minimum effective level

Time

Figure 1.3: The concept of on-demand local drug delivery using a thermoresponsive switch induced by an alternating magnetic field. of near-infrared radiation to trigger the thermoresponsive drug delivery device, a wavelength (λ = 808 nm) has been chosen in the therapeutic window, where the absorption of water and body tissue is relatively low. Unfortunately, the polymer does not absorb significant amounts of radiation at this wavelength. Therefore, a near-infrared light absorbing dye is added to the polymer, which can be easily applied within the core of the matrix or as a coating. Similar to ultrasound, the use of near-infrared as the external trigger is limited to relatively low depths in the body, i.e. short distance from the skin In the work described in this thesis, the increase of temperature and subsequent drug release are induced using an alternating magnetic field, see Figure 1.3. A temperature increase can be induced by an alternating magnetic field via different mechanisms, including induction heating56–58, i.e. eddy currents and hysteresis (for ferromagnetic materials), Néel relaxation and Brown relaxation.59–61 Néel relaxation is the reorientation of the magnetic moment within the particle and Brown relaxation the reorientation of the magnetic particle itself. Currently, Néel and Brown relaxation

6 Introduction are for their medical use in hyperthermia (see Section 1.2.2). Superparamagnetic iron oxide nanoparticles are used to locally increase the temperature of tumors, thereby destroying them or increasing the efficiency of chemotherapy.62–64 These nanoparticles have no remanent magnetization after the applied field is removed and are biocompatible59. Therefore, in this project, the principles of magnetic heating developed for hyperthermia are investigated for the use of on-demand drug delivery using an alternating magnetic field. Consequently, a similar size range of iron oxide nanoparticles and magnetic field characteristics are used in this project.

1.2 Magnetism in Medicine

During the last decades, the application of magnetism in medical diagnostics has significantly increased. While techniques like magnetic resonance imaging, functional magnetic resonance imaging and magnetic separation are nowadays commercially available and widely used in practice, new techniques are still under investigation. Recently, magnetic drug targeting and hyperthermia, the treatment of malignant tumors by temperature increase, have proven to be effective in clinical trials and commercial equipment is being developed. Furthermore, on-demand magnetic drug release is being investigated, where a too high off release rate, i.e. release without the magnetic field, and fast depletion of the system are still the main challenges. A connection between humans and magnetisms has been established for a long time. Thales of Miletus, the first Greek speculative scientist and astronomer believed that as the human soul somehow produced motion, a magnet must also possess a soul, as it also produced motion.65 Since that time, the connection between man and magnetism has been explored in a wide variety of techniques, from the early removal of iron particles from the eye to more recent techniques such a magnetic imaging, drug targeting and hyperthermia.

1.2.1 Magnetic Imaging & Spectroscopy

The most commonly known use of magnetism in medical diagnosis is in magnetic resonance imaging (MRI). However, the principle of this technique, magnetic resonance, can also be used as spectroscopy. The start of the technique was initialized by the discoveries reported independently by Bloch66–68 and Purcell69 in 1946 for which they received the Nobel Prize for Physics in 1952. The principle of magnetic resonance relies on alignment of the nuclear spins using a strong external magnetic ﬁeld and perturbing this alignment using an electromagnetic ﬁeld. If an

7 Chapter 1 electromagnetic field with the same frequency as the Larmor frequency is applied in the transverse direction, the longitudinal nuclear magnetization decreases and a transverse magnetization appears and generates a magnetic resonance signal in a receiver coil. The signal rapidly fades due to two independent processes, longitudinal relaxation (T1 relaxation) and transverse relaxation (T2 relaxation). T1 relaxation is the realignment of the magnetization parallel to the main magnetic field, see Figure 1.4a, and T2 relaxation the signal decrease caused by the loss of phase coherence of the spins, see Figure 1.4b.70,71 As the Larmor frequency depends on the local magnetic field, gradient fields are used to localize the signal of the spins in the sample.72 MRI finds applications in musculoskeletal, oncological, neurological and cardiovascular imaging.

z z

x x

y y

(a)

z z

x x Mxy Mxy

y y

(b)

Figure 1.4: Principle of (a) longitudinal (T1) relaxation and (b) transverse (T2) relaxation in magnetic resonance.

Functional Magnetic Resonance Imaging

Functional magnetic resonance imaging (fMRI) is a type of MRI for studying the blood regulation of the brain with a spatial resolution of 1 mm.73 Most ∼ commonly, fMRI makes use of the blood oxygenated level dependent (BOLD) eﬀect. As (paramagnetic) deoxygenated hemoglobin in the blood enhances relaxation of the

8 Introduction

MR signal, a change in balance between oxygenated and deoxygenated hemoglobin results in a change in image contrast.70,74 Increasing the brain activity causes local changes in blood flow, blood volume and blood oxygenation.75 Therefore, acquiring images at rest and when the patient is thinking or performing a given task makes it possible to relate specific functionalities, e.g. language, memory and perception, with brain regions.75–77 One common use of fMRI in medical treatment is the preservation of functional brain tissue of patients with brain tumors. By identifying the functionality of brain tissue surrounding the tumor, potentially harmful therapy can be directed away from critical areas.78 Furthermore, functional magnetic resonance imaging can be used to show the effect of drugs on the brain function dynamically.79,80

Contrast Enhancement in MRI

The contrast in MRI can be enhanced by using contrast agents. These are chemical compounds that are able to alter the signal by modification of one or several physical features of the resonance effect, e.g. the proton density, longitudinal or transversal relaxation times.72 Contrast agents can be paramagnetic and superparamagnetic. Paramagnetic contrast agents are composed of metal ions with one or more unpaired electrons, e.g. Fe2+, Mn2+ and Gd3+. In aqueous solution these ions form a dipolar magnetic interaction with the nearby water molecules. Random fluctuations in this interaction reduces the longitudinal and transverse relaxation times. Because of their undesirable distribution and high toxicities, these metal ions are used as complexes with supramolecular ligands. As the effect of the ions requires interaction with water molecules, at least one water molecule has to be able to coordinate to the supramolecular structure. Due to the relative small size of the paramagnetic complex, this type of contrast agent is able to easily move from the blood into the interstitium and is cleared by the kidneys. An important application of a paramagnetic contrast agent is the study of the permeability of the blood brain barrier.81 Another group of contrast agents are superparamagnetic compounds, based on iron oxide crystals, either magnetite (Fe3O4) or maghemite (γ-Fe2O3) in the range of 4- 10 nm.83 These crystals consist of a large amount of paramagnetic ions and, as they are ordered, the net magnetization greatly exceeds that of a typical paramagnetic moment. Early studies have shown a dramatic reduction of the transverse relaxation times in liver and spleen.84 In addition, the particles show a contrast effect in longitudinal relaxation in the vascular system,85. For example, nickel enhanced ferrite nanoparticles show a signal increase with particle concentration in T1-

9 Chapter 1

(a) (b)

Figure 1.5: Images of phantom cells with varying concentration of the enhanced ferrite nanoparticles: (a) T1-weighted image and (b) T2-weighted image. The sample concentrations from top to bottom are 0.0714, 0.0357, 0.0179, 0.0089, 0.0045, and 0.0022 µM. Reproduced with permission.82 weighted imaging due to shortinging of the T1 relaxation and a corresponding signal loss in T2-weighted imaging due to T2 shortening, see Figure 1.5. The superparamagnetic iron oxide contrast agents are too large to leak into interstitium and therefore, act as intravascular contrast agents, also known as blood pool agents. The agents are eliminated by the reticuloendothelial system and the half-time in the blood depends strongly on the size and coating of the particles.86 Particles with a short half-time (mins) are primarily used to study the liver, spleen and GI tract, whereas long-circulating particles (hours) find applications in, e.g., imaging of vascular compartments and target specific imaging.86 Target specific imaging is possible by addition of specific ligands and antibodies to the particles.87–89 Therefore, magnetic resonance imaging with contrast agents allows visualization of specific cells or even molecules.90

Magnetic Resonance Spectroscopy

Magnetic resonance spectroscopy is commonly used in chemistry to identify struc- tures and quantify amounts based on their chemical shift. Combining spectroscopy with the ability of magnetic resonance imaging to localize the particular spins, the presence and concentration of biomolecules can be studied at the desired location in the body.91–93 Studying unique biomarkers for diseases can be used for diagnosis, prognosis and follow-up of human disease, e.g. see Figure 1.6, where the presence of N-acetylaspartate within the measured voxel in the human brain can be an

10 Introduction

a b

NAA

-1 -2 -3 -4 -5 -6 -7 PPM

Figure 1.6: In vivo point resolved (single voxel) MRI spectroscopy, with (a) axial and (b) sagital views of a human brain and the outlined voxel for magnetic resonance spectroscopy and (c) 1H spectrum with a readily visible N-acetylaspartate (NAA) peak. An aberrant NAA peak can be an indicator of brain injury or disease.70 indicator of brain injury or disease. Examples of non-neoplastic diseases include multiple sclerosis, Alzheimer’s disease, Parkinson’s disease, epilepsy, schizophrenia, HIV infections and near-drowning syndrome.94 Unfortunately, the spatial resolution of magnetic resonance spectroscopy is relatively low, approximately 2 mm,70 and the measurements are is time consuming.72

1.2.2 Hyperthermia

By elevating the temperature of malignant cells to the range of 42 to 45 ◦C, the growth of the cells can be retarded, arrested or reversed.95 The effect of temperature on the size of malignant tumors was first observed more than 130 years ago when intensified fevers decreased the size of malignant tumors.96 However, normal cells

11 Chapter 1 do not exhibit the same degree of temperature sensitivity as tumor cells.97 Due to this difference in sensitivity, cancer can be treated using local or whole-body temperature increase, so-called hyperthermia. For whole-body hyperthermia, the systemic temperature can be increased using an Aquatherm radiant-heat device.98 A patient, with exception of the head, is placed in a coiled device generating heat. Even though whole-body hyperthermia has been reported to increase the efficiency of cancer treatment using chemotherapy agents,99–101 it is obviously preferable to only increase the temperature of the targeted cells. Ways of local heat generation include microwave radiation, ultrasound, perfusion therapy, interstitial laser photocoagulation and magnetism.102,103 The potential of using magnetism to locally increase the temperature was first shown in 1957 by accumulating magnetite particles in lymph nodes of dogs, dissecting regional nodes and heating them using an alternating magnetic field.62,104 Different mechanisms can be used to heat magnetic particles using an alternating magnetic field, e.g. magnetic hysteresis, Néel relaxations and Brown relaxation.102 For hysteresis, the magnetic particles need to be multidomain particles.105 When these particles, consisting of multiple domains that are magnetically oriented in different directions, are exposed to a magnetic field, the domains all align with the direction of the magnetic field. As the direction of the magnetic field is reversed, the domains realign, thereby creating and subsequently removing domain walls. The amount of energy dissipated by this process per cycle of the magnetic field, i.e. the hysteresis loss, is equal to the surface area of the magnetization loop.56,106 In contrast to these multidomain particles, single domain particles exhibit no remanent magnetization, enhancing suspension stability. These superparamagnetic nanoparticles can be heated with an alternating magnetic field by Néel and Brown relaxation. Néel relaxation is the reorientation of the magnetic moment inside the particles, during which an anisotropic energy barrier is crossed resulting in the dissipation of energy. When the particles are suspended in low viscous media, the particles are able to reorient by Brown relaxation. This reorientation results in friction between the particles and the medium, hence frictional losses occur generating heat. Due to their biocompatibility (e.g. nontoxicity, sufficient chemical stability in the bio-environment, adjustable circulation time in blood and biodegradability), superparamagnetic iron oxide nanoparticles are used in the majority of the investigations to use iron oxide for biomedical applications.59 Regardless of the excellent biocompatibility, the amount of particles for sufficient temperature increase

12 Introduction should be minimized. Therefore, the amount of dissipated heat per gram of iron -1 oxide, the specific absorption rate (SAR, [W giron oxide]), should be maximized. The heating of these particles for magnetic hyperthermia depends, in addition to the magnetic field and medium parameters, on the particle core, hydrodynamic volume and the anisotropy constant, which is dependent on the shape and the coating of the particles.102,107 Furthermore, a narrow particle size distribution significantly increases the specific absorption rate.61 Particle sets with a broad size distribution contain significant amounts of particles that ar too large or too small to be effectively heated. Superparamagnetic iron oxide nanoparticles with a narrow size distribution -1 -1 61 are reported having a large SAR around 600 W giron oxide (11 kA m , 410 kHz). Additionally, the required amount of particles can be minimized by targeting the injected particles to the tumor cells using specific antibodies on the coating of the particles.90,108 Several studies have investigated the possibility of magnetic particle hyperthermia in in vivo experiments in animals.59,62,97,108,109 In the most successful study, complete regression of solid glioma tissue was reported in 87.5 % of female F344 rats, using magnetic cationic liposomes, exposed three times to an alternating magnetic field for 30 min.110 The same group also reported antitumor specific immunity after treatment.63 Even though work is still in progress to deliver magnetic particles specifically and in sufficient amounts to the targeted tumor cell, Phase I in vivo trials on humans took place recently.64,111 In trials on 10 patients with recurrent prostate cancer, a magnetic fluid was injected transperineally into the prostates. A weak alternating magnetic field was applied six times for 60 min using a commercial alternating magnetic field applicator (MFH300F, MagForce Nanotechnologies AG, Berlin, 100 kHz, 2.5-18 kA m-1). The trials have shown feasibility and tolerance in patients as well as that sufficiently high temperatures (55 ◦C) can be achieved with such weak fields. In order to automatically stop the dissipation of heat when the required temperature is reached, it has been suggested to use magnetic particles with a Curie temperature equal or slightly higher than this required temperature. As the Curie temperature is reached, the particles loose their magnetic ordering and stop dissipating heat.112–114

By varying the composition for La1-xSrxMnO3 or Ni1-xCrx, the Curie temperature of these particles can be changed.114,115 However, these particles typically show signiﬁcantly lower speciﬁc absorption rates compared to iron oxide nanoparticles.114 Moreover, there are some concerns regarding the biocompatibility of the used materials.116,117

13 Chapter 1

1.2.3 Magnetic Separation

Magnetic separation has been investigated for several decades. It involves the separation of a wide variety of molecules and cells, including proteins, DNA, RNA, and biomarkers using a static magnetic field.118–120 In general, the various separation systems differ in two features: the size and composition of the magnetic particles and the mode of magnetic separation.121 The first magnetic beads that have been used for magnetic separation are polymer beads in the range of 0.5 to 4.5 µm with incorporated magnetic particles,122 e.g., monodisperse polystyrene particles of 1 µm with identical amounts of magnetic iron.123,124 The advantage of these particles is the large magnetic moment of the bead resulting from the large amount of magnetic material. Therefore, the beads can be easily separated using a low gradient field.124,125 Because of their size, the preferred modes of separation are depletion and negative separation, which will be discussed below.122,125 Disadvantages of these large microbeads include the interference with the viability of the cells to be separated as well as the difficulty to detach the particles due to their multiple point attachment.125 Furthermore, the particles can change the optical properties of the cells and nonspecific entrapment of the particles in aggregates can occur. Because of this, the most commonly used particles for magnetic separation are submicron colloidal particles, consisting of a single magnetic particle specifically coated for stability and targeting. These small particles, in the range of 10 to 100 nm, have a higher stability and the binding reaction to cells is significantly faster than that of large microbeads.122,125 However, due their small size, and consequently small magnetic moment, separation requires strong fields and large field gradients. Therefore, in addition to a strong permanent magnet or electromagnet, the separation uses a column or other kind of confinement, closely packed with a ferromagnetic steel wool or iron spheres.121,122 Small colloidal magnetic particles can be used for positive or negative separation or for depletion. Moreover, similar to the larger microbeads, the coating can be altered to target different cells using ligands and antibodies. Furthermore, the colloidal particles are biodegradable and mild on cells.125 Separation of the targeted cells or molecules is possible using different modes, positive and negative selection as well as depletion. In case of positive selection the desired target cells or molecules are magnetically labeled and retained using a magnet. The supernatant is often discarded, however, it can be collected as well. Positive selection is particulary well-suited for selection of rare cells, e.g.

14 Introduction hematopoietic stem cells.121 The disadvantage of this mode is the potential change of the target by the bond with the magnetic particle. In negative selection all the unwanted cells are magnetically labeled using a mix of particles with various different antibodies or ligands, which generally requires labeling of more cells or molecules. Nonetheless, the targeted cells are untouched and do not have to be detached from the magnetic particles. Depletion is a separation mode, similar to negative separation, where specific unwanted cells are removed and the product, the supernatant, contains the wanted cells and cells or molecules which are neither wanted, nor unwanted. Magnetic separation is possible in a batch, semicontinuous or continuous mode. In a batch process, the mixture, containing the labeled and unlabeled items, is injected into a chamber with a magnet to hold the labeled particles, while the mixture of unlabeled cells is removed. The labeled cells can also be directly measured while the magnet holds these cells in place.127 In semicontinuous magnetic separation, the mixture of labeled and unlabeled cells and molecules is fed to a column that is subjected to a magnetic field. The magnetically labeled cells and molecules are retained, while the unlabeled cells and molecules pass through the column. Subsequently, the magnetic field is removed and the magnetically labeled particles are rinsed from the column. Magnetic separation can also be performed continuously.

Y Y Sample inlet Flow direction Sheath F Sheath 1 drag (+y) + 2 F buoyant (+y)

F F drag (-x) magnetic (+x) gradient

direction F gravity (-y) Field/gradient

direction collection bins X X (a) (b)

Figure 1.7: Continuous magnetic separation with (a) the force diagram for a magnetic microparticle in a ﬂow and (b) schematic diagram of the ﬂow chamber. Reproduced with permission.126

15 Chapter 1

In this mode, the mixture is subjected to a magnetic field perpendicular to the direction of flow. Unlike in the case of semicontinuous separation, the flow is strong enough for the magnetically labeled particles to move, however, the labeled particles change their trajectory compared to unlabeled particles and can therefore be separated in different fractions,126 see Figure 1.7. Moreover, magnetic membrane systems can be used for continuous flow separation.128

1.2.4 Magnetic Drug Targeting

Systemic administration of drugs is often associated with potential problems such as an homogeneous distribution of pharmaceuticals throughout the body, the lack of drug-specific affinity towards a pathological site and the necessity of a large total dose of a drug to achieve high local concentration.129 More than a century ago, German scientist and Nobel Prize winner Paul Ehrlich proposed that if an agent could selectively target a disease-causing organism, then a toxin for that organism could be delivered along with the agent of selectivity. Selective targeting of an agent can for instance be achieved by the use of site specific peptides, proteins and ligands. However, active drug targeting can also occur by attraction of magnetic particles, conjugated to drugs, to a specific site using a magnetic field.130 Magnetic drug targeting by transport of magnetic nanoparticles through the vascular system and concentration of the particles at a particular point in the body with the aid of a magnetic field was first suggested in 1960.131 The principle of magnetic drug targeting is shown in Figure 1.8, where magnetic particles bound with an anti- cancer drug are injected intravascularly and concentrated in a tumor by an external magnetic field. Initially, the active species could not be directly coupled to the magnetic particles. Therefore, both components were incorporated into carrier microspheres and applied in vivo.133–136 However, the microsized particles were often enzymatically and mechanically damaged in vivo, losing their magnetic character.137 The problem of unstable carrier microspheres was solved in 1996 when single magnetic particles were covered in starch and the active species was ionically bound to the coating.138 These particles ( 100 nm) were loaded with epirubicin and used in a Phase I clinical trial ∼ on 14 patients with advanced cancers.139. In about 50 % of the patients, the magnetic particles could be successfully directed to the tumors while organ toxicity did not increase with the treatment. Several companies now produce magnetic nanoparticles for drug targeting, e.g. FerX (USA) and Chemicell (Germany). In a Phase I/II clinical trail using a magnetically targeted carrier bound to doxorubicin particles

16 Introduction

Figure 1.8: Magnetic drug targeting with intravascular administration. An anti- cancer drug (e.g. mitoxantrone) bound to magnetic particles is injected into a blood vessel (here a tumor-feeding blood vessel) of the patient and is concentrated in the target tissue (e.g. tumor) by an external magnetic ﬁeld. Reproduced with permission.132

(MTC-DOX, FerX) one out of four treated patients had a significant reduction in tumor size, while the others showed a stable tumor size during observation of 5 to 17 months.140 In order to capture sufficient particles at the required site using the magnetic field, it is important for the particles to circulate sufficiently long. The clearance of magnetic nanoparticles by the reticuloendothelial system (RES) depends on the surface chemistry, size and magnetic properties of the particles. To retard the detection sensitivity and uptake by the macrophages of the RES, and to avoid particle agglomeration, the nanoparticles are coated with a hydrophilic compound, e.g. dextran, silica, polysaccharides or poly(ethylene glycol). Thereby, the circulatory half-life of the particles is increased from minutes to hours or days. Reducing the particle size decreases detection and can result in superparamagnetic properties. In the case of superparamagnetic particles, the magnetization disappears when no magnetic field is present. Therefore, particle agglomeration, and possible embolization of capillary vessels, is avoided.141 The control of the magnetic particles at the required site requires a significant magnetic field gradient, due to high drag forces of the blood circulation. Most often an external magnetic field is used from

17 Chapter 1 a permanent magnet, e.g. neodymium-iron-boron (NdFeB)142 or samarium-cobalt (SmCo), or a conventional or superconducting electromagnet.143–145 Permanent NdFeB magnets in combination with superparamagnetic iron oxide nanoparticles can reach effective magnetic field gradients up to 15 cm deep in the body.146 Furthermore, dynamic control of magnetic fields created by electromagnets are investigated in an attempt to focus magnetic carriers to targets deep inside the body.147 Moreover, magnetic bandages can be used for prolonged targeting (days) of the magnetic particles in close proximity to the skin (1-2 cm).148 In addition to external magnetic fields, magnetizable implants have been used to attract the magnetic carriers.149–152 The magnetizable implants are able to produce local regions of large attractive forces deep within the body. These magnetic implants are placed in the vicinity of the target by using minimally invasive surgery.141 To create large attractive forces locally without the need of surgery, strongly magnetic seeds could be transdermally injected into or near the target site.153 Compared to the magnetic drug carrier particles, which consist of mostly polymer and drug, these magnetic seeds have a significant magnetic loading. Therefore, the seeds can act as a localized magnetic element to capture the magnetic drug carrier at relatively low external magnetic field strengths.

1.2.5 Magnetic Drug Release

In addition to magnetically targeting drugs to the desired site in the patient’s body, the release of drugs can be induced using a magnetic field. Several types of macroscopic capsules have been developed for externally triggered drug delivery to the gastrointestinal tract. One type uses a static magnetic field to release the drug, which e.g. activates a hydrogen gassing cell using a reed switch to burst the drug container.154 On the other hand, an alternating magnetic field can be used to trigger release from the drug container. The alternating magnetic field is used to heat a part of the container by eddy current155 or hysteresis heating,154 releasing the drug e.g. by melting a polymer coating156 or melting a wire releasing a spring that pushes the drug out of the container.157 Even though drug release can be triggered from the device in the gastrointestinal tract, the system is depleted after single use, which combined with the typical cost of approximately 1000 USD limits applicability. Repetitive on-demand drug release can significantly improve the medical treatments and has been demonstrated using an alternating magnetic field. Early work used a low frequency (several Hz) alternating magnetic field to modulate the release rate of drugs from a polymer matrix. Using a strong permanent magnet embedded into

18 Introduction a bovine serum albumin (BSA) loaded poly(ethylene-co-vinyl acetate) copolymer and exposing the sample multiple times to a strong magnetic field using a rotating table with a frequency of 5 Hz, a 5 to 10 fold increase in release rate has been demonstrated.37 The increase in release has been ascribed to the vibration of the magnet inside the polymer matrix. This motion creates alternating tension and compression shear, which results in a pump-like effect. A comparison of poly(ethylene-co-vinyl acetate) with different Young’s moduli has shown that the polymer has to be able to move sufficiently for the pump like effect, and subsequent on-demand release, to occur.158 However, repetitive enhancement of release has been observed, whereby the high flexibility of the polymer matrix results in high release rates without exposure to the field, in the order of 0.1 mg h-1, even for the large BSA protein. The same principle and polymer matrix has been used in vivo. Exposing an implant loaded with insuline for 1 hour to an alternating magnetic field has shown a nearly 30 % decrease in the blood glucose levels of diabetic rats.38 Using similar conditions, insulin has been triggered from polymer spheres (poly(ethylenimine) cross-linked alginate), with a 50 times increase in release rate induced by exposure to the magnetic field, which the authors ascribe to an increase in water penetration and swelling of the matrix.39 Surprisingly, the increase in release of insuline did not occur during the 1 hour exposure, but only after removal of the 4 Hz magnetic field.

Drug release can also be triggered using reversible temperature sensitive transitions. The most often used transition is the lower critical solution temperature (LCST) phase transition. Polymers showing an LCST phase transition are strongly swollen in water at a temperature below the LCST, whereas entropy driven collapse occurs due to hydrophobic interactions above the LCST.44 Therefore, drug release can be increased by increasing the temperature from below to above the LCST, resulting in squeezing out the incorporated drug.45–49 By incorporation of superparamagnetic iron oxide nanoparticles in a matrix or shell of an LCST polymer, most often poly(N- isopropylacrylamide) or derivatives, a magnetically triggered thermoresponsive drug delivery system can be created.40,41,47,159 Heating the superparamagnetic iron oxide nanoparticles in an alternating magnetic ﬁeld, in the same way as in magnetic hyperthermia, results in squeezing out drugs initially incorporated in the thermoresponsive LCST polymer, see Figure 1.9. Since the LCST phase transition is reversible, subsequent removal of the alternating magnetic ﬁeld results in reswelling of the LCST polymer as the temperature of the system decreases. However, even though the LCST phase transition is reversible, the majority of the incorporated drug

19 Chapter 1

Figure 1.9: Schematic representation of on-demand drug release using a drug incorporated matrix of a polymer with a lower critical solution temperature phase transition, triggered by an alternating magnetic field. Reproduced with permission.159 is often squeezed out during one heating cycle. Therefore, the subsequent heating cycles result in significantly less drug release, limiting the applicability for repetitive on-demand drug delivery.159,160 By formation of nanocomposite polymer membranes, consisting of a poly(vinyl alcohol) gel matrix with channels made of core-shell particles of iron oxide and an LCST polymer, the permeability of the membrane can be controlled externally. Since the LCST polymer prohibits and allows drug transport through the matrix, rather than squeezing out the drug, more reproducible repetitive release can be established.161 A similar blocking effect can be achieved using gelatin, which forms a triple-helix structure below 40 ◦C.43 Increasing the temperature of the polymer by heating the incorporated iron oxide nanoparticles using an alternating magnetic field results in an increase drug release due to higher chain mobility caused by melting of the triple helices.

20 Introduction

1.2.6 Conclusion

Even though a first link between humans and magnetism has been made a long time ago, the application of magnetism in the medical diagnostics has only been flourishing in the last decades. Magnetic resonance imaging (MRI) has become one of the common techniques for musculoskeletal, oncological, neurological and cardiovascular imaging, with novel and improved MRI techniques being developed. Functional MRI has proven its applicability in the identification of brain tissue functionality and may, therefore, help to preserve critical functional areas in patients with cancer tumors. Magnetic resonance spectroscopy can be used for diagnosis, prognosis and follow-up of human disease by studying the presence and concentration of biomarkers. Using superparamagnetic iron oxide nanoparticles, the contrast in MRI can be enhanced and by addition of specific ligands and antibodies, target specific imaging is possible. These superparamagnetic iron oxide particles may generate a significant amount of heat when exposed to an alternating magnetic field. Therefore, these particles are used in local hyperthermia, where tumors are treated by increasing the temperature. Furthermore, using these magnetic particles, a wide variety of biomolecules and cells can be separated with a magnetic field, for the purification of a mixture or the selection of targets. Using a similar principle, these magnetic particles can also be attached to drug molecules, which can be directed towards a target site in the human body, called magnetic drug targeting. Moreover, a magnetic field can be used to trigger on-demand drug release, by vibration of permanent magnets in a drug incorporated polymer matrix or by increasing the temperature magnetically in combination with a reversible phase transition.

1.3 Thesis Outline

In order to develop a system for repetitive on-demand local drug delivery induced by an alternating magnetic field, in the study reported in this thesis the glass transition temperature is used as a thermoresponsive switch, as described in Section 1.1.5 In Chapter 2 commercially available superparamagnetic iron oxide nanoparticles and the magnetic field setup, used throughout this thesis, are characterized and discussed. Furthermore, the influence of the magnetic field strength on the heating of the iron oxide nanoparticles are determined. In Chapter 3 several methods of distributing iron oxide nanoparticles in a poly(methyl methacrylate) matrix are

21 Chapter 1 investigated. Moreover, the influence of the distribution on the magnetic heating of the particles has been demonstrated. The contribution of Néel and Brown relaxation to the heating of the particles, both in liquid and in a polymer, will be discussed in Chapter 4. Using the superparamagnetic iron oxide nanoparticles, repetitive on-demand release of ibuprofen, using an alternating magnetic field as the external trigger, is demonstrated in Chapter 5, using poly(butyl methacrylate-stat- methyl methacrylate as the thermoresponsive polymer. Chapter 6 and 7 describe the on-demand release from a thermoresponsive polymer (poly(styrene-stat-butyl methacrylate)) coated on iron rods and bearing balls, respectively. In Chapter 6 the effect of the alignment between the rod and the direction of the magnetic field on the heating of the thermoresponsive coating is modeled, whereas in Chapter 7 the effect of the alignment of multiple balls on the release of ibuprofen is investigated. In Chapter 8 the synthesis of iron oxide nanoparticles is described. Furthermore, the solubility of ibuprofen in poly(styrene-stat-butyl methacrylate) is addressed. Finally, the perspectives of an alternating magnetic field as external trigger for repetitive on-demand drug release are discussed. As this thesis is constructed from articles, repetition of some material is unavoidable.

22 References

References

[1] Langer, R.S. Chemical Engineering Communications, 6 (1980), 1-48.

[2] Peppas, N.A., Khara, A.R. Advanced Drug Delivery Reviews, 11 (1993), 1-35.

[3] Scott, R.A., Peppas, N.A. Biomaterials, 20 (2000), 1371-1380.

[4] Wood, K.C., Boedicker, J.Q., Lynn, D.M., Hammond, P.T. Langmuir, 21 (2005), 1603-1609.

[5] McHugh, A.J. Journal of Controlled Release, 109 (2005), 211-221.

[6] Brem, H., Walter, K.A., Langer, R. European Journal of Pharmaceutics and Biopharmaceutics, 39 (1993), 2-7.

[7] Langer, R. Pharmacology & Therapeutics, 21 (1983), 35-51.

[8] Stubbe, B.G., de Smedt, S.C., Demeester, J. Pharmaceutical Research, 21 (2004), 1732-1740.

[9] Mohamad, A., Dashevsky. Drug Development and Industrial Pharmacy, 33 (2007), 113-119.

[10] McConville, J.T., Ross, A.C., Chambers, A.R., Smith, G., Florence, A.J., Stevens, H.N.E. European Journal of Pharmaceutics and Biopharmaceutics, 57 (2004), 541-549.

[11] Gazzaniga, A., Palugan, L., Foppoli, A., Sangalli, M.E. European Journal of Pharmaceutics and Biopharmaceutics, 68 (2008), 11-18.

[12] Liu, X., Pettway, G.J., McCauley, L.K., Ma, P.X. Biomaterials, 28 (2007), 4124-4131.

[13] West, J.L. Nature Materials, 2 (2003), 709-710.

[14] Richards Grayson, A.C., Choi, I.S., Tyler, B.M., Wang, P.P., Brem, H., Cima, M.J., Langer, R. Nature Materials, 2 (2003), 767-772.

[15] Siegel, R.A., Firestone, B.A. Macromolecules, 21 (1988), 3254-3259.

[16] Brannon-Peppas, L., Peppas, N.A. Journal of Controlled Release, 8 (1989), 267-274.

23 Chapter 1

[17] Annaka, M., Tanaka, T. Nature, 355 (1992), 430-432.

[18] Gibson, Q.H., Swoboda, B.E.P., Massey, V. The Journal of Biological Chemistry, 239 (1964), 3927-3934.

[19] Fischel-Ghodsian, F., Brown, L., Mathiowitz, E., Brandenburgt, D., Langer, R. Proceedings of the National Academy of Sciences of the Uniteds States of America, 85 (1988), 2403-2406.

[20] Kim, S.W., Pai, C.M., Makino, K., Seminoﬀ, L.A., Holmberg, D.L., Gleeson, J.M., Wilson, D.E., Mack, E.J. Journal of Controlled Release, 11 (1990), 193- 201.

[21] Ito, T., Sato, Y., Yamaguchi, T., Nakao, S.I. Macromolecules, 37 (2004), 3407- 3414.

[22] Ito, T., Yamaguchi, T. Langmuir, 22 (2006), 3945-3949.

[23] Ju, X.J., Liu, L., Xie, R., Niu, C.H., Chu, L.Y. Polymer, 50 (2009), 922-929.

[24] Lee, K.Y., Peters, M.C., Anderson, K.W., Mooney, D.J. Nature, 408 (2000), 998-1000.

[25] Kost, J., Leong, K., Langer, R. Proceedings of the National Academy of Sciences of the Uniteds States of America, 86 (1989), 7663-7666.

[26] Kwok, C.S., Mourad, P.D., Crum, L.A., Ratner, B.D. Journal of Biomedical Materials Research Part A, 57 (2001), 151-164.

[27] Aschkenasya, C., Kost, J. Journal of Controlled Release, 110 (2005), 58-66.

[28] Kim, H.J., Matsuda, H., Zhou, H., Honma, I. Advanced Materials, 18 (2006), 3083-3088.

[29] Kwon, I.C., Bae, Y.H., Kim, S.W. Nature, 354 (1991), 291-293.

[30] Santini Jr, J.T., Cima, M.J., Langer, R. Nature, 397 (1999), 335.

[31] Prescott, J.H., Lipka, S., Baldwin, S., Sheppard Jr, N.F., Maloney, J.M., Cop- peta1, J., Yomtov, B., Staples, M.A., Santini Jr, J.T. Nature Biotechnology, 24 (2006), 437-438.

24 References

[32] Wood, K.C., Zacharia, N.S., Schmidt, D.J., Wrightman, S.N., Andaya, B.J., Hammond, P.T. Proceedings of the National Academy of Sciences of the Uniteds States of America, 105 (2008), 2280-2285.

[33] Mathiowitz, E., Raziel, A., Cohen, M.D., Fischer, E. Journal of Applied Polymer Science, 26 (1981), 809-822.

[34] Sershen, S.R., Westcott, S.L., Halas, N.J., West, J.L. Journal of Biomedical Materials Research Part A, 51 (2000), 293-298.

[35] Sershen, S.R., Westcott, S.L., West, J.L., Halas, H.J. Applied Physics B, 73 (2001), 379-381.

[36] Bikram, M., Gobin, A.M., Whitmire, R.E., West, J.L. Journal of Controlled Release, 123 (2007), 219-227.

[37] Edelman, E.R., Kost, J., Bobeck, H., Langer, R. Journal of Biomedical Materials Research, 19 (1985), 67-83.

[38] Kost, J., Wolfrum, J., Langer, R. Journal of Biomedical Materials Research, 21 (1987), 1367-1373.

[39] Saslawski, O., Weingarten, C., Benoit, J.P., Couvreur, P. Life Sciences, 42 (1988), 1521-1528.

[40] Zhang, J.L., Srivastava, R.S., Misra, R.D.K. Langmuir, 23 (2007), 6342-6351.

[41] Liu, T.Y., Hu, S.H., Liu, D.M., Chen, S.Y., Chen, I.W. Nano Today, 4 (2009), 52-65.

[42] Babincova, M. Pharmazie, 50 (1995), 702-703.

[43] Hu, S.H., Liu, T.Y., Liu, D.M., Chen, S.Y. Macromolecules, 40 (2007), 6786- 6788.

[44] Hoﬀman, A.S., Afrassiabi, A., Dong, L.C. Journal of Controlled Release, 4 (1986), 213-222.

[45] Yoshida, R., Kaneko, Y., Sakai, K., Okano, T., Sakurai, Y., Bae, Y.H., Kim, S.W. Journal of Controlled Release, 32 (1994), 97-102.

[46] Gutowska, A., Bark, J.S., Kwon, I.C., Bae, Y.H., Cha, Y., Kim, S.W. Journal of Controlled Release, 48 (1997), 141-148.

25 Chapter 1

[47] Satarkar, N.S., Hilt, J.Z. Acta Biomaterialia, 4 (2008), 11-16.

[48] Anal, A.K. Recent Patents on Endocrine, Metabolic & Immune Drug Discovery, 1 (2007), 83-90.

[49] Lewis, G., Coughlan, D.C., Lane, M.E., Corrigan, O.I. Journal of Microen- capsulation, 23 (2006), 677-685.

[50] Okano, T., Bae, Y.H., Jacobs, H., Kim, S.W. Journal of Controlled Release, 11 (1990), 255-265.

[51] Brazel, C.S., Peppas, N.A. Journal of Controlled Release, 39 (1996), 57-64.

[52] Vertommen, M.A.M.E., Cornelissen, H.J.L., Dietz, C.H.J.T., Hoogenboom, R., Kemmere, M.F., Keurentjes, J.T.F. Journal of Membrane Science, 322 (2008), 243-248.

[53] Kydonieus, A.F., Decker, S.C., Shah, K.R. Proceedings and program of the 18th International Symposium on Controlled Release of Bioactive Materials, 18 (1991), 417-419.

[54] Bruinewoud, H., Ultrasound-Induced Drug Release from Polymer Matrices: The glass transition temperature as a thermo-responsive switch, Ph.D. thesis, Eindhoven University of Technology, 2005, ISBN: 90-386-2877-3.

[55] Vertommen, M.A.M.E., Near-infrared induced release for localize on-demand drug delivery, Ph.D. thesis, Eindhoven University of Technology, 2009, ISBN: 987-90-386-2044-2.

[56] Bozorth, R.M., Ferromagnetism, D. van Nostrand Company, Inc., Princeton, New Jersey, 1959.

[57] Davies, E.J., Simpson, P., Induction heating handbook, McGraw-Hill, London, 1979.

[58] Davies, E.J., Conduction and induction heating, Peter Peregrinus Ltd., London, 1990.

[59] Chan, D.C.F., Kirpotin, D.B., Bunn Jr., P.A., Physical chemistry and in vivo tissue heating properties of colloidal magnetic iron oxides with increased power absorption rates, in: H¨afeli, U., et al. (eds.) Scientiﬁc and Clinical Applications of Magnetic Carriers, Plenum Press, New York, 1997.

26 References

[60] N´eel, L. Annales Geophysicae, 5 (1949), 99-136.

[61] Hergt, R., Hiergeist, R., Zeisberger, M., Gl¨ockl, G., Weitschies, W., Ramirez, L.P., Hilger, I., Kaiser, W.A. Journal of Magnetism and Magnetic Materials, 280 (2004), 358-368.

[62] Moroz, P., Jones, S.K., Gray, B.N. International Journal of Hyperthermia, 18 (2002), 267-284.

[63] Yanase, M., Shinkai, M., Honda, H., Wakabayashi, T., Yoshida, J., Kobayashi, T. Japanese Journal of Cancer Research, 89 (1998), 775-782.

[64] Johannsen, M., Gneveckow, U., Taymoorian, K., Thiesen, B., Wald¨ofner, N., Scholz, R., Jung, K., Jordan, A., Wust, P., Loening, S.A. International Journal of Hyperthermia, 23 (2007), 315-323.

[65] H¨afeli, U., The history of magnetism is medicine, in: Andr¨a, W., Nowak, H. (eds.) Magnetism in medicine: a handbook, 2nd edition, Wiley, Weinheim, 2007.

[66] Bloch, F., Hansen, W.W., Packard, M. Physical Review, 69 (1946), 127.

[67] Bloch, F. Physical Review, 70 (1946), 460-474.

[68] Bloch, F., Hansen, W.W., Packard, M. Physical Review, 70 (1946), 474-485.

[69] Purcell, E.M., Torrey, H.C., Pound, R.V. Physical Review, 69 (1946), 37-38.

[70] Kherlopian, A.R., Song, T., Duan, Q., Neimark, M.A., Po, M.J., Gohagan, J.K., Laine, A.F. BMC Systems Biology, 2 (2008), 74-91.

[71] Hornak, J.P., The Basics of MRI, http://www.cis.rit.edu/htbooks/mri/, 1996- 2008.

[72] Rodr´ıguez, I., Pérez-Rial, S., González-Jimenez, J., Pérez-Sánchez, J.M., Herranz, F., Beckmann, N., Ru´ız-Cabello, J. Journal of Pharmaceutical Sciences, 97 (2008), 3637-3665.

[73] Speck, O., Schreiber, A., Janz, C., Hennig, J., Functional magnetic resonance imaging, in: Andr¨a, W., Nowak, H. (eds.) Magnetism in medicine: a handbook, 2nd edition, Wiley, Weinheim, 2007.

27 Chapter 1

[74] Ogawa, S., Lee, T.M., Nayak, A.S., Glynn, P. Magnetic Resonance in Medicine, 14 (1990), 68-78.

[75] Kwong, K.K., Belliveau, J.W., Chesler, D.A., Goldberg, I.E., Weisskoﬀ, R.M., brigitte P Poncelet, Kennedy, D.N., Hoppel, B.E., Cohen, M.S., Turner, R., Cheng, H.M., Brady, T.J., Rosen, B.R. Proceedings of the National Academy of Sciences of the Uniteds States of America, 89 (1992), 5675-5679.

[76] Bandettini, P.A., Wong, E.C., Hinks, R.S., Tikofsky, R.S., Hyde, J.S. Magnetic Resonance in Medicine, 25 (1992), 390-.

[77] Ogawa, S., Tank, D.W., Menon, R., Ellermann, J.M., Kim, S.G., Merkle, H., Ugurbil, K. Proceedings of the National Academy of Sciences of the Uniteds States of America, 89 (1992), 5951-5955.

[78] Atlas, S.W., Howard, R.S., Maldjian, J., Alsop, D., Detre, J.A., Listerud, J., D’Esposito, M., Judy, K.D., Zager, E., Stecker, M. Neurosurgery, 38 (1996), 329-338.

[79] Borsook, D., Becerra, L. Nature Review Drug Discovery, 5 (2006), 411-425.

[80] Lannetti, G.D., Zambreanu, L., Wise, R.G., Buchanan, T.J., Huggins, J.P., Smart, T.S., Vennart, W., Tracey, I. Proceedings of the National Academy of Sciences of the Uniteds States of America4, 102 (2005), 18195-18200.

[81] Floris, S., Blezer, E.L.A., Schreibelt, G., D¨opp, E., van der Pol, S.M.A., Schadee-Eestermans, I.L., Nicolay, K., Dijkstra, C.D., de Vries, H.E. Brain, 127 (2004), 616-627.

[82] Shultz, M.D., Calvin, S., Fatouros, P.P., Morrison, S.A., Carpenter, E.E. Journal of Magnetism and Magnetic Materials, 311 (2007), 464-468.

[83] Bjørnerud, A., Johansson, L. NMR in Biomedicine, 17 (2004), 465-477.

[84] Saini, S., Stark, D.D., Hahn, P.F., Wittenberg, J., Brady, T.J., Ferrucci, J.T. Radiology, 162 (1987), 211-216.

[85] Allkemper, T., Bremer, C., Matuszewski, L., Ebert, W., Reimer, P. Radiology, 223 (2002), 432-438.

[86] Weissleder, R., Bogdanov, A., Neuwelt, E.A., Papisov, M. Advanced Drug Delivery Reviews, 16 (1995), 321-334.

28 References

[87] Xu, C., Sun, S. Dalton Transactions, 29 (2009), 5583-5591.

[88] Fang, C., Zhang, M. Journal of Materials Chemistry, 19 (2009), 6258-6266.

[89] Hifumi, H., Yamaoka, S., Tanimoto, A., Akatsu, T., Shindo, Y., Honda, A., Citterio, D., Oka, K., Kuribayashi, S., Suzuki, K. Journal of Materials Chemistry, 19 (2009), 6393-6399.

[90] Amstad, E., Zurcher, S., Mashaghi, A., Wong, J.Y., Textor, M., Reimhult, E. Small, 5 (2009), 1334-1342.

[91] Arias-Mendozaa, F., Browna, T.R. Disease Markers, 19 (2004), 49-68.

[92] Karam, J.A., Mason, R.P., Koeneman, K.S., Antich, P.P., Benaim, E.A., Hsieh, J.T. Journal of Cellular Biochemistry, 90 (2003), 473-483.

[93] Bolan, P.J., Nelson, M.T., Yee, D., Garwood, M. Breast Cancer Research, 7 (2005), 149-152.

[94] Bachert, P., MR spectroscopy, in: Andr¨a, W., Nowak, H. (eds.) Magnetism in medicine: a handbook, 2nd edition, Wiley, Weinheim, 2007.

[95] O’Brien, K.T.O., Mekkaoui, A.M. Journal of Biomechanical Engineering, 115 (1993), 247-253.

[96] Busch, C.J., Einﬂuss heftiger erysipeln auf organisierte neubildungen, in: Andr¨a, C.J. (ed.) Verhandlungen des naturhistorischen vereines der preussis- chen rheinlande und westphalens, Max Cohen and Sohn, Bonn, 1866.

[97] Gordon, R.T., Hines, J.R., Gordon, D. Medical Hypotheses, 5 (1979), 83-102.

[98] Robins, H.I., Woods, J.P., Schmitt, C.L., Cohen, J.D. Cancer Letters, 79 (1994), 137-145.

[99] Robins, H.I., Rushing, D., Kutz, M., Tutsch, K.D., Tiggelaar, C.L., Paul, D., Spriggs, D., Kraemer, C., Gilles, W., Feierabend, C., Arzoomanian, R.Z., Longo, W., Alberti, D., d’Oleira, F., Qu, R.P., Wilding, G., Stewart, J.A. Journal of Clinical Oncology, 15 (1997), 158-64.

[100] Robins, H.I., Katschinski, D.M., Longo, W., Grosen, E., Wilding, G., Gillis, W., Kraemer, C., Tiggelaar, C.L., Rushing, D., Stewart, J.A., Spriggs, D., Love, R., Arzoomanian, R.Z., Feierabend, C., Alberti, D., Morgan, K., Simon, K., d’Oleire, F. Cancer chemotherapy and pharmacology, 43 (1999), 409-414.

29 Chapter 1

[101] Westermann, A.M., Grosen, E.A., Katschinski, D.M., Jäger, D., Rietbroek, R., Schink, J.C., Tiggelaar, C.L., Jäger, E., Zum Vörde sive Vörding, P., Neuman, A., Knuth, A., Dijk, J.D.P.V., Robins, G.J.W.H.I. European Journal of Cancer, 37 (2001), 1111-1117.

[102] Hergt, R., Andr¨a, W., Magnetic hyperthermia and thermoablation, in: Andr¨a, W., Nowak, H. (eds.) Magnetism in medicine: a handbook, 2nd edition, Wiley, Weinheim, 2007.

[103] Falk, M.H., Issels, R.D. International Journal of Hyperthermia, 17 (2001), 1-18.

[104] Gilchrist, R.K., Medal, R., Shorey, W.D., Hanselman, R.C., Parrott, J.C., Taylor, C.B. Annals of surgery, 146 (1957), 596-606.

[105] Crummett, W.P., Western, A.B., University Physics: Models and Applications, WCB/McGraw-Hill, New York, 1994.

[106] Chikazumi, S., Physics of ferromagnetism, 2nd edition, Clarendon Press, Oxford, 1997.

[107] Berkowitz, A.E., Lahut, J.A., Jacobs, I.S., Levinson, L.M., Forester, D.W. Physical Review Letters, 34 (1975), 594-597.

[108] Le, B., Shinkai, M., Kitade, T., Honda, H., Yoshida, J., Wakabayashi, T., Kobayashi, T. Journal of Chemical Engineering of Japan, 34 (2001), 66-72.

[109] Hilger, I., Andr¨a, W., Hergt, R., Hiergeist, R., Schubert, H., Kaiser, W.A. Electromagnetic Heating of Breast Tumors in Interventional Radiology, 218 (2001), 570-575.

[110] Yanase, M., Shinkai, M., Honda, H., Wakabayashi, T., Yoshida, J., Kobayashi, T. Japanese Journal of Cancer Research, 89 (1998), 463-469.

[111] Johannsen, M., Gneveckow, U., Thiesen, B., Taymoorian, K., Cho, C.H., Wald¨ofner, N., Scholz, R., Jordan, A., Loening, S.A., Wust, P. European Urology, 52 (2007), 1653-1662.

[112] Kuznetsov, A.A., Leontiev, V.G., Brukvin, V.A., Vorozhtsov, G.N., Kogan, B.Y., Shlyakhtin, O.A., Tsybin, A.M.Y.O.I., Kuznetsov, O.A. Journal of Magnetism and Magnetic Materials, 311 (2007), 197-203.

30 References

[113] Kuznetsov, O.A., Sorokina, O.N., Leontiev, V.G., Shlyakhtin, O.A., Kovarski, A.L., Kuznetsov, A.A. Journal of Magnetism and Magnetic Materials, 311 (2007), 204-207.

[114] Prasad, N.K., Hardel, L., Duguet, E., Bahadus, D. Journal of Magnetism and Magnetic Materials, 321 (2009), 1490-1492.

[115] Akin, Y., Obaidat, I.M., Issaa, B., Haik, Y. Crystal Research and Technology, 44 (2009), 386-390.

[116] Messer, R.L.W., Lucas, L.C. Dental Materials, 15 (1999), 1-6.

[117] Prasad, N.K., Rathinasamy, K., Panda, D., Bahadur, D. Journal of Biomedical Materials Research Part B: Applied Biomaterials, 85B (2008), 409-416.

[118] Uhlen, M. Nature, 340 (1989), 733-734.

[119] Melville, D., Paul, F., Roath, S. Nature, 255 (1975), 663-750.

[120] He, X., Chen, Y., Wang, K., Wu, P., Gong, P., Huo, H. Nanotechnology, 18 (2007), 365604/1-365604/6.

[121] Apel, M., Heinlein, U.A.O., Miltenyi, S., Schmitz, J., Campbell, J.D.M., Magnetic cell separation for research and clinical applications, in: Andr¨a, W., Nowak, H. (eds.) Magnetism in medicine: a handbook, 2nd edition, Wiley, Weinheim, 2007.

[122] Said, T.M., Agarwal, A., Zborowski, M., Grunewald, S., Glander, H.J., Paasch, U. Journal of Andrology, 29 (2008), 134-142.

[123] Ugelstad, J., S¨oderberg, L., Berge, A., Bergstr¨om, J. Nature, 303 (1983), 95-96.

[124] Kemshead, J.T., Ugelstad, J. Molecular and Cellular Biochemistry, 67 (1985), 11-18.

[125] Miltenyi, S., Muller, W., Weichel, W., Radbruch, A. Cytometry, 11 (1990), 231-238.

[126] Carr, C., Espy, M., Nath, P., Martin, S.L., Ward, M.D., Martin, J. Journal of Magnetism and Magnetic Materials, 321 (2009), 1440-1445.

[127] Tibbe, A.G.J., de Grooth, B.G., Greve, J., Dolan, G.J., Rao, C., Terstappen, L.W.M.M. Cytometry, 47 (2002), 163-172.

31 Chapter 1

[128] Earhart, C.M., Wilson, R.J., White, R.L., Pourmand, N., Wang, S.X. Journal of Magnetism and Magnetic Materials, 321 (2009), 1436-1439.

[129] Torchilin, V.P. European Journal of Pharmaceutical Sciences, 11 (2000), S81- S91.

[130] Kim, D.K., Dobson, J. Journal of Materials Chemistry, 19 (2009), 6294-6307.

[131] Freeman, M.W., Arrott, A., Watson, J.H.L. Journal of Applied Physics, 31 (1960), 404S-405S.

[132] Siemens. Spring (2007).

[133] Widder, K.J., Senyei, A.E., Scarpelli, D.G. Proceedings of the Society for Experimental Biology and Medicine, 158 (1978), 141-146.

[134] Widder, K.J., Morris, R.M., Poore, G., Howard Jr, D.P., Senyei, A.E. Proceedings of the National Academy of Sciences of the Uniteds States of America, 78 (1981), 579-581.

[135] Gupta, P.K., Hung, C.T., Rao, N.S. Journal of Pharmaceutical Sciences, 78 (1989), 290-294.

[136] Morimoto, Y., Sugibayashi, K., Okumura, M., Kato, Y. Journal of Pharma- cobio-Dynamics, 3 (1980), 264-267.

[137] Alexiou, C., Jurgons, R., Schmid, R.J., Bergemann, C., Henke, J., Erhardt, W., Huenges, E., Parak, F. Journal of Drug Targeting, 11 (2003), 139-149.

[138] L¨ubbe, A.S., Bergemann, C., Huhnt, W., Fricke, T., Riess, H., Brock, J.W., Huhn, D. Cancer Research, 56 (1996), 4694-4701.

[139] Lübbe, A.S., Bergemann, C., Riess, H., Schriever, F., Reichardt, P., Possinger, K., Matthias, M., Dörken, B., Herrinann, F., Gürtler, R., Hohenberger, P., Haas, N., Sohr, R., Sander, B., Lemke, A.J., Ohlendorf, D., Huhnt, W., Huhn, D. Cancer Research, 56 (1996), 4686-4693.

[140] Wilson, M.W., Kerlan Jr, R.K., Fidelman, N.A., Venook, A.P., LaBerge, J.M., Koda, J., Gordon, R.L. Radiology, 230 (2004), 287-293.

[141] Arruebo, M., Fern´andez-Pacheco, R., Ibarra, M.R., Santamar´ıa, J. Nano Today, 2 (2007), 22-32.

32 References

[142] Darton, N.J., Hallmark, B., James, T., Agrawal, P., Mackley, M.R., Slater, N.K.H. Journal of Magnetism and Magnetic Materials, 321 (2009), 1571-1574.

[143] Gillies, G.T., Ritter, R.C., Broaddus, W.C., Grady, M.S., Howard III, M.A., McNeil, R.G. Review of Scientiﬁc Instruments, 65 (1994), 533-562.

[144] Cha, Y., Chen, L., Askew, T., Veal, B., Hull, J. Journal of Magnetism and Magnetic Materials, 311 (2007), 312-317.

[145] Takeda, S., Mishima, F., Fujimoto, S., Izumi, Y., Nishijima, S. Journal of Magnetism and Magnetic Materials, 311 (2007), 367-371.

[146] Neuberger, T., Sch¨opf, B., Hofmann, H., Hofmann, M., von Rechenberg, B. Journal of Magnetism and Magnetic Materials, 293 (2005), 483-496.

[147] Shapiro, B. Journal of Magnetism and Magnetic Materials, 321 (2009), 1594- 1599.

[148] H¨afeli, U.O., Gilmour, K., Zhou, A., Lee, S., Hayden, M.E. Journal of Magnetism and Magnetic Materials, 311 (2007), 323-329.

[149] Chen, H., Ebner, A.D., Rosengart, A.J., Kaminski, M.D., Ritter, J.A. Journal of Magnetism and Magnetic Materials, 284 (2004), 181-194.

[150] Yellen, B.B., Forbes, Z.G., Halverson, D.S., Fridman, G., Barbee, K.A., Chorny, M., Levy, R., Friedman, G. Journal of Magnetism and Magnetic Materials, 293 (2005), 647-654.

[151] Avil´es, M.O., Chen, H., Ebner, A.D., Rosengart, A.J., Kaminski, M.D., Ritter, J.A. Journal of Magnetism and Magnetic Materials, 311 (2007), 306-311.

[152] Fernández-Pacheco, R., Marquina, C., Valdivia, J.G., Gutiérrez, M., Romero, M.S., Cornudella, R., Laborda, A., Viloria, A., Higuera, T., Gar´ıa, A., Garc´ıa de Jalón, J.A., Ibarra, M.R. Journal of Magnetism and Magnetic Materials, 311 (2007), 318-322.

[153] Avil´es, M.O., Ebner, A.D., Ritter, J.A. Journal of Magnetism and Magnetic Materials, 321 (2009), 1586-1590.

[154] Andr¨a, W., Werner, C., Remote-controlled drug delivery in the gastrointestinal tract, in: Andr¨a, W., Nowak, H. (eds.) Magnetism in medicine: a handbook, 2nd edition, Wiley, Weinheim, 2007.

33 Chapter 1

[155] Eriksen, S.P., Swintowsky, J.V., Serfass, E.J., Lin, T.H., Abrams, J. Journal of Pharmaceutical Sciences, 50 (1961), 151-156.

[156] Minakawa, S., Henmi, H. European Patent, 0,543,498 A1 (1992).

[157] Hosoya, T., Tanaka, F., Nogichi, K. United States Patent, 4,439,197.

[158] Kost, J., abd Enora Kunica, R.N., Langer, R. Journal of Biomedical Materials Research, 19 (1985), 935-940.

[159] Satarkar, N.S., Hilt, J.Z. Journal of Controlled Release, 130 (2008), 246-251.

[160] M¨uller-Schulte, D., Schmitz-Rode, T. Journal of Magnetism and Magnetic Materials, 302 (2006), 267-271.

[161] Csetneki, I., Filipcsei, G., Zr´ınyi, M. Macromolecules, 39 (2006), 1939-1942.

34 Chapter 2

Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles

Commercially available superparamagnetic iron oxide nanoparticles (SPION) of 12 nm are characterized with respect to their physical, magnetic, and heating properties in an alternating magnetic field. For this purpose, a special magnetic field setup has been developed and characterized. Particles of the water-based ferrofluid, EMG705, and particles coated for a hydrocarbon carrier have been characterized using several techniques, including transmission electron microscopy and superconducting quantum interference device magnetometry, showing the superparamagnetic behavior of the particles. The specific absorption rate of the particles has been confirmed to have square dependency on the magnetic field strength. The difference in heating of the two samples could be explained by the difference in particle size distribution.

This chapter has been published as: S.A. Rovers, L.A.M. v.d. Poel, C.H.J.T. Dietz, J. Noijen, R. Hoogenboom, M.F. Kemmere, K. Kopinga and J.T.F. Keurentjes, Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles, Journal of Physical Chemistry C, 113 (2009), 14638-14643. Chapter 2

2.1 Introduction

Superparamagnetic iron oxide nanoparticles (SPION) are of great interest in current research and have been studied for a multitude of applications, including magnetic resonance imaging,1,2 drug targeting,2–4 magnetic separation,5–7 and hyperthermia.8–10 In hyperthermia, an AC magnetic field is used to induce a temperature increase. This magnetic heating of SPION results from two relaxation processes, namely Néel and Brown relaxation.7,8,11 Néel relaxation is the reorientation of the magnetic moment within the particles in which an anisotropy barrier is crossed, thereby causing a temperature increase. Brown relaxation is the reorientation of the magnetic particle itself in a fluid, resulting in friction between the particle and the fluid. Even though extensive research is carried out on both the synthesis of particles that generate high amounts of thermal energy in an alternating magnetic field8,9 as well as the large scale production of iron oxide nanoparticles,12,13 only a limited number of iron oxide nanoparticles is commercially available. In the present work, two different types of commercially available iron oxide nanoparticles have been studied with regard to physical and heating properties in an alternating magnetic field setup. Furthermore, the experimental setup to study the heating in an alternating magnetic field is discussed and evaluated in detail.

2.2 Materials & Methods

2.2.1 Materials

The commercially available iron oxide nanoparticles investigated in this study were purchased from FerroTec, Germany, and were used without further puriﬁcation. Both water-based ferroﬂuid (EMG705) as well as particles for a hydrocarbon carrier (EMG1200) were investigated. Tetrahydrofuran (THF) (99+ %) was purchased from Sigma Aldrich.

2.2.2 Magnetic Field

Setup

In this work, a custom built setup was used to generate an alternating magnetic field with a nominal field strength of 2850 A m-1 with a maximum frequency of 745 kHz, which decreases slightly with decreasing field strength. This setup consists

36 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles of three basic elements, see Figure 2.1. Part A consisted of a gold-coated hollow copper solenoid of 33 turns around a polycarbonate tube with an internal diameter of 60 mm. Two capacitors were placed in series across the solenoid in order to create an inductor-capacitor (LC) circuit. The center of the solenoid was at earth potential for radio frequency (RF). An RF signal was generated by a solid state push-pull oscillator (Part B) containing four IRFP3600 MOSFETS, which was connected to two taps on the solenoid, located one turn above and below the center, respectively. The frequency of this oscillator was determined by the LC circuit containing the solenoid. The DC voltage of this oscillator was supplied by a Delta Electronics SM7020-D power supply (70 V, 20 A), Part C. A feed-back circuit was used to control the average current through the MOSFETS. The output power of the oscillator, and hence the magnetic field strength, was adjusted by changing the setting of the power supply. The solenoid was kept at room temperature by flowing high electrical resistance cooling fluid through the tube of the coil. The temperature of the samples, placed in the alternating magnetic field, was measured using a fluoroptic temperature probe and recorded, Part D.

(a) (b)

Figure 2.1: Magnetization setup with (a) photograph of the solenoid and (b) schematic representation with (A) the solenoid, (B) the push-pull oscillator, (C) the DC power supply, and (D) data acquisition.

37 Chapter 2

Field Characterization

The magnetic ﬁeld strength and frequency were measured using a custom built pickup coil, consisting of a single turn with a diameter of 31.5 mm, connected to a Tektronix TDS210 oscilloscope. In the lateral direction, the pickup coil was located in the center of the solenoid of the setup. In the axial direction, the position of the pickup coil has been varied in order to determine the magnetic ﬁeld strength relative to the axial center of the solenoid.

2.2.3 Characterization of the Particles

The presence of surfactants that stabilize the commercially available iron oxide nanoparticles was studied using a TA Instruments Q500 thermogravimetric analyzer ◦ -1 -1 (TGA), with a heating rate of 10 C min under N2-flow (60 mL min ). Furthermore, the core size of the iron oxide nanoparticles was determined by transmission electron microscopy (TEM). TEM samples were prepared by diluting EMG705 ferrofluid (100x w/w) or suspending EMG1200 particles in THF (0.03 wt%) and placing a single drop on a carbon-coated copper grid. In this work, a FEI Tecnai G2 Sphera TEM operating at 200 kV was used. Furthermore, agglomeration of the nanoparticles was investigated by dynamic light scattering (DLS) using a Coulter N4 Plus. The commercially available superparamagnetic iron oxide particles were characterized using a small-angle X-ray diffraction technique to determine the class of iron oxide as well as the crystallinity of the particles. Moreover, the peak broadening effect, due to a small particle size, was used to calculate the crystallite size. For this purpose, the full width at half the peak maximum of the six most significant peaks was corrected for the peak broadening due to the inconstancy of the Lorentz-polarization. Thereafter, the Scherrer’s Equation was used to calculate the crystallite size:

0.93 λ dcrys = , (2.1) B1/2 cosθ where dcrys is the crystallite size [nm], λ the used X-ray wavelength [nm], B1/2 the peak width at half-maximum [-] and θ is the Bragg’s angle. The measurements were performed on a Rigaku Geigerﬂex diﬀractometer (copper, 40 kV, 25 mA, λ = 0.154056 nm). The magnetization of the commercial iron oxide nanoparticles was determined at room temperature with a helium cooled MPMS 50 SQUID (superconducting quantum interference device) magnetometer.

38 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles

2.2.4 Temperature Measurements

Heating experiments were performed to determine the heat generated by both types of commercial iron oxide nanoparticles, EMG705 and EMG1200. EMG1200 nanoparticles were suspended in THF (27 wt%) using an overhead mixer. Subsequently, 0.5 g of ferrofluid was placed in a glass tube with an inner diameter of 6 mm and a height of 40 mm. The temperature was measured by placing a Luxtron Fluoroptic temperature probe in the ferrofluid. Thereafter, the samples were placed in the center of the custom built magnetic field setup, and the heating was measured at seven different magnetic field strengths. The amount of heat per gram of iron oxide generated, i.e., the specific absorption rate (SAR), was calculated based on the initial heating rate of the sample, the iron oxide content and the specific heat using Equation 2.2.

C dT SAR = p (2.2) f dt w ini

-1 Where SAR is the amount of heat generated per gram of iron oxide [W giron oxide], ◦ -1 -1 Cp the speciﬁc heat of the sample, 3.42 and 1.54 J C gsample for the EMG705 ferroﬂuid and EMG1200 particles suspended in THF, respectively, fw the iron oxide dT ◦ -1 weight fraction in the sample [-] and dt ini the initial temperature increase [ C s ]. 2.3 Results & Discussion

2.3.1 Characterization of the Magnetic Field

The magnetic field, generated by the magnetic field setup, has been characterized using the single turn pickup coil. To verify the time needed for the setup to create and maintain a stable alternating magnetic field for further measurements and heating experiments, the magnetic field strength and frequency have been determined as a function of time. A steady magnetic field amplitude has been found approximately 11 min after the setup is started, see Figure 2.2. As a considerable amount of power is supplied to the setup, the temperature of the setup, and in particular that of the uncooled capacitors, is significantly increased. As an effect, the capacitance and the loss factor of the capacitors change, changing the resonance frequency and the magnetic field amplitude of the setup during the initial stages of the experiment. Consequently, all heating measurements described in this chapter have been started after 20 min of starting up the setup, when steady state has been reached.

39 Chapter 2

3500 750 ]

−1 740

3000 730

720 2500 710 Field strength Frequency

700 Frequency [kHz] 2000

690 Magnetic field strength [A m

1500 680 0 200 400 600 800 1000 1200 1400 Time [s]

Figure 2.2: Magnetic ﬁeld strength and frequency of the setup in time.

Furthermore, the magnetic field strength was determined in the axial direction relative to the vertical center of the solenoid. As expected, the maximum field strength has been found at the center of the solenoid, see Figure 2.3. The field strength as a function of axial position relative to the center can be calculated by Equation 2.3, derived from the Biot-Savart law.14

H √l2 +4r2 (l 2x) (2x + l) H = max − + (2.3) 2l " (l 2x)2 +4r2 (2x + l)2 +4r2 # −

p -1 p Where Hmax is the maximum field strength [A m ], l is the length the solenoid (0.19 m), r the radius of the solenoid (0.038 m) and x is the position from the center [m]. The experimentally found magnetic field strength corresponds well with the calculations based on the Boit-Savart law, see Figure 2.3. To maintain the field strength during measurements within 95 and 98 % of the maximum field strength, the sample has to remain within a 8.9 and 6.1 cm window, respectively. During the measurements described in this chapter, all samples were within 1 cm of the center of the solenoid and, therefore, the magnetic field strength is considered to be constant.

2.3.2 Characterization of the Particles

Thermogravimetric analysis (TGA) has been performed to investigate the presence of stabilizing surfactants on the iron oxide nanoparticles. The TGA results show

40 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles

3500 ] −1

3000

2500

2000 Measured field strength Calculated field strength

Magnetic field strength [A m 95 % of maximum field strength 98 % of maximum field strength 1500 −6 −4 −2 0 2 4 6 8 Axial position relative to center [cm]

Figure 2.3: Magnetic field strength of the setup at different axial positions relative to the center, with the calculated field strength from the Biot-Savart law. that the EMG705 particles appear to have a single surfactant that is lost in between approximately 120 ◦C and 220 ◦C, see Figure 2.4. Due to the hydrophilic nature of the iron oxide, the surfactant is presumably oriented in a double layer and contains either an amine, alcohol, or acid group.15–18 In contrast, TGA analysis of the particles for organic media, EMG1200, shows a secondary plateau from approximately 270 ◦C and 350 ◦C, indicating the presence of a second surfactant. The surfactant content of EMG705 and EMG1200 nanoparticles is 21 and 16 wt%, respectively. The

100

90 EMG705 nanoparticles EMG1200 nanoparticles 85 Residual weight [%] 80

75 0 100 200 300 400 500 Temperature [°C]

Figure 2.4: Thermogravimetric analysis of EMG705 and EMG1200 nanoparticles.

41 Chapter 2

(a) (b)

Figure 2.5: Transmission electron microscopy image of (a) EMG705 and (b) EMG1200 nanoparticles. particle size distribution of the iron oxide nanoparticles has been determined using transmission electron microscopy. TEM images show nearly spherical particles, see Figure 2.5. Both types of particles, the water-based EMG705 as well as the particles for hydrocarbon carrier, EMG1200, have a similar particle size in the superparamagnetic range.19 The TEM images have been analyzed to determine the particle size distribution, from which it has been found that both types of particles have a rather broad particle size distribution, see Figure 2.6 and Table 2.1. The EMG705 ferroﬂuid has a small fraction of relatively large particles, whereas EMG1200 has a small fraction of smaller particles, see Figure 2.6. Furthermore, agglomerates have been found for both types of particles in ﬂuid using dynamic light scattering, with a size of 118 48.4 and 175 67.3 nm, respectively. ± ±

Table 2.1: Particle and crystallite size of commercial iron oxide nanoparticles.

TEM XRD crystals [nm] [nm] EMG705 12.1 3.0 13.4 ± EMG1200 11.4 2.6 9.2 ±

X-ray diﬀraction has been used to determine the type of iron oxide of the commercial nanoparticles. The diﬀraction spectra of both samples correspond to magnetite,

Fe3O4, spectra according reference spectra of the International Centre for Diﬀraction Data, see Figure 2.7. From the X-ray diﬀraction spectra, the crystallite size of the

42 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles

60 EMG705 EMG1200 50

Intensity [%] 20

0 0 5 10 15 20 Particle size [nm]

Figure 2.6: Particle size distribution of EMG705 and EMG1200 nanoparticles obtained by TEM analysis. iron oxide nanoparticles has been calculated using Equation 2.1.20,21 The mean area diameter corresponds, within experimental error, with the TEM data, Table 2.1. The magnetization of both types of iron oxide nanoparticles has been measured using a SQUID Magnetometer. The magnetization curves clearly shows the superparamagnetic behavior of both types of nanoparticles, see Figure 2.8. The magnetization

100 100

80 80

60 60

40 40 Intensity [%] Intensity [%]

20 20

0 0

20 30 40 50 60 70 80 20 30 40 50 60 70 80 2θ [°] 2θ [°] (a) (b)

Figure 2.7: X-ray diﬀraction spectra of (a) EMG705 and (b) EMG1200 nanoparticles with corresponding spectra of the International Centre of Diﬀraction Data.

43 Chapter 2

400 EMG705 nanoparticles 300 EMG1200 nanoparticles ]

−1 200

100

−100

−200 Magnetization [kA m −300

−400 −1000 −500 0 500 1000 Magnetic field strength [mT]

Figure 2.8: Magnetization curve of EMG705 and EMG1200 nanoparticles. can be ﬁtted using the Langevin equation:22,23

mH k T M = M coth - b , (2.4) s k T mH b where m is the magnetic moment:

πµ M d3 m = 0 b , (2.5) 6 where Ms is the saturation magnetization of the particles, kb the Boltzmann constant, µ is the permeability of free space (4π 10-7), M is the saturation 0 · b magnetization of bulk magnetite (4.2 105 A m-1)24 and d is the particle diameter [m]. · In the strong ﬁeld, the limit, Equation 2.4 combined with Equation 2.5 reduces to23,25

1 M = M - α , (2.6) s H where 6 MskbT α = 3 , (2.7) π µ0Mbd For weak magnetic ﬁelds the Langevin equation gives the initial susceptibility

3 π µ0MsMbd χi = , (2.8) 18 kbT

44 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles where ∂M χ . (2.9) i ≡ ∂H H=0 Using the magnetization curves of the nanoparticles, the initial susceptibility, the saturation magnetization and the coefficient α have been calculated to be 22.2, 366 kA m-1 and 5.5 109 A2 m-2 for the EMG705 particles and 14.4, 365 kA m-1 · and 8.2 109 A2 m-2 for the EMG1200 particles, respectively. Using the limit for · strong fields, the particle size has been calculated using Equation 2.7 to be 9.7 nm and 8.5 nm for EMG705 and EMG1200 respectively. However, calculating the particle size using the limit for weak fields, Equation 2.8, d = 13.6 nm and 11.7 nm have been found for EMG705 and EMG1200 nanoparticles, respectively. The observed difference between the values calculated using the strong and weak field limitation can be explained by the fact that the approach to saturation is more sensitive to the smaller particles that are present, whereas the main contribution to the magnetization at weak fields originates from the larger particles.25,26 These results are in good agreement with the particle size obtained from TEM and X-ray diffraction.

2.3.3 Inﬂuence of Field Strength on the Speciﬁc Absorption Rate

Samples of the commercially available iron oxide nanoparticles have been heated in an alternating magnetic field at different magnetic field strengths. The temperature increases of the water-based ferrofluid, EMG705, and the nanoparticles for hydrocarbon carriers, EMG1200, suspended in THF, have been measured using a fluoroptic temperature probe, see Figure 2.9. From the initial heating rate, the specific absorption rate has been calculated using Equation 2.2. In addition, the specific absorption rate can be calculated based on particle properties using27 2 2(πmHfτN ) SAR = 2 2 , (2.10) τN kbTV (1+(2πf) τN ) where f is the frequency of the alternating field [Hz] and τN is the Néel relaxation time. Equation 2.11 is used because Néel relaxation is the only relaxation process that contributes to the heating of the suspended particles at this frequency, see

45 Chapter 2

60 2220 A m−1 55 2440 A m−1 2650 A m−1 50 2850 A m−1 Initial heating rates 45

Temperature [°C] 30

20 0 10 20 30 40 50 60 Time [s]

(a)

60 2220 A m−1 55 2440 A m−1 2650 A m−1 50 2850 A m−1 Initial heating rates 45

Temperature [°C] 30

20 0 10 20 30 40 50 60 Time [s]

(b)

Figure 2.9: Heating of (a) EMG705 aqueous ferrofluid (25 wt% particles) and (b) EMG1200 nanoparticles in THF (27 wt%) at different field strengths, with corresponding initial heating rates.

Chapter 4. The N´eel relaxation time is given by:11

KV k T τ = τ e b (2.11) N 0 ·

-9 -3 8,27 where τ 0 = 10 s, K is the anisotropy constant of magnetite of 8 kJ m , V 3 -1 the volume of the particle core [m ], kb the Boltzmann constant [J K ] and T the temperature [K].

46 Characterization and magnetic heating of commercial superparamagnetic iron oxide nanoparticles

Using an average particle size of 12 nm and a frequency of 745 kHz,28 a square frequency dependency is expected from Equation 2.10 for (2πf)2τ 2 1. Therefore, N ≪ the specific absorption rate has been divided by the frequency squared to determine the dependency of the specific absorption rate on the magnetic field strength. For both commercial types of particles, a quadratic field dependence has been observed, see Figure 2.10, in accordance with Equation 2.10. However, the specific absorption rate of the EMG1200 particles has been found to be approximately 1.55 times lower than that of the EMG705 particles.

−11 x 10 1.6 EMG705 1.4 EMG1200 Square field dependency ]

−2 1.2

Hz 1 −1

0.8 [W g

−2 0.6

0.4 SAR f

0.2

0 0 2 4 6 8 2 2 −2 6 H [A m ] x 10

Figure 2.10: Specific absorption rate of EMG705 and EMG1200 ferrofluids, at different field strengths.

Using Equations 2.10 and 2.11 and the particle size distribution, see Figure 2.6, the specific absorption rate has been calculated to estimate the effect of the difference in particles size distribution. Even though the absolute values resulting from these calculations are approximately ten times higher than the measured values, the calculated relative ratio between the samples of 1.6 agrees with the observed ratio of 1.55. The difference in observed, and calculated values might be due to agglomeration of the particles as observed by DLS, which reduces the heating of the particles.29

2.4 Conclusion

Commercially available iron oxide nanoparticles have been characterized in detail and were investigated with respect to heating in an alternating magnetic ﬁeld. A custom built magnetic ﬁeld setup has been developed and characterized,

47 Chapter 2 demonstrating that it provides a stable alternating magnetic field for all measurements. The types of investigated nanoparticles, water-based ferrofluid, EMG705, and particles coated for a hydrocarbon carrier, EMG1200, have a similar particles size with a broad size distribution in the superparamagnetic size range. From the magnetization of the EMG705 and EMG1200 nanoparticles, described by the Langevin equation, the particle size could be calculated, which was in good agreement with the particle size determined by transmission electron microscopy and X-ray diffraction. Both samples of commercial iron oxide nanoparticles consisted of magnetite with approximately 20 wt% of surfactant to stabilize the particles in suspension. The heating of the particles in fluid, occurring by Néel relaxation, has been measured at various magnetic field strengths, showing a quadratic dependence of heating on the field strength. The difference in heating of the EMG705 and EMG1200 particles could be explained by the dependence of the specific absorption rate on the particle size. Due to the presence of a fraction of smaller particles in the EMG1200 particles and a fraction of larger particles in the EMG705 particles, the specific absorption rate of EMG705 is higher.

48 References

References

[1] Jung, C.W., Jacobs, P. Magnetic Resonance Imaging, 13 (1995), 661-674.

[2] Gupta, A.K., Gupta, M. Biomaterials, 26 (2005), 3995-4021.

[3] Zhang, J.L., Srivastava, R.S., Misra, R.D.K. Langmuir, 23 (2007), 6342-6351.

[4] Alexiou, C., Jurgons, R., Schmid, R.J., Bergemann, C., Henke, J., Erhardt, W., Huenges, E., Parak, F. Journal of Drug Targeting, 11 (2003), 139-149.

[5] Melville, D., Paul, F., Roath, S. Nature, 255 (1975), 663-750.

[6] Bahaj, A.S., James, P.A.B., Moesschler, F.D. Journal of Applied Physics, 83 (1998), 6444-6446.

[7] Hergt, R., Hiergeist, R., Zeisberger, M., Gl¨ockl, G., Weitschies, W., Ramirez, L.P., Hilger, I., Kaiser, W.A. Journal of Magnetism and Magnetic Materials, 280 (2004), 358-368.

[8] Hergt, R., Hiergeist, R., Hilger, I., Kaiser, W.A., Lapatnikov, Y., Margel, S., Richter, U. Journal of Magnetism and Magnetic Materials, 270 (2004), 345-357.

[9] Chan, D.C.F., Kirpotin, D.B., Bunn Jr., P.A., Physical chemistry and in vivo tissue heating properties of colloidal magnetic iron oxides with increased power absorption rates, in: H¨afeli, U., et al. (eds.) Scientiﬁc and Clinical Applications of Magnetic Carriers, Plenum Press, New York, 1997.

[10] Hiergeist, R., Andr¨a, W., Buske, N., Hergt, R., Hilger, I., Richter, U., Kaiser, W. Journal of Magnetism and Magnetic Materials, 201 (1999), 420-422.

[11] N´eel, L. Annales Geophysicae, 5 (1949), 99-136.

[12] Park, J., An, K., Hwang, Y., Park, J.G., Noh, H.J., Kim, J.Y., Park, J.H., Hwang, N.M., Hyeon, T. Nature Materials, 3 (2004), 891-895.

[13] Lee, Y., Lee, J., Bae, C.J., Park, J.G., Noh, H.J., Park, J.H., Hyeon, T. Advanced Functional Materials, 15 (2005), 503-509.

[14] Crummett, W.P., Western, A.B., University Physics: Models and Applications, WCB/McGraw-Hill, New York, 1994.

[15] Boal, A.K., Das, K., Gray, M., Rotello, V.M. Chemistry of Materials, 14 (2002), 2628-2636.

49 Chapter 2

[16] Cheng, F.Y., Su, C.H., Yang, Y.S., Yeh, C.S., Tsai, C.Y., Wu, C.L., Wu, M.T., Shieh, D.B. Biomaterials, 26 (2005), 729-738.

[17] Molday, R.S., Mackenzie, D. Journal of Immunological Methods, 52 (1982), 353-367.

[18] Zhang, L., He, R., Gu, H.C. Applied Surface Science, 253 (2006), 2611-2617.

[19] Dutz, S., Hergt, R., Mürbe, J., Müller, R., Zeisberger, M., Andrä, W., Töpfer, J., Bellemann, M.E. Journal of Magnetism and Magnetic Materials, 308 (2007), 305-312.

[20] Scherrer, P. Nachrichten von der Gesellschaft der Wissenschaften, 26 (1918), 98-100.

[21] Langford, J.I., Wilson, A.J.C. Journal of Applied Crystallography, 11 (1978), 102-113.

[22] Parvin, K., Ma, J., Ly, J. Journal of Applied Physics, 95 (2004), 7121-7123.

[23] Rosensweig, R.E., Ferrohydrodynamics, Cambridge University Press, Cam- bridge, 1985.

[24] Kim, T., Shima, M. Journal of Applied Physics, 101 (2007), 09M516.

[25] Krekhova, M., Lattermann, G. Journal of Materials Chemistry, 18 (2008), 2842- 2848.

[26] Chantrell, R.W., Popplewell, J., Charles, S.W. IEEE Transactions on Magnet- ics, MAG-14 (1978), 975-977.

[27] Hergt, R., Andr¨a, W., d’Ambly, C.G., Hilger, I., Kaiser, W.A., Richter, U., Schmidt, H.G. IEEE Transactions on Magnetics, 34 (1998), 3745-3754.

[28] Due to a decrease in temperature of the setup at decreased ﬁeld strength, the frequency of the setup decreases slightly with decreasing ﬁeld strength.

[29] B¨uscher, K., Helm, C.A., Gross, C., Gl¨ockl, G., Romanus, E., Weitschies, W. Langmuir, 20 (2004), 2435-2444.

50 Chapter 3

Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate)

The effect of the distribution on the heating of superparamagnetic iron oxide nanoparticles in a polymer matrix has been investigated in an alternating magnetic field. Commercially available particles have been distributed using 30 and 50 wt% loading in a poly(methyl methacrylate) (p(MMA)) matrix by different preparation methods, resulting in different distributions of the particles. Freeze drying a mixture of ferrofluid and p(MMA) latex followed by compounding is found to diminish particle aggregation leading to an optimal distribution. Subsequently, the heating of the particles in the nanocomposites has been investigated in an alternating magnetic field of 2850 A m-1 with a 745 kHz frequency. These heating experiments show significantly higher specific absorption rates (SAR) of the incorporated iron oxide particles in case of the freeze drying method due to the improved distribution of the particles when compared to direct compounding or solvent casting. Furthermore, the higher particle loading provides faster heating, although the SAR decreases due to the presence of larger aggregates.

This chapter is accepted for publication as: S.A. Rovers, C.H.J.T. Dietz, L.A.M. v.d. Poel, R. Hoogenboom, M.F. Kemmere and J.T.F. Keurentjes, Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate) in an alternating magnetic ﬁeld, Journal of Physical Chemistry C. Chapter 3

3.1 Introduction

In recent years, remote heating of a polymer in a magnetic field has been studied by a number of groups for several applications including remote shape changing1,2 and externally triggered drug delivery.3–5 In drug delivery, magnetic particles are often coated with a polymer that exhibits a lower critical solution temperature (LCST). Magnetically heating the particles to above the LCST results in drug release from the polymer coating. Even though this polymer phase transition is reversible, often the majority of the incorporated drug is squeezed out in a single dose.5 For repetitive drug release, the glass transition of a polymer can be used as a reversible switch for externally triggered on-demand drug delivery.6 Below the glass transition temperature (Tg) of the polymer (T < Tg) the diffusion of the incorporated drug is low. By magnetically heating the polymer above the Tg (T > Tg) the diffusion coefficient of the drug is increased significantly. As the drugs are not squeezed out of the polymer and the drug concentration in the polymer matrix is constant due to the presence of drug crystals that replenish the released drug, this system can be used for multiple doses. Moreover, magnetic particles have been incorporated into polymer matrices to distinguish between Néel and Brown relaxation of superparamagnetic iron oxide nanoparticles.7–9 Néel and Brown relaxation are two processes by which such nanoparticles can be heated in an alternating magnetic field. Néel relaxation is the reorientation of the magnetic moment within the particle in order to stay aligned with the changing field direction. Thereby, an anisotropy barrier is crossed, resulting in a temperature increase. The Néel relaxation time is given by:10

KV τ = τ e kbT (3.1) N 0 ·

-9 where τ0 is the exponential prefactor of 10 s, K is the anisotropy constant of -3 7,11 3 magnetite of 8 kJ m , V the volume of the particle core [m ], kb the Boltzmann constant [J K-1] and T the temperature [K]. Brown relaxation is the reorientation of the magnetic particle itself in a fluid, resulting in friction and consequently, heating of the particle.12 By dispersing iron oxide particles in a polymer matrix, Brown relaxation can be excluded and Néel relaxation is the only occurring relaxation process. By comparing the heating of the particles in a ferrofluid to that in the polymer matrix, the contribution of both processes in the ferrofluid can be determined. The decreased specific absorption rate of incorporated particles compared to particles in fluid is often accredited

52 Influence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate) to the loss of Brown relaxation.8 However, the effect of the distribution of the particles in the polymer matrix on the heating of such composites has not been discussed in the literature. In the present work, the effect of different preparation methods on the distribution of commercially available iron oxide nanoparticles into a poly(methyl methacrylate) (p(MMA)) matrix has been investigated. Subsequent studies on heating the resulting nanocomposites in an alternating magnetic field have been performed to evaluate the effect of particle distribution on the specific absorption rate (SAR). Using incorporated iron oxide nanoparticles to specifically heat a polymer above its Tg, the particle distribution should be optimum for heating. A maximum specific absorption rate of the particles would minimize the amount of required particles, minimizing potential hindering of drug diffusion. It is expected from literature13 that the agglomeration of iron oxide particles will increase interparticles interactions, decreasing the heating by Néel relaxation. Presumably, the iron oxide particles require addition chemical or physical stabilization during processing to prevent significant particles agglomeration.

3.2 Materials & Methods

3.2.1 Materials

The iron oxide nanoparticles used in this study were purchased from Ferrotec, Germany. Poly(methyl methacrylate) (p(MMA)) (MW 120,000), tetrahydrofuran (THF) (99+ %), methyl methacrylate (MMA) (99.5+ %) and lauric acid (99.5+ %) were purchased from Sigma Aldrich. Tetramethyl ammonium hydroxide (25 % in water) and potassium persulfate (99+ %) were purchased from VWR International.

3.2.2 Distribution of EMG705 Particles in Polymer Matrix

The water-based iron oxide nanoparticles, EMG705, was distributed in p(MMA) in two ways. In the first method 2.5 g of p(MMA) was added to a pre-heated custom built double cone screw compounder with a volume of 5 cm3, set at 190 ◦C and rotating at 100 rpm. Subsequently, 7.5 mL of EMG705 ferrofluid (33 wt% particles) was added drop wise. After 15 minutes, cylindrical bars of approximately 10 mm with a diameter of 3 mm were extruded from the compounder. Additionally, EMG705 nanoparticles were distributed by mixing a p(MMA) latex and the EMG705 ferrofluid. Subsequently, this aqueous mixture was freeze dried using a Labconco Freezone 4.5 in combination with a Chemstar 1402N vacuum

53 Chapter 3 pump, operated at 84 10-3 mbar. The resulting powder was compounded as · described above. The p(MMA) latex was prepared by emulsion polymerization at 80 ◦C using 50 g water, 5 g tetramethyl ammonium hydroxide, 1.8 g lauric acid, 70 g methyl methacrylate (MMA) and 1 g potassium persulfate under argon atmosphere. The resulting p(MMA) particles were measured by dynamic light scattering (DLS) and were found to have a size of 54.8 14.4 nm. ±

3.2.3 Distribution of EMG1200 Particles in Polymer Matrix

The iron oxide coated for organic media, EMG1200, was distributed in p(MMA) by solvent casting. For this purpose, EMG1200 particles were suspended in THF for 2 hours by mixing in an ultrasound bath. Subsequently, a solution of p(MMA) in THF was added in the appropriate ratio. After mixing for 1 hour, the suspension was casted into a petri dish and covered with perforated aluminum foil to slowly evaporate THF and to prevent the formation of bubbles. Subsequently, the casted ﬁlm was removed, grinded and compounded into cylindrical bars, as described above.

3.2.4 Characterization

The size of the iron oxide nanoparticles was determined by transmission electron microscopy (TEM). Images were taken of the EMG705 and EMG1200 particles. TEM samples were prepared by diluting EMG705 ferrofluid (100x w/w) and suspending EMG1200 particles in THF (0.03 wt%) and placing a single drop on a carbon coated copper grid. The presence and size of agglomerates in the EMG705 ferrofluid and EMG1200 nanoparticles suspended in THF were determined using DLS, furthermore, the agglomerates in the EMG705 ferrofluid were confirmed by cryo-TEM images. DLS was performed using a Coulter N4 Plus particle size analyzer measuring 3 times for 900 seconds at 20 ◦C at an angle of 90 ◦ and analyzed using the cumulant method. Cryo-TEM samples were prepared on carbon coated copper grids by injection into liquid ethane using a Vitrobot Mark III. In this work a FEI Tecnai G2 Sphera cryo-TEM was used, operating at 200 kV. The distribution of the iron oxide in p(MMA) was examined by different microscopy techniques; light (LM) and transmission electron microscopy (TEM). For light microscopy, 2-5 µm thick coupes were made of the composite samples using a Leica RM2165 rotary microtome. Subsequently, light microscopy analysis was done using a Zeiss Axioplan 2 imaging microscope. Moreover, the particle distribution of EMG705 containing p(MMA) composite prepared by freeze drying was studied using TEM as

54 Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate) well. A Leica RM2165 rotary microtome was used to cut coupes with an approximate thickness of 50 nm and the samples were placed on a carbon coated copper grid. The viscosity of p(MMA) was measured at a shear rate of 100 s-1 and 190 ◦C using a TA Instruments AR-G2 rheometer.

3.2.5 Temperature Measurements

Temperature measurements were performed to determine the amount of heat generated by the superparamagnetic iron oxide nanoparticles in fluid and distributed in the p(MMA), induced by an AC magnetic field. Therefore, the heating profiles of iron oxide suspended in fluid were measured by placing a Luxtron Fluoroptic temperature probe in a glass tube with an inner diameter of 6 mm and a height of 40 mm containing 0.5 g of EMG705 ferrofluid or EMG1200 in THF (30 wt%). The heating of the iron oxide imbedded p(MMA) was measured by placing 3 cylindrical samples with a length of 10 mm and a diameter of 3 mm, around the Luxtron probe. Subsequently, the samples were placed in a custom built setup generating an AC magnetic field of 2850 A m-1 with a frequency of 745 kHz. The amount of heat generated per gram of iron oxide nanoparticles was calculated from the initial heating rate of the samples, the iron oxide content and the specific heat as measured by DSC analysis, Equation 3.2.

C dT SAR = p (3.2) f dt w ini Where the speciﬁc absorption rate (SAR) is the amount of heat generated per gram -1 ◦ -1 -1 of iron oxide [W giron oxide], Cp the speciﬁc heat of the sample [J C gsample], fw dT the iron oxide weight fraction of the sample [-] and dt the initial temperature ◦ ini increase [ C s-1].

3.3 Results & Discussion

3.3.1 Characterization of Iron Oxide Nanoparticles

TEM/DLS

The particle size of the superparamagnetic iron oxide nanoparticles used has been determined by transmission electron microscopy, see Figure 3.1. Both EMG705 and EMG1200 particles have a size in the superparamagnetic region, with a diameter of 12.1 3.0 nm and 11.4 2.6 nm, respectively.14 ± ±

55 Chapter 3

(a) (b)

Figure 3.1: TEM images of (a) EMG705 nanoparticles and (b) EMG1200 particles.

Figure 3.2: Cryo-TEM image of EMG705 nanoparticles.

Dynamic light scattering analysis has shown that EMG705 nanoparticles in ferroﬂuid are clustered in agglomerates of 118 48.4 nm. Using cryo-TEM images, ± these agglomerates have been conﬁrmed, see Figure 3.2. A size distribution could not be reliably determined from cryo-TEM due to the relatively small number of particles that are visualized. In addition, by analyzing a thin slice of several hundreds of nm of solution, size-discrimination can occur. In the suspension of EMG1200 particles in THF agglomerates of 175 67.3 nm have been found by DLS. ±

Heating of EMG Nanoparticles

The heating of iron oxide nanoparticles in ﬂuid has been determined by measuring the temperature increase in an alternating magnetic ﬁeld of 2850 A m-1 with a frequency of 745 kHz.

56 Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate)

EMG1200 in THF 35 EMG705 in water Initial heating rates

30 Temperature [°C]

20 0 10 20 30 40 50 Time [s]

Figure 3.3: Heating of iron oxide nanoparticles in ﬂuids.

The initial temperature increase of the water-based EMG705 ferrofluid is significantly lower than that of the EMG1200 dispersed in THF, 0.642 and 0.866 ◦C s-1, respectively, see Figure 3.3 and Table 3.1. However, due to a substantially higher heat capacity of water compared to THF, calculation of the specific absorption rates shows that the SAR of the EMG705 particles is higher than the EMG1200 particles, -1 7.82 and 5.32 W giron oxide, respectively. The increased SAR of the EMG705 particles compared to the EMG1200 particles can be explained by the slightly larger size of the EMG705 particles. As Néel relaxation is dependent on the particle size and the optimal size at the frequency used, 17.4 nm, is higher than the particles used, the heating rate due to Néel relaxation is higher for larger particles.11

Table 3.1: Initial heating rate and speciﬁc absorption rate of commercially available iron oxide nanoparticles in ﬂuid.

Particles Initial heating rate Speciﬁc absorption rate ◦ -1 -1 [ C s ] [Wgiron oxide] EMG705 in ferroﬂuid 0.642 7.82 EMG1200 in THF 0.866 5.32

57 Chapter 3

3.3.2 Characterization of EMG Nanoparticles Incorporated in p(MMA)

Eﬀect of Preparation Method on the Distribution of 50 wt% EMG Nanoparticles in p(MMA)

As agglomeration of the iron oxide nanoparticles is expected to decrease the specific absorption rate, the distribution of the iron oxide nanoparticles in the p(MMA) matrix has been studied using several preparation methods. The organic media particles, EMG1200, have been incorporated by solvent casting together with p(MMA) and subsequent compounding. The water-based particles, EMG705, have been distributed by direct injection of the ferrofluid and p(MMA) in the compounder as well as by freeze drying a mixture of ferrofluid and p(MMA) latex and subsequent compounding. Figure 3.4 shows the light microscopy images of p(MMA) containing 50 wt% of iron oxide nanoparticles. In the case of direct injection of ferrofluid in the compounder prefilled with p(MMA), agglomerates with a large size distribution from 1 - 50 µm have been observed, see Figure 3.4a. Most likely, these agglomerates are formed during the fast evaporation of the water as the ferrofluid is injected in the compounder operating at 190 ◦C. During the mixing of the particles in the polymer, the clusters remain as the compounder is not capable of disintegrating the clusters. As an approximation, the size reduction of the agglomerates in the compounder has been estimated as a result of rupture and erosion mechanisms. For rupture, a model derived for agglomerates of small spherical particles in a newtonian fluid has been used, in which the ratio between the shear force (Fh) and the cohesive force (F c) is described by15 F h = Z sin2θsinφcosφ. (3.3) Fc · in which, 8 ǫ 24z2d Z = χµγ . (3.4) 9 1-ǫ A Where θ and φ are the Euler angles of orientation, χ is a numerical constant of 12.23,16 µ is the medium viscosity [Pa s], γ is the shear rate [s-1], ǫ is the porosity [-], assuming 0.26 for the iron oxide agglomerates, z is the physical adsorption separation distance of typically 0.4 nm,15 d is the primary particle size [m] and A is the Hamaker constant of 15.2 10-20 J for iron oxide in p(MMA),17–19. In order to rupture the · agglomerates under shear flow, the force of shear flow is required to be larger than the cohesive force, so that F /F 1. h c ≥

58 Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate)

(a) (b)

Figure 3.4: Light microscopy images of the distribution of 50 wt% EMG nanoparticles in p(MMA) by different preparation methods: (a) EMG705 by direct injection of the ferrofluid in the compounder, (b) EMG1200 by solvent casting, (c) EMG1200 by solvent casting and subsequent compounding and (d) EMG705 by freeze drying a mixture of ferrofluid and p(MMA) latex and subsequent compounding.

Assuming optimum positioning of the agglomerate for rupture, rupture will occur when Z 2. Using the viscosity of p(MMA) measured to be 5000 Pa s at 190 ◦C ≥ under the maximum shear rate experienced, 100 s-1, Z has been calculated to be 0.6. Therefore, no rupture of the agglomerates is expected due to the shear force. Moreover, the local tensile stress around agglomerates in a viscoelastic matrix has been found to be lower than in a newtonian ﬂuid, further suppressing rupture of the agglomerates.20 However, erosion is feasible at signiﬁcantly lower shear rates than rupture, although this, is a considerably slower process.21 The size reduction due to erosion can be

59 Chapter 3 described by21 ′′ R0-R(t) k = Waµγt (3.5) R0 τc ′′ where R0 and R(t) are the agglomerate radius at time 0 and t, respectively, k is a constant independent of agglomerate and medium, Wa is the work of adhesion, which is square root dependent on the surface tension of the medium and τc is the cohesivity, where 1-ǫ A τ . (3.6) c ∝ ǫ 24z2d Calculating an average shear rate of 50 s-1 for 15 minutes, the dimensionless shearing time is 45000. For agglomerates of titanium oxide in p(DMS), it has been found that the size reduction due to erosion is approximately 0.5 % at this shearing time.21 Taking into account the differences in both the agglomerates and the medium, the size reduction due to erosion in the present case has been estimated. For the agglomerates these hold the Hamaker constant, 2.7 10-20 J for titanium oxide18,19 · as well as the primary particle size and porosity, 167 nm and 0.574, respectively, for the titanium oxide reported. Differences of the medium taken into account are the viscosity, 10 Pa s for p(DMS) and surface tension of the medium, 40 and 18.4 mJ m-2 for p(MMA) and p(DMS), respectively,22. This results in an estimation of the size reduction due to erosion of 1.2 %. Therefore, no significant size reduction of the agglomerates is expected in the present case by shear flow in the compounder. This analysis clearly indicates that clustering of the nanoparticles has to be prevented during processing since formed agglomerates cannot be ruptured afterwards. When iron oxide nanoparticles are solvent casted into the polymer and subsequently compounded, large agglomerates have been found as well, see Figure 3.4c. It can be speculated that during the solvent casting process the surfactant is probably able to dissolve in the polymer matrix, thus detaching from the iron oxide surface.23 Therefore, the nanoparticles are not sufficiently stabilized during slow solvent evaporation and the agglomerates are formed during solvent casting. The high particle loading and the loose packing of the agglomerates in the sample are evident from Figure 3.4b. As in the case of injection of the ferrofluid, the compounder is not able to fragment these clusters. However, in case that the EMG705 ferrofluid and a p(MMA) latex are premixed and freeze dried prior to compounding, no agglomerates are observed by light microscopy, see Figure 3.4d. Using transmission electron microscopy, however, small sized agglomerates could be observed, as depicted in further on in Figure 3.6b The size of these clusters ( 500 nm) is slightly larger than for the clusters found in the ∼

60 Influence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate) initial ferrofluid ( 100 nm). It is thought that due to the high iron oxide loading and, ∼ thus, close proximity of particles, the particles might agglomerate to some extent. Nonetheless, since both the latex and iron oxide particles in solution are stabilized by surfactants, the agglomeration is suppressed. During subsequent freeze drying, the particle position remains unaffected, preventing the formation of agglomerates. As the dried mixture is compounded above the glass transition temperature of the latex, the latex particles flow together forming a solid matrix between the iron oxide particles.

Eﬀect of Loading on the Distribution of Iron Oxide Nanoparticles Incorporated in a p(MMA) Matrix

To decrease particle agglomeration, the distribution of 30 wt% iron oxide particles in p(MMA) has been investigated using the different preparation methods, identical to the distribution methods used for 50 wt% particles. Similar to the distribution of 50 wt%, the direct injection of ferrofluid into the compounder and solvent casting the particles into p(MMA) and subsequent compounding result in substantial agglomeration of the iron oxide particles, see Figures 3.5a and c, respectively. Nevertheless, it should be noted that the agglomerates are smaller (up to 25 µm). However, freeze drying a mixture of ∼ ferrofluid and p(MMA) latex and subsequent compounding, results in an excellent distribution of the particles. Identical to the distribution of 50 wt%, the stabilization of both the latex and iron oxide particles by different surfactants and the lack of movement during the freeze drying process, results in good distribution of the iron oxide particles. As it is not possible to identify any agglomerates using light microscopy, TEM has been used to further investigate the particle distribution. The TEM images show very small agglomerates in the order of 100 to 200 nm, see Figure 3.6a. These agglomerates have a similar size as the agglomerates that have been observed in the ferrofluid, measured by DLS. Therefore, it can be concluded that the very small agglomerates have not been formed during the incorporation into the p(MMA) matrix.

Eﬀect of Particle Distribution on Heating Iron Oxide Nanoparticles p(MMA) Nanocomposites

To investigate the eﬀect of iron oxide particle distribution in the p(MMA) matrix on the heating of the particles in the nanocomposites containing 50 wt% in an AC

61 Chapter 3

(a) (b)

(c) (d) Figure 3.5: Light microscopy images of the distribution of 30 wt% EMG nanoparticles in p(MMA) by different preparation methods: (a) EMG705 by direct injection of the ferrofluid in the compounder, (b) EMG1200 by solvent casting, (c) EMG1200 by solvent casting and subsequent compounding and (d) EMG705 by freeze drying a mixture of ferrofluid and p(MMA) latex and subsequent compounding.

(a) (b) Figure 3.6: TEM images of p(MMA) containing (a) 30 wt% and (b) 50 wt% of EMG705 nanoparticles prepared by the freeze drying method.

62 Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate)

200

160 EMG705 Freeze Drying EMG705 Direct Injection 120 EMG1200 Solvent Casting Initial heating rates

80 Temperature [°C] 40

0 0 50 100 150 200 250 300 Time [s]

Figure 3.7: Heating of iron oxide - p(MMA) composites containing 50 wt% of diﬀerent iron oxide nanoparticles and using diﬀerent preparation methods.

Table 3.2: Initial heating rate and speciﬁc absorption rate of iron oxide nanoparticles incorporated in p(MMA) (50 wt% loading) by diﬀerent preparation methods.

Preparation method Initial heating rate Specific absorption rate ◦ -1 -1 [ C s ] [Wgiron oxide] EMG705 Freeze Drying 2.051 6.53 EMG705 Direct Injection 0.590 1.88 EMG1200 Solvent Casting 0.288 0.92 magnetic field, temperature measurements have been performed. For this purpose, 3 cylindrical bars have been placed around a fluoroptic temperature probe in an alternating magnetic field. The heating of the nanocomposite prepared by freeze drying the mixture of p(MMA) latex and ferrofluid shows a significantly higher initial temperature increase than the samples prepared by the other two methods, see Figure 3.7 and Table 3.2. Due to the substantially smaller agglomerates, the Néel relaxation of the iron oxide nanoparticles experience less interparticle interactions, resulting in faster heating of the composite.13 Furthermore, the reduced heating of the composite prepared by solvent casting compared to the heating of the composite prepared by direct injection of the

63 Chapter 3 ferrofluid can be explained by the difference in iron oxide particle size as has been observed and described in the heating of the particles in fluid. Moreover, slightly increased agglomeration in the case of solvent casting prepared composite would be able to explain the reduced heating. Unfortunately, this could not be quantified.

Eﬀect of Loading on the Heating of Iron Oxide Nanoparticles Incorpo- rated in p(MMA)

The effect of loading on the heating of iron oxide nanoparticles incorporated in p(MMA) in an AC magnetic field has been determined by a direct comparison of the heating of nanocomposites prepared by the freeze drying methode containing 30 and 50 wt% of iron oxide in an alternating magnetic field.

160

140

120

100 50 wt% EMG705 30 wt% EMG705 80 Initial heating rates

Temperature [°C] 60

20 0 20 40 60 80 100 120 140 Time [s]

Figure 3.8: Heating of iron oxide - p(MMA) composites containing 30 and 50 wt% of EMG705 nanoparticles prepared by the freeze drying method.

As can be expected, the nanocomposite containing 50 wt% iron oxide particles has a higher initial heating rate than the composites containing 30 wt%, see Figure 3.8 and Table 3.3. The composite with the higher loading is capable of reaching higher temperatures as well. However, when comparing the specific absorption rates it has been found that in the composite with a loading of 30 wt% more thermal energy is produced per unit of mass of iron oxide incorporated, Table 3.3. This can be explained by the improved distribution of the iron oxide particles in the p(MMA) matrix in the case of 30 wt% loading. The specific absorption rate found with 30 wt% loading is identical to that of the initial ferrofluid, demonstrating similar

64 Inﬂuence of distribution on the heating of superparamagnetic iron oxide nanoparticles in poly(methyl methacrylate)

Table 3.3: Initial heating rate and speciﬁc absorption rate of p(MMA) containing diﬀerent loading of iron oxide nanoparticles prepared by the freeze drying method.

Iron oxide loading Initial heating rate Specific absorption rate ◦ -1 -1 weight % [ C s ] [Wgiron oxide] 30 1.421 7.83 50 2.051 6.53 particle agglomeration. Thus, the specific absorption rate due to Néel relaxation is identical to the overall specific absorption rate, justifying the conclusion that only Néel relaxation contributes to heating the ferrofluid, also see Chapter 4

3.4 Conclusion

Different preparation methods of incorporating superparamagnetic iron oxide nanoparticles into a p(MMA) matrix have been investigated. The distribution of the nanoparticles has been examined by light microscopy and transmission electron microscopy. The method of mixing a ferrofluid with a polymer latex solution and subsequent freeze drying and compounding the resulting powder, resulted in none to very little agglomeration of the iron oxide particles in the polymer matrix compared to the nanosized agglomerates in the initial ferrofluid. The absence of agglomeration is expected to be due to the combination of stabilizing both the polymer and iron oxide particles by surfactants and the lack of mobility during freeze drying. Furthermore, the effect of the distribution of the particles on the heating thereof has been determined by comparing the heating of nanocomposites with different particle distributions. It has been observed that improvement of the distribution results in a significantly better heating of the composite, i.e., the specific absorption rate of the iron oxide nanoparticles increases, due to a decrease in interparticle interactions. Moreover, it has been found that a lower loading of the particles results in a decrease in agglomeration as the probability of agglomeration decreases. Therefore, a higher specific absorption rate has been observed in composites with a lower iron oxide loading. Nonetheless, it is worth mentioning that despite the lower specific absorption rate, the heating rate and the attainable temperature increase with increasing particle loading.

65 Chapter 3

As the decrease in speciﬁc absorption rate of superparamagnetic iron oxide nanoparticles when incorporated in a polymer matrix is often accredited to the loss in Brown relaxation, this contribution clearly demonstrates that the distribution of the particles in the matrix should be taken into account as well.

66 References

References

[1] Mohr, R., Kratz, K., Weigel, T., Lucka-Gabor, M., Moneke, M., Lendlein, A. PNAS, 103 (2006), 3540-3545.

[2] Schmidt, A.M. Macromolecular Rapid Communications, 27 (2006), 1168-1172.

[3] Schmidt, A.M. Journal of Magnetism and Magnetic Materials, 289 (2005), 5-8.

[4] Wakamatsu, H., Yamamoto, K., Nakao, A., Aoyagi, T. Journal of Magnetism and Magnetic Materials, 302 (2006), 327-333.

[5] M¨uller-Schulte, D., Schmitz-Rode, T. Journal of Magnetism and Magnetic Materials, 302 (2006), 267-271.

[6] Keurentjes, J.T.F., Kemmere, M.F., Bruinewoud, H., Vertommen, M.A.M.E., Rovers, S.A., Hoogenboom, R., Stemkens, L.F.S., P´eters, F.L.A.M.A., Tielen, N.J.C., van Asseldonk, D.T.A., Gabriel, A., Joosten, B., Marcus, M.A.E. Angewandte Chemie International Edition, 48 (2009), 9867-9870.

[7] Hergt, R., Hiergeist, R., Hilger, I., Kaiser, W.A., Lapatnikov, Y., Margel, S., Richter, U. Journal of Magnetism and Magnetic Materials, 270 (2004), 345-357.

[8] Chan, D.C.F., Kirpotin, D.B., Bunn Jr., P.A., Physical chemistry and in vivo tissue heating properties of colloidal magnetic iron oxides with increased power absorption rates, in: H¨afeli, U., et al. (eds.) Scientiﬁc and Clinical Applications of Magnetic Carriers, Plenum Press, New York, 1997.

[9] Rovers, S.A., Hoogenboom, R., Kemmere, M.F., Keurentjes, J.T.F. Journal of Physical Chemistry C, 112 (2008), 15643-15646.

[10] N´eel, L. Annales Geophysicae, 5 (1949), 99-136.

[11] Hergt, R., Andr¨a, W., d’Ambly, C.G., Hilger, I., Kaiser, W.A., Richter, U., Schmidt, H.G. IEEE Transactions on Magnetics, 34 (1998), 3745-3754.

[12] Hergt, R., Hiergeist, R., Zeisberger, M., Gl¨ockl, G., Weitschies, W., Ramirez, L.P., Hilger, I., Kaiser, W.A. Journal of Magnetism and Magnetic Materials, 280 (2004), 358-368.

[13] B¨uscher, K., Helm, C.A., Gross, C., Gl¨ockl, G., Romanus, E., Weitschies, W. Langmuir, 20 (2004), 2435-2444.

67 Chapter 3

[14] Dutz, S., Hergt, R., Mürbe, J., Müller, R., Zeisberger, M., Andrä, W., Töpfer, J., Bellemann, M.E. Journal of Magnetism and Magnetic Materials, 308 (2007), 305-312.

[15] Manas-Zloczower, I., Mixing in high-intensity batch mixers, in: Rauwendaal, C. (ed.) Mixing in polymer processing, Marcel Dekker, Inc., New York, 1991.

[16] Nir, A., Acrivos, A. Journal of Fluid Mechanics, 59 (1973), 209-223.

[17] Rosensweig, R.E., Ferrohydrodynamics, Cambridge University Press, Cam- bridge, 1985.

[18] Drummond, C.J., Georgaklis, G., Chan, D.Y.C. Langmuir, 12 (1996), 2617- 2621.

[19] Visser, J. Advances in Colloid and Interface Science, 3 (1972), 331-363.

[20] Astruc, M., Vervoort, S., Nouatin, H.O., Coupez, T., Puydt, Y.D., Navard, P., Peuvrel-Disdier, E. Rheologica Acta, 42 (2003), 421-431.

[21] Lee, Y.J., Feke, D.L., Manas-Zloczower, I. Chemical Engineering Science, 48 (1993), 3363-3372.

[22] Zisman, W.A., Contact angle wettability and adhesion, in: Advances in Chemistry Series 43, American Chemical Society, Washington, D.C., 1964.

[23] Povey, A.C., Nixon, J.R., O’Neill, I.K. Journal of Microencapsulation, 4 (1987), 299-314.

68 Chapter 4

Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate)

Commercially available superparamagnetic iron oxide nanoparticles (SPION) of 12 nm are investigated with respect to the contribution of both Néel and Brown relaxation to the magnetic heating of these particles. For this purpose experiments have been performed to heat the particles, suspended in a liquid as well as incorporated in a poly(methyl methacrylate) (p(MMA)) matrix, using an alternating (AC) magnetic field of 745 kHz at 3 different field strengths up to 2850 A m-1. It has been shown that the specific absorption rates (SAR) of the particles in the ferrofluid are identical within experimental error to the SAR of the particles incorporated in p(MMA) at all measured field strengths. As Brown relaxation is not possible if the SPION are incorporated in the polymer matrix, it can be concluded that Néel relaxation is the only relaxation process that contributes to heating the ferrofluid at 745 kHz as well. This is confirmed by calculation of the relaxation times for Néel and Brown relaxation. Furthermore, these calculations suggest that Néel relaxation is also the predominant relaxation process for small superparamagnetic particles at lower frequencies.

This chapter has been published as: S.A. Rovers, C.H.J.T. Dietz, R. Hoogenboom, M.F. Kemmere and J.T.F. Keurentjes, Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate), Journal of Physical Chemistry C, 112 (2008), 15643-15646. Chapter 4

4.1 Introduction

Superparamagnetic iron oxide nanoparticles (SPION) are of great interest and have been studied for a multitude of applications, including magnetic resonance imaging,1,2 drug targeting,2–4 magnetic separation5–7 and hyperthermia.8–10 In hyperthermia, an AC magnetic field is used to induce a temperature increase. This magnetic heating of SPION results from two relaxation processes, namely Néel and Brown relaxation.8,11 Néel relaxation is the reorientation of the magnetic moment within the particles, in which an anisotropy barrier is crossed, thereby causing a temperature increase. Brown relaxation is the reorientation of the magnetic particle itself in a fluid, resulting in friction between the particle and the fluid. The Néel relaxation time is given by:11

KV τ = τ e kbT (4.1) N 0 · and the Brown relaxation time by:6

3ηVH τB = (4.2) kbT

-9 -3 8 where τ 0 = 10 s, K is the anisotropy constant of 8 kJ m , V the volume of 3 -1 the particle core [m ], kb the Boltzmann constant [J K ], T the temperature [K], -1 -1 η the viscosity of the carrier liquid [kg m s ] and VH the hydrodynamic volume of the particle [m3]. If SPION are dispersed in a highly viscous medium, e.g. a gel, the particles are not able to rotate. Therefore, Brown relaxation is generally excluded in this case.6 However, in low viscous media the relative contribution of Néel and Brown relaxation is not entirely evident. Generally, Néel relaxation is found to be the predominant, though not exclusive, relaxation mechanism. The decrease of the specific absorption rate (SAR) in a gel compared to a low viscous fluid is thereafter often attributed to the loss of Brown relaxation.6 The most frequently used method to discriminate between the two relaxation processes is measuring the temperature increase of both the ferrofluid as well as the gel containing the iron oxide nanoparticles.9 Furthermore, a distinction has been made using calculations of the specific absorption rates based on magnetic susceptibility measurements of particles in ferrofluid and incorporated in gel.8 In the present work, the heating of commercially available magnetic nanoparticles in a liquid and incorporated in a polymer matrix has been investigated at different magnetic field strengths. In the case of polymer incorporated particles, Brown

70 Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate) relaxation can be completely excluded, allowing the direct assessment of the relaxation process in the ferroﬂuid.

4.2 Materials & Methods

4.2.1 Materials

The iron oxide nanoparticles used in this study, aqueous ferroﬂuid EMG705, were used as purchased from Ferrotec, Germany. Methyl methacrylate (MMA) (99.5+ %) and lauric acid (99.5+ %) were purchased from Sigma Aldrich. Tetramethyl ammonium hydroxide (25 % in water) and potassium persulfate (99+ %) were purchased from VWR International.

4.2.2 Distribution of EMG705 Particles in Polymer

EMG705 nanoparticles were distributed in the polymer by mixing the aqueous ferroﬂuid with a p(MMA) latex and subsequent freeze drying using a Labconco Freezone 4.5 in combination with a Chemstar 1402N vacuum pump, operated at 84 10-3 mbar. Thereafter, the resulting powder was compounded into cylindrical · bars of approximately 10 mm length with a diameter of 3 mm, using a pre-heated custom built double screw compounder with a volume of 5 cm3, set at 190 ◦C and rotating at 100 rpm. The p(MMA) latex was made by emulsion polymerization at 80 ◦C using 50 g water, 5 g tetramethyl ammonium hydroxide, 1.8 g lauric acid, 70 g methyl methacrylate and 1 g potassium persulfate and was performed in an argon atmosphere with a reaction time of 1 hour. The diameter of the resulting p(MMA) particles was measured by dynamic light scattering (DLS) and was found to have a size of 55 14 nm. ±

4.2.3 Characterization

The size of the superparamagnetic iron oxide particles used in this study was determined using transmission electron microscopy (TEM). A single drop of diluted ferroﬂuid (100x w/w) was placed on a carbon coated copper grid. The distribution of the particles within the poly(methyl methacrylate) was studied with TEM as well. A Leica RM2165 rotary microtome was used to cut slices with an approximate thickness of 50 nm and the samples were placed on a carbon coated copper grid. TEM analysis was performed with a FEI Tecnai G2 Sphera cryo-TEM operated at

71 Chapter 4

200 kV. Dynamic light scattering was used to investigate the p(MMA) latex particle size as well as the presence of clusters of the iron oxide nanoparticles in the ferroﬂuid. A Coulter N4 Plus particle size analyzer was used, measuring 3 times for 900 seconds at 20 ◦C at an angle of 90 ◦. The presence of clusters was conﬁrmed by cryo-TEM analysis. The cryo-TEM samples were prepared on carbon coated copper grids by injection into liquid ethanol using a Vitrobot Mark III. The magnetization of the iron oxide nanoparticles and the particles incorporated in p(MMA) was determined at room temperature with a helium cooled MPMS 50 SQUID (superconducting quantum interference device) magnetometer.

4.2.4 Temperature Measurements

Heating experiments were performed to determine the amount of heat generated by the iron oxide nanoparticles, both in the ferrofluid as well as in the nanocomposite. In case of the ferrofluid, a Luxtron Fluoroptic temperature probe was placed in a glass tube with an inner diameter of 6 mm and a height of 40 mm containing 0.5 g EMG705 ferrofluid. The temperature of the nanocomposite was measured by placing 3 cylindrical bars around and in direct contact with the Luxtron probe. Subsequently, the samples were placed in a custom built setup generating an AC magnetic field with a frequency of 745 kHz. All samples were measured 3 times at different magnetic field strengths (2010, 2440 and 2850 A m-1). The amount of heat per gram of iron oxide generated, i.e. the specific absorption rate (SAR), was calculated based on the initial heating rate of the sample, the iron oxide content and the specific heat using Equation 4.3.

C dT SAR = p (4.3) f dt w ini

-1 Where SAR is the amount of heat generated per gram of iron oxide [W giron oxide], ◦ -1 -1 Cp the speciﬁc heat of the sample, (3.18 and 1.31 J C gsample for the ferroﬂuid and nanocomposite, respectively), fw the iron oxide weight fraction of the sample [-] and dT ◦ -1 dt ini the initial temperature increase [ C s ].

72 Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate)

(a) (b) Figure 4.1: Images of EMG705 by (a) conventional TEM of the nanoparticles and (b) cryo-TEM of the ferroﬂuid.

4.3 Results & Discussion

4.3.1 Characterization

The commercially available superparamagnetic iron oxide nanoparticles are characterized using TEM, magnetization measurements and DLS. TEM analysis shows that the core size of the EMG705 ferrofluid particles is in the superparamagnetic range, with an average size of 12.1 3.0 nm,12 see Figure 4.1a. Moreover, the magnetization ± curve of the particles shows zero coercivity and remanence, see Figure 4.2. The coercivity and remanence of a sample are the magnetic field required to reduce the magnetization of a sample to zero and the residual magnetization of a sample, respectively, after the magnetization of the sample has been driven into saturation. DLS analysis of the ferrofluid shows that the particles in the fluid are agglomerated in clusters of 118.1 48.4 nm. These clusters have been confirmed by cryo-TEM ± analysis of the EMG705 ferrofluid, see Figure 4.1b. The iron oxide nanoparticles incorporated in p(MMA) are observed to be in clusters of the same size order as in the ferrofluid, see Figure 4.3. These clusters are well distributed throughout the p(MMA) matrix. The decrease in specific absorption rate after immobilization of the nanoparticles in gel is often attributed to the loss of Brown relaxation upon immobilization.6,9 However, the effect of the distribution and agglomeration of the particles in the gel has never been studied in these investigations. Nonetheless, in the current work the particles in both the ferrofluid and the p(MMA) are present as 100 nm clusters ∼ allowing a direct comparison of the specific absorption rates.

73 Chapter 4

80 EMG705 nanoparticles

] p(MMA) containing EMG705 60

40 −1 iron oxide 20

−20

−40

−60 Magnetization [emu g −80

−10 −5 0 5 10 Magnetic field strength [kGs]

Figure 4.2: Magnetization curves of EMG705 nanoparticles and particles incorporated in p(MMA).

Figure 4.3: TEM image of EMG705 nanoparticles incorporated in p(MMA).

4.3.2 Temperature Measurements

Heating experiments are performed to assess the contribution of Néel and Brown relaxation to the heating of SPION in a ferrofluid by a magnetic field. The heating of SPION in a ferrofluid and incorporated in a p(MMA) matrix is measured 3 times using a fluoroptic fiber in an alternating magnetic field at different field strengths. The heating experiments show a significant temperature increase of both the ferrofluid and SPION containing p(MMA) depending on the magnetic field strength, see Figure 4.4 and 4.5. Initial heating rates have been found to be identical within experimental error for the 3 repetitive measurements for each sample and field strength.

74 Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate)

100 2010 A m−1 90 2440 A m−1 2850 A m−1 80 Initial heating rates

Temperature [°C] 40

20 0 50 100 150 200 Time [s]

Figure 4.4: Heating of EMG705 ferrofluid containing 33 wt% iron oxide nanoparticles at different field strengths, with the corresponding initial heating rates.

140 2010 A m−1 2440 A m−1 120 −1 2850 A m Initial heating rates 100

60 Temperature [°C]

20 0 20 40 60 80 100 120 140 Time [s]

Figure 4.5: Heating of p(MMA) containing 30 wt% EMG705 nanoparticles at diﬀerent ﬁeld strengths, with the corresponding initial heating rates.

Maximum initial heating rates of 0.64 and 1.42 ◦C s-1 are found for the ferrofluid and the nanocomposite, respectively, by applying the maximum field strength of 2850 A m-1. As expected, decreasing the field strength results in lower initial heating rates for both the ferrofluid and the composite. The calculated specific absorption rates at various field strengths are identical for both samples, Table 4.1. Since the iron oxide particles that are incorporated in p(MMA) cannot rotate, it is concluded that the heating contribution of Brown relaxation can be excluded in the ferrofluid as

75 Chapter 4

Table 4.1: Initial heating and specific absorption rate of iron oxide nanoparticles in ferrofluid and incorporated in p(MMA) at different magnetic field strengths.

Field strength Position Initial heating Speciﬁc absorption rate -1 ◦ -1 -1 [A m ] [ C s ] [Wgiron oxide] Ferroﬂuid 0.242 0.001 2.94 2010 Incorporated in p(MMA) 0.533 ± 0.006 2.94 ±

Ferroﬂuid 0.382 0.003 4.66 2440 ± Incorporated in p(MMA) 0.844 0.001 4.65 ±

Ferrofluid 0.642 0.003 7.82 2850 Incorporated in p(MMA) 1.42 ± 0.02 7.83 ± well. In order to verify the experimentally observed data with known theory of Néel and Brown relaxation, calculations of the relaxation times have been performed. Assuming a hydrodynamic particle size of 3 times the particle core size measured by TEM as determined previously for similar particles,8 the relaxation times for Néel and Brown relaxation in ferrofluid can be calculated using Equations 4.1 and 4.2. Using T = 300 K and η = 1.01 10-3 kg m-1 s-1, the calculation results in τ = 6.07 10-9 · · and 1.84 10-5 s, for Néel and Brown relaxation, respectively. Therefore, Brown · relaxation is not able to follow the fast change in magnetic field at a frequency of 745 kHz (τ = 2.14 10-7 s). Based on these calculations, it can be concluded that · Néel relaxation will be the main relaxation process in the ferrofluid, which is in accordance with the heating experiments.

4.4 Conclusion

Experimental results of heating superparamagnetic iron oxide nanoparticles in a ferrofluid as well as incorporated in poly(methyl methacrylate) reveal identical specific absorption rates for both samples at 3 different magnetic field strengths. Due to the exclusion of Brown relaxation for the particles incorporated in p(MMA), it is concluded that heating by Brown relaxation can also be neglected in the ferrofluid and, as a consequence, Néel relaxation is the only occurring relaxation process that contributes to heating. At a frequency of 745 kHz this is confirmed by calculations of

76 Relaxation processes of superparamagnetic iron oxide nanoparticles in liquid and incorporated in poly(methyl methacrylate) the relaxation times, which show that Brown relaxation is not able to follow the fast change of field direction. Moreover, these calculations also suggest that for 12 nm sized particles in low viscous media, Néel relaxation will also predominate at lower frequencies due to the significantly shorter relaxation times.

77 Chapter 4

References

[1] Jung, C.W., Jacobs, P. Magnetic Resonance Imaging, 13 (1995), 661-674.

[2] Gupta, A.K., Gupta, M. Biomaterials, 26 (2005), 3995-4021.

[3] Zhang, J.L., Srivastava, R.S., Misra, R.D.K. Langmuir, 23 (2007), 6342-6351.

[4] Alexiou, C., Jurgons, R., Schmid, R.J., Bergemann, C., Henke, J., Erhardt, W., Huenges, E., Parak, F. Journal of Drug Targeting, 11 (2003), 139-149.

[5] Melville, D., Paul, F., Roath, S. Nature, 255 (1975), 663-750.

[6] Hergt, R., Hiergeist, R., Zeisberger, M., Gl¨ockl, G., Weitschies, W., Ramirez, L.P., Hilger, I., Kaiser, W.A. Journal of Magnetism and Magnetic Materials, 280 (2004), 358-368.

[7] Bahaj, A.S., James, P.A.B., Moesschler, F.D. Journal of Applied Physics, 83 (1998), 6444-6446.

[8] Hergt, R., Hiergeist, R., Hilger, I., Kaiser, W.A., Lapatnikov, Y., Margel, S., Richter, U. Journal of Magnetism and Magnetic Materials, 270 (2004), 345-357.

[10] Hiergeist, R., Andr¨a, W., Buske, N., Hergt, R., Hilger, I., Richter, U., Kaiser, W. Journal of Magnetism and Magnetic Materials, 201 (1999), 420-422.

[11] N´eel, L. Annales Geophysicae, 5 (1949), 99-136.

[12] Dutz, S., Hergt, R., Mürbe, J., Müller, R., Zeisberger, M., Andrä, W., Töpfer, J., Bellemann, M.E. Journal of Magnetism and Magnetic Materials, 308 (2007), 305-312.

78 Chapter 5

Repetitive on-demand drug release from iron oxide incorporated polymeric matrices

Drug release from a polymeric matrix has been externally triggered using an alternating magnetic field in order to develop an on-demand drug delivery device. Superparamagnetic iron oxide nanoparticles have been distributed in a poly(methyl methacrylate) core, coated with a thermoresponsive layer of poly(butyl methacrylate-stat-methyl methacrylate) containing ibuprofen as a model drug. The effect of the loading of iron oxide particles in the core and of ibuprofen in the coating on the drug release has been investigated. The release rate of ibuprofen significantly increases upon exposure to the magnetic field and increases with increasing iron oxide loading. An increase in ibuprofen concentration in the polymer matrix increases the release rate in both the on and off situation. Chapter 5

5.1 Introduction

For several decades, a major challenge in the field of controlled drug delivery has been the controlled delivery of an active in a pulsatile manner. This challenge has been addressed from two directions, of which one is the formation of a system that releases its drug load at a predetermined time or in pulses with predetermined intervals.1–3 A different approach is the use of external or internal stimuli including pH,4–6 ultrasound,7–10 electrical,11,12 mechanical13 and magnetic triggers.14–16 Internally stimulated systems have the advantage of self regulation of the drug release. However, after admission the release can not be adjusted apart from removal of the dosage form. Externally stimulated systems are able to release a drug on-demand when the physician or patient requires drug release. In several studies using a magnetic field to trigger the release of the active substance, magnetic heating of superparamagnetic iron oxide nanoparticles has been applied. The nanoparticles have no remanent magnetization after the applied field is removed, are biocompatible17 and generate heat primarily due to Néel relaxation, see Chapter 4 In most cases the lower critical solution temperature (LCST) of a polymer is used as a thermoresponsive switch to induce drug release, whereby the alternating magnetic field heats the nanoparticles, and, subsequently, the polymer collapses when heated above its LCST.18–21 Collapse of the polymer chains above the LCST results in shrinkage of the polymer layer. Even though the LCST is a reversible mechanism, the majority of the incorporated drug is often squeezed out during one exposure to the magnetic field. Intense heating of such systems can be used to irreversibly rupture the system, triggering a burst-like drug release.20 As another approach, magnetoliposomes have been triggered for instant release using an alternating magnetic field.22,23 In this study, a reversible thermoresponsive switch is used for multiple pulsatile drug release using superparamagnetic iron oxide in combination with an alternating magnetic field as an external trigger. The thermoresponsive switch is based on the increase of the diffusion coefficient of a solute in a polymer matrix with increasing 24 temperature, especially around the glass transition, Tg, of the polymer. Ibuprofen has a clear plasticizing effect on p(BMA-MMA), lowering the bulk glass transition temperature.25 However, the rate determining diffusion coefficient is located at the outer skin layer of the thermoresponsive coating. As the ibuprofen concentration in the phosphate buffered saline (PBS) solution is very low, sink conditions can be assumed and, apart from the partition coefficient, the concentration of ibuprofen at the outer skin of the coating is negligibly low. Therefore, the glass transition temperature of the outer skin is equal to the Tg of the pure polymer and the rate

80 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices determining diffusion coefficient is equal to the diffusion coefficient at negligibly low ibuprofen concentration. Consequently, the plasticizing effect of the ibuprofen does not influence the release from the sample and the higher concentration of ibuprofen merely increases the driving force for diffusion.

5.2 Materials & Methods

5.2.1 Materials

The EMG705 iron oxide nanoparticles used in this study were purchased from Ferrotec, Germany. Poly(methyl methacrylate) (p(MMA)) (MW 120,000), poly(butyl methacrylate-stat-methyl methacrylate) (p(BMA-MMA)) (MW 150,000, 75 mol% BMA), methyl methacrylate (MMA) (99.5+ %), lauric acid (99.5+ %) and phosphate buffered saline (PBS) were purchased from Sigma Aldrich. The phosphate buffered saline solution contained NaCl (120 mM), KCl (2.7 mM) and phosphate buffer (10 mM) with a pH of 7.4 at 25 ◦C. Tetramethyl ammonium hydroxide (25 % in water), potassium persulfate (99+ %) and trifluoracetic acid (TFA) (99.8+ %) were purchased from VWR International. Acetonitrile (HPLC - Supra gradient) and dichloromethane (DCM) (99.9+ %) were purchased from Biosolve. The monomer units of the various polymers and the structure of ibuprofen are shown in Figure 5.1.

CH3

CH H C H C H C 3 3 3 H2C CH3

n n m O O O O O O

H3C H2C H3C

CH 2

H2C CH O H3C C CH 3 OH (a) (b) (c)

Figure 5.1: Molecular structure of (a) p(MMA), (b) p(BMA-MMA) and (c) of the model drug ibuprofen.

81 Chapter 5

5.2.2 Preparation of Iron Oxide Containing Heatable Core

The water-based iron oxide nanoparticles, EMG705, were distributed in p(MMA) in two ways as reported in detail in Chapter 3. In the first method, 2.5 g of p(MMA) was added to a pre-heated custom built double cone screw compounder with a volume of 5 cm3, set at 190 ◦C and rotating at 100 rpm. Subsequently, 7.5 mL of EMG705 ferrofluid (33 wt% particles) was added drop wise. After 15 minutes, cylindrical bars of approximately 10 mm length with a diameter of 3 mm were extruded from the compounder. Alternatively, EMG705 nanoparticles were distributed by mixing a p(MMA) latex and the EMG705 ferrofluid. This aqueous mixture was freeze dried using a Labconco Freezone 4.5 in combination with a Chemstar 1402N vacuum pump, operated at 84 10-3 mbar. The resulting powder was compounded as described above. The · p(MMA) latex was prepared by emulsion polymerization at 80 ◦C using 50 g water, 5 g tetramethyl ammonium hydroxide, 1.8 g lauric acid, 70 g methyl methacrylate (MMA) and 1 g potassium persulfate in an argon atmosphere. The resulting p(MMA) particles were measured by dynamic light scattering (DLS) and were found to have a size of 54.8 14.4 nm. ±

5.2.3 Preparation of Thermoresponsive Release Coating

To release the model drug, ibuprofen, using an alternating magnetic ﬁeld, the cylindrical bars of p(MMA) containing the iron oxide nanoparticles were coated with a 0.3 mm thick layer of p(BMA-MMA) containing ibuprofen. The coating was applied by dip coating the bar in a polymer solution (10 wt% in DCM) containing the appropriate amount of ibuprofen. In determined cases, dip coating was repeated with p(BMA-MMA) without ibuprofen after evaporation of the DCM. Subsequently, the DCM was allowed to evaporate during several days prior to a release experiment.

5.2.4 Release Measurements

Release experiments were performed by placing the desired number of samples in a flask containing 10 to 30 mL phosphate buffered saline and placing these flasks in a water bath set at the required temperature of 20 or 37 ◦C. The samples were pretreated by moving them once for 30 minutes to a secondary water bath set at 65 ◦C. Subsequently, the PBS buffer was refreshed and the samples were placed again in the initial water bath at the required temperature. Triggering drug release from the samples was done by placing the sample flask for 15

82 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices to 60 minutes in an alternating magnetic field setup. The setup generated a magnetic field with a field strength of 2850 A m-1 and a frequency of 745 kHz. During the release experiments performed at 37 ◦C, a small water bath was placed within the magnetic field setup. The sample flask was placed in this water bath during the triggering in order to maintain a steady buffer temperature during the experiment. The amount of released ibuprofen was measured by taking samples from the PBS solution, which were analyzed by high performance liquid chromatography (HPLC). An equal amount as the sample of fresh buffer was added to the flask to maintain a constant volume. The HPLC analysis was done on a Shimadzu VP Series HPLC using a solution of 50/50 vol% acetonitrile and water with 0.05 vol% trifluoracetic acid as the mobile phase, with a flow of 1.2 mL min-1 (LC-10AD VPA solvent pump). A reversed phase C18 column (Discovery HS, 150 x 4.6 mm - 3 µm) was used as the stationary phase and an UV-spectrophotometer (Shimadzu, SPD-10Avp) was used to measure the ibuprofen content at a wavelength of 223 nm.

5.3 Results & Discussion

Release experiments have been performed on samples consisting of a p(MMA) core with iron oxide nanoparticles, which are dip coated with a thermoresponsive layer of p(BMA-MMA) containing ibuprofen as a model drug. As the samples are exposed to an alternating magnetic ﬁeld, the iron oxide in the composite core is heated. Subsequently, the temperature of the thermoresponsive coating is increased, enhancing drug release.

5.3.1 Eﬀect of Iron Oxide Concentration

Due to heating of the iron oxide nanoparticles in the core, the temperature of the thermoresponsive p(BMA-MMA) coating increases. As a result, the diffusion coefficient of the incorporated ibuprofen increases and release of ibuprofen from the sample is enhanced, see Figure 5.2. As the flask is removed from the magnetic field and is placed back in the 20 ◦C water bath, the sample is cooled and the diffusion coefficient decreases again, lowering the release rate to the rate prior to the magnetic field exposure. Multiple exposures to the magnetic field in the following days result in similar release rates, showing that samples can be used for reproducible multiple drug release doses. In fact, this is the first proof-of-principle of magnetically induced on-demand drug release based on the diffusivity of a thermoresponsive coating. As expected, 50 wt% concentration of iron oxide nanoparticles in the core results in

83 Chapter 5

60 5 ] −1 55 4

50 3 45 2 40

1 35 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 30 0 65 75 85 95 105 115 Time [h]

(a)

60 5 ] −1 55 4

50 3 45 2 40

1 35 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 30 0 65 75 85 95 105 115 Time [h]

(b)

Figure 5.2: On-demand release from p(BMA-MMA) samples including 5 wt% ibuprofen coated on p(MMA) cores (prepared by direct compounding) containing (a) 30 wt% and (b) 50 wt% of iron oxide nanoparticles in a 20 ◦C bath exposed to the magnetic field for 60 minutes. The release shown here is the cumulative release of three identical samples in a single jar. higher release rates compared to 30 wt% iron oxide as the magnetic field is applied, see Figure 5.2. Because of the higher concentration of iron oxide, the temperature increase of the core (Chapter 3) and, subsequently, of the thermoresponsive coating is higher, resulting in a higher diffusion coefficient. However, when the magnetic field

84 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices is not applied, the release rate is independent of the amount of iron oxide, as the temperatures of both samples are equal. Therefore, the ratio between the release rate upon triggering by the magnetic field and the release rate without triggering, the on/off ratio, is significantly higher for the 50 wt% iron oxide samples (on/off ratio = 25) compared to the 30 wt% iron oxide samples (on/off ratio = 6.5).

5.3.2 Eﬀect of Ibuprofen Concentration

As ibuprofen is released from the sample by diffusion, the release rate depends on the diffusion coefficient and on the concentration of the ibuprofen dissolved in the polymer matrix. Therefore, on-demand release from p(BMA-MMA) coated cores (50 wt% iron oxide, incorporated by direct compounding) has been performed using four identical samples in 10 mL PBS solution, triggering for 60 minutes from 20 ◦C. By increasing the ibuprofen concentration from 5 to 9 and 17 wt% the driving force for mass transport by diffusion increases, due to the larger difference in concentration inside and outside the sample. As a consequence, the release rate in both the on and off situation increases, see Figure 5.3, and the on/off ratio is independent of the ibuprofen concentration, remaining at a value of around 10.

5.3.3 Release from Coatings with High Drug Loading

Due to the release of ibuprofen from the samples, the concentration of ibuprofen dissolved in the polymer matrix decreases with time and, subsequently, the driving force for diffusional transport decreases. A constant driving force is important to achieve reproducible on-demand drug delivery. By using a high concentration of ibuprofen in the thermoresponsive coating, the released amount of ibuprofen is negligible compared to the total amount, ensuring a constant, maximum driving force for diffusion. Apart from this, the temperature in the off situation of the final application within the human body will be 37 ◦C. Therefore, experiments with a high ibuprofen concentration have been performed keeping the PBS solution at this temperature both in the water bath as well as during exposure to the magnetic field. A flask containing one sample of p(BMA-MMA) coated on a p(MMA) core (containing 50 wt% iron oxide) in 30 mL PBS has been exposed to the magnetic field for 15 minutes. The p(BMA-MMA) coating has been incorporated using an ibuprofen concentration of 29 wt%. The composite core has been prepared using a freeze drying method, resulting in an increased heating of the core material as described

85 Chapter 5

300 30 ] −1 25 250 20

200 15

10 150 5 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 100 0 140 150 160 170 180 190 Time [h]

(a)

800 30 ] −1 25 750 20

700 15

10 650 5 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 600 0 140 150 160 170 180 190 Time [h]

(b)

1000 30 ] −1 25 950 20

900 15

10 850 5 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 800 0 140 150 160 170 180 190 Time [h]

(c)

Figure 5.3: On-demand release from p(BMA-MMA) coated p(MMA) cores containing 50 wt% of iron oxide particles and diﬀerent ibuprofen contents of (a) 5 wt%, (b) 9 wt% and (c) 17 wt% from 20 ◦C exposed to the magnetic ﬁeld for 60 minutes. The release shown here is the cumulative release of three identical samples in a single jar.

86 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices

1900 250 ] −1 1700 200

1500 150

1300 100

1100 50 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 900 0 25 35 45 55 65 75 Time [h]

Figure 5.4: On-demand release from one sample of p(BMA-MMA) with a high ibuprofen loading (29 wt%) coated p(MMA) core containing 50 wt% of iron oxide particles in a 37 ◦C bath exposed to the magnetic ﬁeld for 15 minutes.

5 From 20 °C, directly compounded core From 37 °C, freeze dried core 4

1 Temperature increase [°C]

0 0 50 100 150 200 250 Time [s]

Figure 5.5: The temperature increase at the outer layer of the coating of samples with a core prepared using direct compounding in a 20 ◦C bath and samples with a core prepared by the freeze drying method from 37 ◦C, modeled using Comsol. in Chapter 3. As both the concentration of ibuprofen and the temperature of the coating and, thus, the diffusion coefficient of the ibuprofen are higher than in the case of 20 ◦C as the starting temperature, the release rate is significantly higher in both the on and

87 Chapter 5 off situation, see Figure 5.4. Since the heating rate of the p(MMA) composite core is significantly affected by the method of preparation, the temperature increase of the outer layer of the coating is increased, as modeled using Comsol 3.5, Figure 5.5, resulting in a measured on/off ratio of around 16.5, despite the higher off release. In this finite element model the temperature of the outer layer of the coating has been evaluated using a core prepared by direct compounding (heat generated during exposure = 0.09 W) and a core prepared by the freeze drying method (heat generated = 0.31 W). Additional parameters used can be found in the Appendix.

1600 250 ] −1 1400 200

1200 150

1000 100

800 50 Release Rate Ibuprofen [µg h Cumulative Release Ibuprofen [µg] 600 0 45 55 135 145 155 165 Time [h]

Figure 5.6: On-demand release from one sample of p(BMA-MMA) with a high ibuprofen loading (29 wt%) coated on ap(MMA) core containing 50 wt% of iron oxide particles in a 37 ◦C bath exposed to the magnetic ﬁeld for 15 minutes. Samples were additionally coated with an empty p(BMA- MMA) layer.

To decrease the off release at 37 ◦C it is attempted to add an additional coating layer of p(BMA-MMA) without ibuprofen over the ibuprofen containing layer to increase the distance for diffusional transport of ibuprofen. Indeed, with this additional coating layer, the release rate when the sample is not exposed to the magnetic field is decreased, see Figure 5.6. However, the increased diffusion distance in combination with the increase of distance between the heat source, i.e. the core containing the iron oxide particles, and the outer layer of the coating also results in a lower release rate upon exposure to the magnetic field, resulting in an on/off ratio of only 4.5. In order to calculate the effect of the additional distance on the temperature at the outer layer of the thermoresponsive coating, the heating has again been modeled using a finite element method using Comsol 3.5. Even though the sample has been

88 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices

41.5 Without empty layer 41 With empty layer 40.5 40 39.5 39 38.5 38 Temperature [°C] 37.5 37

36.5 0 50 100 150 200 250 Time [s]

Figure 5.7: Effect of an additional layer of p(BMA-MMA) (thickness of 20 % of the initial coating) on the temperature at the outer layer of the thermoresponsive layer, modeled using Comsol. exposed to the magnetic field for 60 minutes, it is sufficient to model the exposure for 2 minutes, from t = 20 to t = 140 seconds, since the sample reaches a steady state temperature within 1 minute, see Figure 5.7. This figure shows that the additional layer has only a minor effect on the outer temperature reached, which indicates that the increase in the distance for diffusion is the predominante cause of the decrease in release rate of ibuprofen in the on state.

5.4 Conclusion

A drug delivery device has been prepared from which repetitive drug release can be reproducibly induced using an alternating magnetic field as an external trigger. The on-demand drug release was performed using the increase of the diffusion coefficient of a solute in a polymer matrix with increasing temperature. The heating of the polymer by the magnetic field has been accomplished by incorporation of superparamagnetic iron oxide nanoparticles in the core of the matrix. A significant increase of the release rate of the model drug, ibuprofen, has been achieved upon exposure of the sample to the alternating magnetic field. After the magnetic field has been removed, the release rate decreases back to the rate prior to application of the magnetic field, demonstrating the reversibility of the thermoresponsive switch. Multiple consecutive exposures to the external trigger

89 Chapter 5 result in a similar increase of the release rate. A higher concentration of iron oxide nanoparticles in the core of the device increases the heat generation and therefore, increases the release rate upon exposure, while the release rate without exposure is not influenced. However, an increase of the concentration of ibuprofen in the thermoresponsive coating increases the release rate in both the on and off situation. Furthermore, increasing the experimental temperature from 20 ◦C to the temperature of the final application, 37 ◦C, increases both the on and off release leading to an on/off ratio of 16.5. Adding an additional empty coating layer decreased the on release rate more than the off release rate, which was demonstrated to be due to the increased diffusion distance.

90 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices

Appendix

Table 5.1: Input parameters for modeling in Comsol.

Parameter Symbol Value Unit

Sample length l 0.01 [m] Sample diameter d 0.003 [m]

Characteristic length Lk d [m] ◦ Ambient temperature Tbl 20or37 [ C]

Powerdirect compounded core P 0.09 [W]

Powerfreeze dried core P 0.311 [W] -3 Density core ρcore 1705 [kg m ] 26 -3 Density Tg layer ρTg layer 1066 [kg m ] -1 -1 Thermal conductivity core kcore 0.29 [W m K ] 26 -1 -1 Thermal conductivity Tg layer kcore 0.19 [W m K ] -1 -1 Heat capacity core Cp,core 1260 [Jkg K ] 26 -1 -1 Heat capacity Tg layer Cp,core 1500 [Jkg K ] Emissivity ǫ 0.9 [-]

Table 5.2: Water properties and dimensionless numbers for modeling in Comsol.27

Parameter Symbol Value Unit

Nu·kwater -2 -1 Heat transfer coeﬃcient hf [W m K ] Lk

Nusselt correlation Nu 0.48 Ra0.25 [-] ·

Rayleigh number Ra Gr Pr [-] ·

· · 3 · βwater g Lk ∆T Grashof number Gr 2 [-] νwater

· Prandtl number Pr µwater Cp,water [-] kwater

Table continued on next page

91 Chapter 5

Table 5.2 – continued from previous page Parameter Symbol Value Unit

Temperature diﬀerence ∆T T T∞ [K] −

T −T∞ Temperature boundary layer Tbl 2 [K]

1.3759888667 10-9 T 6 · · bl -2.4790027049 10-6 T 5 · · bl +1.8549599965 10-3 T 4 · · bl Heat capacity water C -7.3792022835 10-1 T 3 [J kg-1 K-1] p,water · · bl +1.6461861532 102 T 2 · · bl -1.9530990636 104 T · · bl +9.6735169253 105 ·

2,5941160868 10-17 T 6 · · bl -4,9752344767 10-14 T 5 · · bl +3,9741981497 10-11 T 4 · · bl Kinematic viscosity water ν -1.6926235746 10-8 T 3 [m2 s-1] water · · bl +4.0545248068 10-6 T 2 · · bl -5.1803600574 10-4 T · · bl +2.7589834589 10-2 ·

-1.1838177032 10-7 T 4 · · bl +1.6622737154 10-4 T 3 · · bl Density water ρ -9.0246260933 10-2 T 2 [kg m-3] water · · bl +2.1777961060 101 T · · bl -9.4391841907 102 ·

-4.239026 10-12 T 4 · · bl +5.923936 10-9 T 3 · · bl -2 2 -1 Thermal expansion βwater -3.120298 6 Tbl [K ] coeﬃcient water · · +7.398328 10-4 T · · bl -6.645996 10-2 ·

Table continued on next page

92 Repetitive on-demand drug release from iron oxide incorporated polymeric matrices

Table 5.2 – continued from previous page Parameter Symbol Value Unit

2,5318197442 10-14 T 6 · · bl -4,8556640447 10-11 T 5 · · bl +3.8786380338 10-8 T 4 · · bl Viscosity water µ -1.6519196941 10-5 T 3 [Pa s] water · · bl · +3.9570647567 10-3 T 2 · · bl -5.0559891863 10-1 T · · bl +2.6929010495 101 ·

-1,4860681136 10-11 T 6 · · bl +2.8303246844 10-8 T 5 · · bl -2.2438215478 10-5 T 4 · · bl Thermal conductivity water k +9.4777682617 10-3 T 3 [W m-1 K-1] water · · bl -2.2496574292 100 T 2 · · bl +2.8451298262 102 T · · bl -1.4977764053 104 ·

93 Chapter 5

References

[1] Richards Grayson, A.C., Choi, I.S., Tyler, B.M., Wang, P.P., Brem, H., Cima, M.J., Langer, R. Nature Materials, 2 (2003), 767-772.

[2] Qiu, L.Y., Zhub, K.J. International Journal of Pharmaceutics, 219 (2001), 151- 160.

[3] Moriyama, K., Yui, N. Journal of Controlled Release, 42 (1996), 237-248.

[4] Zhang, J., Peppas, N.A. Macromolecules, 33 (2000), 102-107.

[5] Heller, J., Trescony, P.V. Journal of Pharmaceutical Sciences, 68 (1979), 919- 921.

[6] Brannon-Peppas, L., Peppas, N.A. Biomaterials, 11 (1990), 635-644.

[7] Kwok, C.S., Mourad, P.D., Crum, L.A., Ratner, B.D. Journal of Biomedical Materials Research Part A, 57 (2001), 151-164.

[8] Kim, H.J., Matsuda, H., Zhou, H., Honma, I. Advanced Materials, 18 (2006), 3083-3088.

[9] Aschkenasya, C., Kost, J. Journal of Controlled Release, 110 (2005), 58-66.

[10] Kost, J., Leong, K., Langer, R. Proceedings of the National Academy of Sciences of the Uniteds States of America, 86 (1989), 7663-7666.

[11] Kwon, I.C., Bae, Y.H., Kim, S.W. Nature, 354 (1991), 291-293.

[12] Santini Jr, J.T., Cima, M.J., Langer, R. Nature, 397 (1999), 335.

[13] Lee, K.Y., Peters, M.C., Anderson, K.W., Mooney, D.J. Nature, 408 (2000), 998-1000.

[14] Edelman, E.R., Kost, J., Bobeck, H., Langer, R. Journal of Biomedical Materials Research, 19 (1985), 67-83.

[15] Kost, J., Wolfrum, J., Langer, R. Journal of Biomedical Materials Research, 21 (1987), 1367-1373.

[16] Saslawski, O., Weingarten, C., Benoit, J.P., Couvreur, P. Life Sciences, 42 (1988), 1521-1528.

94 References

[17] Chan, D.C.F., Kirpotin, D.B., Bunn Jr., P.A., Physical chemistry and in vivo tissue heating properties of colloidal magnetic iron oxides with increased power absorption rates, in: H¨afeli, U., et al. (eds.) Scientiﬁc and Clinical Applications of Magnetic Carriers, Plenum Press, New York, 1997.

[18] M¨uller-Schulte, D., Schmitz-Rode, T. Journal of Magnetism and Magnetic Materials, 302 (2006), 267-271.

[19] Zhang, J.L., Srivastava, R.S., Misra, R.D.K. Langmuir, 23 (2007), 6342-6351.

[20] Liu, T.Y., Hu, S.H., Liu, D.M., Chen, S.Y., Chen, I.W. Nano Today, 4 (2009), 52-65.

[21] Csetneki, I., Filipcsei, G., Zr´ınyi, M. Macromolecules, 39 (2006), 1939-1942.

[22] Babincova, M. Pharmazie, 50 (1995), 702-703.

[23] Babincova, M., Cicmanec, P., Altanerova, V., Altaner, C., Babinec, P. Bioelec- trochemistry, 55 (2002), 17-19.

[24] Keurentjes, J.T.F., Kemmere, M.F., Bruinewoud, H., Vertommen, M.A.M.E., Rovers, S.A., Hoogenboom, R., Stemkens, L.F.S., P´eters, F.L.A.M.A., Tielen, N.J.C., van Asseldonk, D.T.A., Gabriel, A., Joosten, B., Marcus, M.A.E. Angewandte Chemie International Edition, 48 (2009), 9867-9870.

[25] Siepmann, F., Le Brun, V., Siepmann, J. Journal of Controlled Release, 115 (2006), 298-306.

[26] CRC Polymers, Polymers: A Properties Database, CRC Press, Online Edition, 2007.

[27] Perry, R.H., Green, D.W., Perry’s Chemical EngineersHandbook,´ McGraw-Hill, Singapore, 1997.

95 Chapter 5

96 Chapter 6

Repetitive on-demand drug release from polymeric matrices containing a cylindrical iron core

For the purpose of repetitive on-demand drug delivery, a system has been developed using an alternating magnetic field as external trigger. A thermoresponsive polymer, poly(styrene-stat-butyl methacrylate), containing ibuprofen as a model drug, has been coated on an iron rod. During exposure of the device to the alternating magnetic field (on situation) the drug release rate significantly increases compared to the release rate without exposure (off situation), up to factors exceeding 2000. When the magnetic field trigger is removed, the release rate decreases back to the rate prior to exposure, because of the reversible glass transition temperature switch. Drug release has been triggered multiple times on and off, showing the ability of repetitive on-demand drug delivery. Chapter 6

6.1 Introduction

For several decades, in the field of controlled delivery efforts have been made to release a drug in a pulsatile manner. One way to do this is the formation of a system that releases its payload at a predetermined time or in pulses with predetermined intervals.1–3 A different approach is the use of external or internal stimuli including pH,4–6 near-infrared radiation,7–9 ultrasound,10–13 electrical,14,15 mechanical16 and magnetic triggers.17–19 Internally stimulated systems have the advantage of self regulation of the drug release, however, similar to predetermined release systems, after admission the release can not be adjusted, apart from removal. Externally stimulated systems are able to release drug on-demand as the physician or patient requires. In vitro and in vivo on-demand release has been reported by exerting pressure on, and squeezing additional drug out of, pores in a matrix by movement of a magnet in the drug delivery device using a low frequency magnetic field.17,18 Furthermore, the lower critical solution temperature (LCST) phase transition of a polymer has been used in on-demand drug delivery, using an alternating magnetic field as external trigger. Iron oxide nanoparticles, incorporated in the polymer are heated remotely by the magnetic field, subsequently heating the polymer above its LCST resulting in collapse of the polymer.20–23 Collapse of the polymer chains above the LCST results in shrinkage of the polymer and squeezing out of the drug incorporated within the polymer. Even though the LCST is a reversible mechanism, the majority of the incorporated drug is squeezed out during one exposure to the magnetic field in most examples. Additionally, intense heating of such systems is used to irreversibly rupture the system, triggering a burst-like drug release.22 Moreover, magnetoliposomes have been triggered for instantaneous release using an alternating magnetic field.24,25 In recent work, we demonstrated a reversible thermoresponsive switch based on the glass transition temperature of the polymer, Tg, for multiple pulsatile drug release using an external trigger.26 Also see Chapters 5 and 7. The thermoresponsive switch used in these studies is based on the significant increase in the diffusion coefficient of a solute in a polymer matrix with increasing temperature, especially around the 26 glass transition temperature, Tg, of the polymer. Even though the model drug ibuprofen has a clear plasticizing effect on the polymer used,27 lowering the bulk glass transition temperature, the rate determining diffusion coefficient is found at the outer skin layer of the thermoresponsive coating. As the ibuprofen concentration in the phosphate buffered saline (PBS) solution is very low, sink conditions can be assumed and, apart from the partition coefficient, the concentration of ibuprofen

98 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core at the outer skin of the coating is negligibly low. Therefore, the glass transition temperature of the outer skin is equal to the Tg of the pure polymer and the rate determining diffusion coefficient is equal to the diffusion coefficient at negligibly low ibuprofen concentration. Previously used triggers for external heating of the polymer include near-infrared radiation, for which a near-infrared dye is coated on the thermoresponsive polymer.26 Since the penetration depth of near-infrared radiation is small, limiting the application, a magnetic field has also been used as trigger to increase the temperature of the polymer. Therefore, the thermoresponsive polymer is coated on the core of a polymer matrix that contains superparamagnetic iron oxide nanoparticles (Chapter 5) However, only limited drug release was achieved due to insufficient heating of the outer switching layer. To further increase repetitive on-demand drug release, the thermoresponsive coating has been applied on an iron rod in the present work. Upon exposure of the sample to an alternating magnetic field, the temperature of the iron rod is increased because of induction heating. Subsequently, the thermoresponsive coating is heated from below to above the glass transition temperature of the polymer, significantly increasing the diffusion coefficient of ibuprofen in the polymer, resulting in drug release.

6.2 Materials & Methods

6.2.1 Materials

Iron rod cores, with a diameter of 1.1 mm and a length of 15 mm, were cut from 50 mm Kangaro paperclips purchased from Central point, the Netherlands. The iron composition was confirmed using a Curie temperature measurement using a TA Instruments Q500 thermogravimetric analyzer. Poly(styrene-stat-butyl methacrylate) (p(S-BMA)) (Mw 200,000, 50 % styrene) was purchased from Scientific Polymer Products, USA). Poly(ethylene-stat-vinyl acetate) (p(E-VA)) (40 wt% vinyl acetate) and phosphate buffered saline (PBS) were purchased from Sigma Aldrich. The phosphate buffered saline solution contained NaCl (120 mM), KCl (2.7 mM) and phosphate buffer (10 mM) with a pH of 7.4 at 25 ◦C. Acetonitrile (HPLC - Supra gradient) and dichloromethane (DCM) (99.9+ %) were purchased from Biosolve. Trifluoracetic acid (TFA) (99.8+ %) was purchased from VWR International and ibuprofen was obtained from Fargon.

99 Chapter 6

6.2.2 Diﬀusion Measurements

The apparent diffusion coefficient of ibuprofen in poly(styrene-stat-butyl methacrylate) at different initial ibuprofen concentrations and different temperatures was determined by measuring the release of ibuprofen from polymer discs in time. For this purpose, different concentrations of ibuprofen, 5, 10, 16.7, 22, and 26.7 wt%, were incorporated in p(S-BMA) discs by solvent casting the discs using dichloromethane (10 wt% polymer). The samples were dried for several days in air followed by 24 hours under reduced pressure. Discs with a diameter of 2.54 cm and a thickness of 0.5 mm were cut from the films and set of two discs were fused together with a press at 100 ◦C into 1 mm thick discs for the diffusion experiments. The release of ibuprofen from the polymer was measured by placing a disc in a flask containing 50 mL of phosphate buffered saline and placing the flask in a water bath set at the required temperature. The amount of released ibuprofen was measured by taking samples from the PBS solution, which were analyzed by high performance liquid chromatography (HPLC). After taking a sample, an equal amount of fresh buffer was added to the flask to maintain a constant volume. The HPLC analysis was done on a Shimadzu VP Series HPLC using a solution of 50/50 vol% acetonitrile and water with 0.05 vol% trifluoracetic acid as the mobile phase, with a flow rate of 1.2 mL min-1 (LC-10AD VPA solvent pump). A reversed phase C18 column (Discovery HS, 150 x 4.6 mm - 3 µm) was used a the stationary phase and an UV-spectrophotometer (Shimadzu, SPD-10Avp) was used to measure the ibuprofen content at a wavelength of 223 nm. For a constant diffusion coefficient, diffusion from a disc can be described by the one dimensional version of Fick’s second law,28

∂C ∂2C = D (6.1) ∂t ∂x2 where C is the concentration of ibuprofen [mol m-3], t is the time [s], D the diﬀusion coeﬃcient [m2 s-1] and x axial position [m]. Since the PBS volume was large in comparison to the disc, sink conditions were assumed. Furthermore, an initially homogenous distribution of ibuprofen in the polymer matrix was assumed as well. In such a case, Fick’s second law can be described by an analytical solution, Equation 6.2:29 ∞ 2 2 Mt 8 -D·(2n+1) π t =1 e L2 (6.2) M∞ − (2n + 1)2 · n=0 X

100 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core

Where Mt and M∞ are the cumulative release of ibuprofen in time and the total amount of ibuprofen in the polymer disc, respectively, and L is the full thickness of the disc. Up to Mt/M∞ = 0.6 this analytical solution can be approximated by

M D t t =4 · (6.3) M∞ · π L2 r · Therefore, the square root of the apparent diffusion coefficient is equal to the slope -1 1/2 -1 of a graph of Mt/M∞ versus 4π t L . Since the apparent diffusion coefficient depends on the concentration of ibuprofen, due to the plasticizing effect of ibuprofen, the initial slope has been used to calculate the diffusion coefficient.

6.2.3 Preparation of Release Coating

The thermoresponsive polymer, poly(styrene-stat-butyl methacrylate), containing ibuprofen as a model drug, was applied as a 0.7 mm coating on an iron rod with a diameter of 1.1 mm and a length of 15 mm. The coating was applied by dip coating the rod in a polymer solution (10 wt% in DCM) containing the appropriate amount of ibuprofen. In several cases, an additional 0.6 mm coating of poly(ethylene-stat-vinyl acetate) without ibuprofen was applied as a thermal isolation layer by dip coating. Because of the low glass transition temperature of p(E-VA), the diﬀusion coeﬃcient of ibuprofen is assumed to be much higher in p(E-VA) than in the thermoresponsive layer. After dip coating, the DCM was allowed to evaporate during several days prior to a release experiment. Oversaturated samples, containing 40 wt% ibuprofen, were stored for approximately 4 weeks before the release experiments, to allow the ibuprofen to crystallize.

6.2.4 On-demand Release Measurement

Release experiments were performed by placing a sample in a holder to keep the cylinder axis aligned with the magnetic field direction. Subsequently, the sample was placed in a jar containing 30 mL phosphate buffered saline. Thereafter, the jar was placed in a water bath set at the required temperature of 25 or 37 ◦C. The samples were pretreated by moving them twice for 30 minutes to a secondary water bath set at 65 ◦C. The samples were placed in the initial water bath for 60 minutes in between. Subsequently, the PBS buffer was refreshed and the samples were placed again in the initial water bath at the required temperature. Triggering the on-demand drug release from the samples was done by placing the

101 Chapter 6 sample flask for 15 minutes in a small water bath in an alternating magnetic field setup. The setup generated an magnetic field with a field strength of 2850 A m-1 and a frequency of 745 kHz and the small water bath was set at the same temperature as the initial bath. The sample flask was placed within this water bath during the experiment in order to maintain a steady buffer temperature during the experiment. The amount of ibuprofen released was measured by taking samples from the PBS solution, which were analyzed by HPLC. After sampling an equal amount of fresh buffer was added to the flask to maintain a constant volume.

6.3 Results & Discussion

6.3.1 Diﬀusion of Ibuprofen in Poly(styrene-stat-butyl methacrylate)

Apparent diffusion coefficients of ibuprofen in poly(styrene-stat-butyl methacrylate) (p(S-BMA)) at different initial ibuprofen concentrations and different temperatures have been determined. For that purpose, the release of ibuprofen from p(S-BMA) discs into phosphate buffered saline (PBS) has been measured. By plotting the -1 1/2 -1 fractional release of ibuprofen, Mt/M∞, against 4π t L , the slope is equal to the square root of the diffusion coefficient. As the diffusion coefficient depends on concentration of ibuprofen, the initial slope has been taken in order to calculate the diffusion coefficient at the initial concentration at the particular temperature, see Figures 6.1a-e. Since ibuprofen acts as a plasticizer, i.e. lowering the glass transition temperature of the polymer, and the movement of the polymer chains increases with temperature, the diffusion coefficient increases with concentration and temperature. The apparent diffusion coefficients of all samples measured above the glass transition temperature have been fitted using an exponential function of the temperature, with both the pre-exponential and growth factors depending on the initial concentration of ibuprofen. The diffusion coefficient of the sample with an ibuprofen concentration of 5 wt% measured at 20 ◦C was not included, since ◦ the Tg of this sample is approximately 30 C. The results of these fits are give by Equation 6.4: D =1.219 10-51 e1.420·C e(-0.0039·C+0.2498)·T (6.4) · · · where C is the initial concentration [wt%] and T is the temperature [K]. The diffusion coefficient can be extrapolated to 0 wt% concentration of ibuprofen for temperatures above the glass transition temperature of the polymer, see Figure 6.1f.

102 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core

0.16 0.25 20 °C 20 °C 0.14 30 °C 30 °C 40 °C 40 °C 50 °C 0.2 50 °C 0.12 60 °C 60 °C 70 °C 70 °C 0.1 0.15 [−] [−]

inf 0.08 inf / M / M t t 0.1

M 0.06 M

0.04 0.05 0.02

0 0 0 0.005 0.01 0.015 0.02 0 0.005 0.01 0.015 0.02 4 π−1 t1/2 L−1 [10−8 s1/2 m−1] 4 pi−1 t1/2 L−1 [10−8 s1/2 m−1]

(a) (b)

0.35 0.35 20 °C 20 °C 30 °C 30 °C 0.3 40 °C 0.3 40 °C 50 °C 50 °C 0.25 60 °C 0.25 60 °C 70 °C 70 °C

[−] 0.2 [−] 0.2 inf inf / M / M

t 0.15 t 0.15 M M 0.1 0.1

0.05 0.05

0 0 0 0.005 0.01 0.015 0.02 0 0.005 0.01 0.015 0.02 4 pi−1 t1/2 L−1 [10−8 s1/2 m−1] 4 pi−1 t1/2 L−1 [10−8 s1/2 m−1]

−12 0.4 10 20 °C 30 °C 0.35 ]

40 °C −1

50 °C s −14 2 0.3 60 °C 10 70 °C 0.25 [−]

inf −16 0.2 10 / M t

M 0.15 5 wt% −18 0.1 10 10 wt% 16.7 wt% 22 wt% 0.05 Diffusion coefficient [m 28.7 wt%

−20 0 wt% (extrapolated) 0 10 0 0.005 0.01 0.015 0.02 20 30 40 50 60 70 4 pi−1 t1/2 L−1 [10−8 s1/2 m−1] Temperature [°C]

(e) (f)

Figure 6.1: Diffusion of ibuprofen in p(S-BMA) at different temperatures with initial concentrations of (a) 5, (b) 10, (c) 16.7, (d) 22 and (e) 28.7 wt% and the apparent diffusion coefficient at different initial concentration.s

For temperatures below the Tg of the polymer this procedure would result in a significant overestimation of the diffusion coefficient due to the plasticizing effect of the ibuprofen since all data used in the fit was measured above the glass transition temperature.

103 Chapter 6

6.3.2 On-demand Release from Poly(styrene-stat-butyl methacrylate)

The thermoresponsive polymer, poly(styrene-stat-butyl methacrylate) (p(S-BMA)), has been coated (0.7 mm) on iron rods with a diameter of 1.1 mm. The polymer coating contained diﬀerent concentrations of ibuprofen as a model drug, i.e. 16.7, 28.6 and 40 wt%, for on-demand release using an alternating magnetic ﬁeld as external trigger.

Release from 37 ◦C

On-demand release experiments have been performed by placing a flask containing 30 mL phosphate buffered saline and the sample, vertically aligned in a holder (parallel with the magnetic field), from a water bath of 37 ◦C in an alternating magnetic field setup for 15 minutes. The flask has been put in a small water bath, kept at 37 ◦C in the magnetic field setup in order to maintain the temperature of the buffer at 37 ◦C. Upon exposure of the sample to the alternating magnetic field, the iron rod in the center is heated by induction. Consequently, the temperature of the thermoresponsive coating is increased, significantly enhancing the drug release from the sample, see Figure 6.2. By removing the sample from the magnetic field setup and placing it back into the initial 37 ◦C water bath, the induction heating stops and the temperature of the thermoresponsive coating decreases. Therefore, the diffusion coefficient of ibuprofen in the polymer decreases and hence the release of ibuprofen from the sample decreases. When the sample is again exposed to the alternating magnetic field, the release rate of ibuprofen increases, showing the possibility of repetitive on-demand release. As expected, increasing the concentration of ibuprofen in the thermoresponsive p(S-BMA) coating increases the release rate during exposure to the magnetic field, the on situation, as well as without exposure, the off situation. This is due to the increase in driving force for diffusion with concentration. However, the release rate during exposure to the magnetic field increases less with concentration compared to the release rate in the off situation, i.e. without the magnetic field. The on/off ratio decreases with increasing concentration of ibuprofen from 2180 to 530 and 200 for concentrations of 16.7, 28.6 and 40 wt%, respectively. Due to the plasticizing effect of ibuprofen, the bulk Tg of the p(S- BMA) coating containing 28.6 wt% ibuprofen ( -6 ◦C) is lower than that containing ∼ 16.7 wt% ibuprofen ( 12 ◦C). At the outer layer of the thermoresponsive coating, ∼

104 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core

16 18 16 18 ] ]

14 16 −1 14 16 −1 14 14 12 12 12 12 10 10 16.7 wt% ibuprofen 10 16.7 wt% ibuprofen 10 8 without isolation 8 with isolation 8 8 6 6 6 6 4 4 4 4

2 2 2 2 Release Rate Ibuprofen [mg h Release Rate Ibuprofen [mg h Cumulative Release Ibuprofen [mg] 0 0 Cumulative Release Ibuprofen [mg] 0 0 50 100 150 200 250 300 50 100 150 200 250 300 Time [h] Time [h]

(a) (b)

16 18 16 18 ] ]

14 16 −1 14 16 −1 14 14 12 12 12 12 10 10 28.7 wt% ibuprofen 10 28.7 wt% ibuprofen 10 8 without isolation 8 with isolation 8 8 6 6 6 6 4 4 4 4

16 18 16 18 ] ]

14 16 −1 14 16 −1 14 14 12 12 12 12 10 10 40 wt% ibuprofen 10 40 wt% ibuprofen 10 8 without isolation 8 with isolation 8 8 6 6 6 6 4 4 4 4

(e) (f)

Figure 6.2: On-demand release of ibuprofen from p(S-BMA) coated (0.7 mm) on an iron rod with an ibuprofen concentration of (a-b) 16.7, (c-d) 28.7 and (e-f) 40 wt% from 37 ◦C. Samples (b), (d) and (f) have been additionally coated with 0.6 mm p(E-VA). the concentration of ibuprofen is negligible and, therefore, the glass transition ◦ temperature is assumed to be equal to the Tg of the pure polymer (44.6 C from DSC measurements). The glass transition temperature of the pure polymer at the outer

105 Chapter 6 layer might be slightly lower than measured by DSC due to the surface effect.30 Since the off temperature of 37 ◦C is close to the glass transition temperature at the outer layer of the thermoresponsive coating, it appears that the release of ibuprofen is not only limited by the diffusion at the outer layer, but in addition, the diffusion in the bulk is also important. Similarly, the bulk diffusion is important in the on situation, as the temperature of the outer layer is well above the glass transition temperature of the pure polymer. The ratio between the bulk diffusion at high temperature (on situation) and moderate temperature (off situation) is larger at a low concentrations of ibuprofen, as can be seen by the larger slope at lower ibuprofen concentrations at Figure 6.1f. Therefore, the on/off ratio at low concentration is higher for release from 37 ◦C. Increasing the ibuprofen concentration to oversaturation (40 wt%) results in ibuprofen dissolved in the p(S-BMA) matrix as well as the presence of ibuprofen crystals in the matrix, which are visible by eye as well as detectable by DSC. By pretreatment of the samples using two cycles of release triggered by a water bath, a similar boundary layer is created as in the samples with an ibuprofen concentration below saturation (16.7 and 28.6 wt%). However, the crystals that are present close to the boundary can redissolve and provide a constant high driving force for diffusion possibly causing a higher initial off release, resulting a lower on/off ratio of the on- demand drug release. This hypothesis is supported by the fact that the off release decreases after each triggering due to depletion of the crystals close to the boundary. In addition to the thermoresponsive coating, the samples have been coated with 0.6 mm of poly(ethylene-stat-vinyl acetate) (p(E-VA)) as an isolation layer to increase the effective temperature at the outer layer of the glass transition switch. Because of the low glass transition temperature of p(E-VA), approximately -35 ◦C, the diffusion coefficient of ibuprofen is assumed to be much higher compared to that in p(S-BMA) to not decrease the release rate of the ibuprofen from the sample. Surprisingly, the results from the release experiments of isolation coated samples show similar release rates upon exposure as the nonisolated samples. This while the release rate without the presence of the magnetic field is higher in the case of the additional layer of isolation. Therefore, the on/off ratio is lower than that of the nonisolated samples, with ratios 870, 300 and 170, for samples with an initial ibuprofen concentration of 16.7, 28.6 and 40 wt% in the thermoresponsive p(S-BMA) coating, respectively. The behavior can be explained by noting that even though the diffusion coefficient of ibuprofen in p(E-VA) is several orders of magnitude larger than that in p(S-BMA), the concentration of ibuprofen in the outer layer

106 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core of the thermoresponsive p(S-BMA) coating can probably no longer be assumed to be negligible. Therefore, in the thermoresponsive switch, the glass transition temperature of the polymer is suppressed, resulting in a higher diffusion coefficient of ibuprofen and, consequently, a higher release rate in the off situation. On the other hand, because of the isolating effect of the p(E-VA) coating, the temperature reached by the outer layer of the thermoresponsive p(S-BMA) coating is expected to exceed the temperature of this outer layer when no isolation is used. The temperature of the outer layer, with and without the additional isolation layer has been modeled using Comsol. See the Appendix for the used parameters. The model shows a significant temperature difference between with and without the isolation layer, see Figure 6.3. Nonetheless, the release rate upon exposure, i.e. during heating of the sample, was found to be similar to that of the samples without isolation layer. Due to the significantly lower solubility of ibuprofen in p(E-VA) compared to that in p(S-BMA),31 the release rate might be suppressed, in particular at high temperatures, because of the significantly higher release of ibuprofen into the p(E-VA) layer. Moreover, the temperature reached far exceeds the melting temperature of the model drug, ibuprofen, possibly resulting in additional unexpected alteration of the release kinetics.

160 With isolation Without isolation

140

120

100

80 Temperature [°C] 60

0 50 100 150 200 250 Time [s]

Figure 6.3: Eﬀect of an additional isolation layer of p(E-VA) (0.6 mm) on the temperature at the outer layer of the thermoresponsive layer, modeled using Comsol. Initial temperature 37 ◦C.

107 Chapter 6

Release from 25 ◦C

In order to decrease the release rate of ibuprofen in the absence of the alternating magnetic field, experiments have been performed from 25 ◦C. For this purpose, both the initial water bath and the water bath within the magnetic field setup have been set at 25 ◦C and release experiments have been performed using samples consisting of an iron rod core, coated with a 0.7 mm thick layer of ibuprofen incorporated p(S-BMA), with an additional coating of 0.6 mm p(E-VA) as an isolation layer. Modeling the temperature increase in the case of heating from 25 ◦C in Comsol (data not shown) has shown similar results as the case of heating from 37 ◦C. Therefore, the temperature at the outer layer of the thermoresponsive p(S-BMA) coating is expected to significantly exceed the Tg of the polymer, resulting in a similar release rate upon exposure to the alternating magnetic field, see Figure 6.4. Since the lower temperature of 25 ◦C is significantly below the glass transition temperature of the pure polymer, the off release is predominantly controlled by the diffusion through the outer layer. The diffusion coefficient at 25 ◦C is lower than at 37 ◦C and, therefore, the on/off release ratios with an ibuprofen concentration of 28.6 and 40 wt% increase to 990 and 790, respectively. Since the off release is almost independent of the ibuprofen concentration in the bulk whereas in the on situation the bulk diffusion is important, the on/off ratio at 16.7 wt% ibuprofen is lower (500) than that of 28.6 wt%, due to the lower Tg of the bulk at 28.6 wt% and, hence, higher bulk diffusion. However, no explanation has been found for the lack of increase of the on/off ratio compared to the release experiment of the p(E-VA) isolated sample with 16.7 wt% ibuprofen from 37 ◦C.

6.3.3 Eﬀect of Alignment with Magnetic Field

The effect of the alignment of the sample with respect to the direction of the alternating magnetic field has been quantitatively investigated using a finite element method. The induction heating power generated in the iron rod was modeled using a three dimensional model in the AC/DC module of Comsol 3.5 at different angles of alignment of the rod with the magnetic field, assuming a constant relative 32 permeability, µr, of 1350, which is one third of µr,max of commercially available iron. The hysteresis loss has been modeled using the law of Steinmetz.33 Subsequently, the temperature increase of the outer layer of the thermoresponsive coating, i.e. the position important for on-demand release, was modeled for an exposure of 120 seconds. This exposure time is sufficient for modeling, as the sample reaches

108 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core

16 18 ]

14 16 −1 14 12 12 10 10 8 8 6 6 4 4

2 2 Release Rate Ibuprofen [mg h Cumulative Release Ibuprofen [mg] 0 0 50 100 150 200 250 300 Time [h]

(a)

16 18 ]

14 16 −1 14 12 12 10 10 8 8 6 6 4 4

2 2 Release Rate Ibuprofen [mg h Cumulative Release Ibuprofen [mg] 0 0 50 100 150 200 250 300 Time [h]

(b)

16 18 ]

14 16 −1 14 12 12 10 10 8 8 6 6 4 4

2 2 Release Rate Ibuprofen [mg h Cumulative Release Ibuprofen [mg] 0 0 50 100 150 200 250 300 Time [h]

(c)

Figure 6.4: On-demand release of ibuprofen from p(S-BMA) coated (0.7 mm) on an iron rod with an ibuprofen concentration of (a) 16.7, (b) 28.7 and (c) 40 wt% from 25 ◦C. The samples have been additionally coated with 0.6 mm p(E-VA). a steady-state temperature after 30 s. ∼ Exposure of the cylindrical iron rod parallel to the direction of the alternating

109 Chapter 6 magnetic field results in an eddy current flow in a circular plane perpendicular to the direction of the magnetic field and, therefore, perpendicular to the axis of the rod.34 Consequently, the rod is heated by the joule effect. Furthermore, as iron is ferromagnetic, hysteresis contributes to the induction heating as well. As the sample is tilted, i.e. the angle between the axis of the iron rod and the direction of the magnetic field becomes larger, the cross sectional area, perpendicular to the direction of the magnetic field, becomes larger. Therefore, the induced current density decreases, and subsequently, the generated power decreases, see Figure 6.5. Cross sectional areas close to the ends of the rod are not ovular as is the case around the center of the rod. The induced current paths are squared off, increasing the local current density and, consequently, locally the generated power. With increasing angle between the axis of the rod and the direction of the magnetic field, the fraction of the rod carrying these squared off cross sections increases, increasingly balancing the decrease of generated power. As expected, the demagnetization effect in the iron rod is negligible, due to the small diameter of the rod. Due to the significantly higher thermal conductivity of the iron rod compared the polymer coating, any inhomogeneity in the heating can be neglected.

5.5

4.5

3.5 Power [W] 3

2.5

2 0 15 30 45 60 75 90 Angle [°]

Figure 6.5: Eﬀect of the alignment of the iron rod with the direction of the magnetic ﬁeld on the power generated by induction in an iron rod, modeled using Comsol.

Using the generated power output calculated using the 3 dimensional AC/DC module of Comsol 3.5, the eﬀect of the angle between the axis of the iron rod and the magnetic ﬁeld on the temperature of the outer layer of the thermoresponsive coating has been modeled. Because of the relatively small reduction of the generated

110 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core power at small angles, the decrease in temperature of the outer layer is minor at small angles, see Figure 6.6. However, at large angles, i.e. the axis of the rod nearly perpendicular to the magnetic field, the effect of the angle is relatively large, due to a large dependence of the characteristic length for heat transfer at these angles. As the release experiments described in this work have been performed in a vertical solenoid, the iron rod is horizontal when it is perpendicular to the direction of the magnetic field.

75 0° (parallel) 70 90° (perpendicular)

65 Increasing angle 60

Temperature [°C] 45

35 0 50 100 150 200 250 Time [s]

Figure 6.6: Eﬀect of the alignment of the iron rod with the direction of the magnetic ﬁeld on the temperature at the outer layer of the thermoresponsive layer, modeled using Comsol.

Since the temperature of the outer layer of the thermoresponsive coating is important for on-demand release, it is required to align the rod with the magnetic field in order to reach the maximum temperature and, subsequently, obtain the maximum on/off ratio. As a temperature difference of 20 ◦C at the outer layer has an effect of several ∼ orders of magnitude on the diffusion coefficient, especially around the glass transition temperature the release rate of ibuprofen is significantly affected by a variation of the alignment, making the application of rods in drug delivery implants less robust.

6.4 Conclusion

On-demand release of ibuprofen as model drug has been shown using an alternating magnetic ﬁeld as external trigger from 25 and 37 ◦C. Ibuprofen has been released from a thermoresponsive polymer, based on the reversible change of the diﬀusion

111 Chapter 6 coefficient of a solute around its glass transition temperature. Using an iron rod as core material, coated with poly(styrene-stat-butyl methacrylate) as the thermoresponsive polymer, repetitive on-demand drug delivery has been shown upon exposure of the device to an alternating magnetic field, with on/off ratios exceeding 2000. The on/off ratio, the ratio between the release rate upon exposure and without the magnetic field, and the release rate itself is affected by the concentration of the incorporated ibuprofen. The dependence of the on/off ratio on the concentration is different for different start temperatures. At 25 ◦C, the off release is predominantly controlled by the diffusion coefficient at the outer layer, and the bulk diffusion only affects the on release, whereas at 37 ◦C the bulk diffusion is important in the off state as well as the on situation, due to small difference between the off temperature and the glass transition temperature of the pure polymer. An additional coating of poly(ethylene-stat-vinyl acetate) as an isolation layer has no significant effect on the release rate of ibuprofen upon exposure to the magnetic field trigger, however, it increases the release without exposure, decreasing the on/off ratio. Modeling data quantitatively shows the effect of the alignment of the iron rod with respect to the direction of the magnetic field on the inductive heating of the rod. The dependence of the heating on the angle originates from a decrease of induced current density and an increase in end effects with increasing angle. Furthermore, the subsequent effect of the angle on the temperature of the outer layer of the thermoresponsive coating, which is the switching layer for the on-demand release, has been modeled. At small angles the effect is relatively small. However, at larger angles the temperature of the sample is significantly lower than that of an aligned sample, up to 20 ◦C. This can result in a difference of several orders of magnitude ∼ for the diffusion coefficient of ibuprofen in the polymer.

112 Reversible on-demand drug release from polymeric matrices containing a cylindrical iron core

Appendix

Table 6.1: Input parameters for modeling in Comsol.

Parameter Symbol Value Unit

- Magnetic ﬁeld strength Hmax 2850 [Am ] Frequency f 745 [kHz]

Relative permeability core µr 1350 [-] Hysteresis loss33 W 0.002 (B 104)1.6 f 10-7 [W m-3] hys · · · · Sample length l 0.015 [m] Sample diameter d 0.0011 [m] Characteristic length L L sin(α) +d sin(90-α) [m] k tot · · Angle with respect to ﬁeld α 0to90 [◦] ◦ Ambient temperature Tbl 37 [ C] Power (values from AC/DC model) P 2.59 to 4.73 [W] -3 Density core ρcore 7870 [kg m ] 35 -3 Density Tg layer ρTg layer 1066 [kg m ] 35 -3 Density isolation ρisolation 1066 [kg m ] -1 -1 Thermal conductivity core kcore 80.4 [W m K ] 35 -1 -1 Thermal conductivity Tg layer kTg layer 0.19 [W m K ] 35 -1 -1 Thermal conductivity isolation kisolation 0.19 [W m K ] -1 -1 Heat capacity core Cp,core 460 [Jkg K ] 35 -1 -1 Heat capacity Tg layer Cp,Tg layer 1500 [Jkg K ] 35 -1 -1 Heat capacity isolation Cp,isolation 1500 [Jkg K ] Emissivity ǫ 0.9 [-]

For the used water properties and dimensionless numbers see the appendix of Chapter 5.

113 Chapter 6

References

[1] Richards Grayson, A.C., Choi, I.S., Tyler, B.M., Wang, P.P., Brem, H., Cima, M.J., Langer, R. Nature Materials, 2 (2003), 767-772.

[2] Qiu, L.Y., Zhub, K.J. International Journal of Pharmaceutics, 219 (2001), 151- 160.

[3] Moriyama, K., Yui, N. Journal of Controlled Release, 42 (1996), 237-248.

[4] Zhang, J., Peppas, N.A. Macromolecules, 33 (2000), 102-107.

[5] Heller, J., Trescony, P.V. Journal of Pharmaceutical Sciences, 68 (1979), 919- 921.

[6] Brannon-Peppas, L., Peppas, N.A. Biomaterials, 11 (1990), 635-644.

[7] Sershen, S.R., Westcott, S.L., Halas, N.J., West, J.L. Journal of Biomedical Materials Research Part A, 51 (2000), 293-298.

[8] Sershen, S.R., Westcott, S.L., West, J.L., Halas, H.J. Applied Physics B, 73 (2001), 379-381.

[9] Bikram, M., Gobin, A.M., Whitmire, R.E., West, J.L. Journal of Controlled Release, 123 (2007), 219-227.

[10] Kwok, C.S., Mourad, P.D., Crum, L.A., Ratner, B.D. Journal of Biomedical Materials Research Part A, 57 (2001), 151-164.

[11] Kim, H.J., Matsuda, H., Zhou, H., Honma, I. Advanced Materials, 18 (2006), 3083-3088.

[12] Aschkenasya, C., Kost, J. Journal of Controlled Release, 110 (2005), 58-66.

[13] Kost, J., Leong, K., Langer, R. Proceedings of the National Academy of Sciences of the Uniteds States of America, 86 (1989), 7663-7666.

[14] Kwon, I.C., Bae, Y.H., Kim, S.W. Nature, 354 (1991), 291-293.

[15] Santini Jr, J.T., Cima, M.J., Langer, R. Nature, 397 (1999), 335.

[16] Lee, K.Y., Peters, M.C., Anderson, K.W., Mooney, D.J. Nature, 408 (2000), 998-1000.

114 References

[17] Edelman, E.R., Kost, J., Bobeck, H., Langer, R. Journal of Biomedical Materials Research, 19 (1985), 67-83.

[18] Kost, J., Wolfrum, J., Langer, R. Journal of Biomedical Materials Research, 21 (1987), 1367-1373.

[19] Saslawski, O., Weingarten, C., Benoit, J.P., Couvreur, P. Life Sciences, 42 (1988), 1521-1528.

[20] M¨uller-Schulte, D., Schmitz-Rode, T. Journal of Magnetism and Magnetic Materials, 302 (2006), 267-271.

[21] Zhang, J.L., Srivastava, R.S., Misra, R.D.K. Langmuir, 23 (2007), 6342-6351.

[22] Liu, T.Y., Hu, S.H., Liu, D.M., Chen, S.Y., Chen, I.W. Nano Today, 4 (2009), 52-65.

[23] Csetneki, I., Filipcsei, G., Zr´ınyi, M. Macromolecules, 39 (2006), 1939-1942.

[24] Babincova, M. Pharmazie, 50 (1995), 702-703.

[25] Babincova, M., Cicmanec, P., Altanerova, V., Altaner, C., Babinec, P. Bioelec- trochemistry, 55 (2002), 17-19.

[26] Keurentjes, J.T.F., Kemmere, M.F., Bruinewoud, H., Vertommen, M.A.M.E., Rovers, S.A., Hoogenboom, R., Stemkens, L.F.S., P´eters, F.L.A.M.A., Tielen, N.J.C., van Asseldonk, D.T.A., Gabriel, A., Joosten, B., Marcus, M.A.E. Angewandte Chemie International Edition, 48 (2009), 9867-9870.

[27] Siepmann, F., Le Brun, V., Siepmann, J. Journal of Controlled Release, 115 (2006), 298-306.

[28] Fournier, R.L., Basic transport phenomena in biomedical engineering, 2nd edition, Taylor & Francis, New York, 2007.

[29] Crank, J., The mathematics of diﬀusion, Oxford University Press, Oxford, 1975.

[30] Fischer, H. Macromolecules, 38 (2005), 844-850.

[31] Bruinewoud, H., Ultrasound-Induced Drug Release from Polymer Matrices: The glass transition temperature as a thermo-responsive switch, Ph.D. thesis, Eindhoven University of Technology, 2005, ISBN: 90-386-2877-3.

115 Chapter 6

[32] Stanley, J.K., Electrical and magnetic properties of metals, American Society for Metals, Metals Park, Ohio, 1963.

[33] Bozorth, R.M., Ferromagnetism, D. van Nostrand Company, Inc., Princeton, New Jersey, 1959.

[34] Davies, E.J., Simpson, P., Induction heating handbook, McGraw-Hill, London, 1979.

[35] CRC Polymers, Polymers: A Properties Database, CRC Press, Online Edition, 2007.

116 Chapter 7

Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core

A system for multiple on-demand drug release has been prepared using an alternating magnetic field as external trigger. The core/shell samples have been composed of a macroscopic spherical iron core coated with a thermoresponsive polymer, poly(styrene-stat-butyl methacrylate), containing ibuprofen as a model drug. During exposure of the samples to the magnetic field (on situation), the release rate of ibuprofen is significantly increased, up to 35 times the release rate without the magnetic field (off situation). Using 1 sample or 2 samples in line with the magnetic field does not influence the on/off ratio of the system, showing the possibility of using multiple samples to increase the drug dose. Increasing the concentration of ibuprofen in the polymer layer is shown to increase the release rate in both the on and off situation. Increasing the size of the iron core and, consequently, decreasing the polymer thickness, only increases the release rate during exposure resulting in higher on/off ratios. Chapter 7

7.1 Introduction

Since several decades, a major challenge in the field of controlled drug delivery is the controlled delivery of an active in a pulsatile manner. This challenge has been addressed from two directions of which one is the formation of a system that releases its load at a predetermined time or in pulses with predetermined intervals.1–3 A different approach is the use of external or internal stimuli including pH,4–6 near-infrared radiation,7–9 ultrasound,10–13 electrical,14,15 mechanical16 and magnetic triggers.17–19 Internally stimulated systems have the advantage of self regulation of the drug release, however, after admission the release can not be adjusted, apart from removal of the device. Externally stimulated systems are able to release drug on-demand as the physician or patient requires. The use of a low frequency magnetic field has been reported for in vitro and in vivo on-demand release by exerting pressure on, and squeezing additional drug out of, pores in the polymer matrix by movement of a magnet in the drug delivery device.17,18 Furthermore, an alternating magnetic field has been used to remotely heat a polymer showing a lower critical solution temperature (LCST) incorporated with iron oxide nanoparticles. The alternating magnetic field heats the nanoparticles, and, subsequently, the polymer collapses when heating above its LCST.20–23 Collapse of the polymer chains above the LCST results in shrinkage of the polymer layer. However, even though the LCST is a reversible mechanism, the majority of the incorporated drug is frequently squeezed out during one single exposure to the magnetic field. Apart from this, intense heating of such systems is used to irreversibly rupture the system, triggering a burst-like drug release.21,22 Also, magnetoliposomes have been triggered for instant release using an alternating magnetic field, causing the liposomes to burst.24,25 In Chapters 5 and 6, as well as a recent paper,26 we demonstrated a reversible thermoresponsive switch based on the glass transition temperature of the polymer,

Tg, for multiple pulsatile drug release using an external trigger. The thermoresponsive switch used in these studies is based on the significant reversible increase in the diffusion coefficient of a solute in a polymer matrix with increasing 26 temperature, especially around the glass transition, Tg, of the polymer. Even though ibuprofen has a clear plasticizing effect on the used polymers,27 lowering the glass transition temperature, the rate determining diffusion coefficient is found at the outer skin layer of the thermoresponsive coating. As the ibuprofen concentration in the PBS solution is very low, sink conditions can be assumed and, apart from the partition coefficient, the concentration of ibuprofen at the outer skin of the coating is negligibly low. Therefore, the glass transition temperature of the outer skin is equal

118 Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core

to the Tg of the pure polymer and the rate determining diffusion coefficient is equal to the diffusion coefficient at negligibly low ibuprofen concentration. Previously used external triggers include near-infrared radiation, for which a near-infrared dye was incorporated in the thermoresponsive polymer.26 Since the penetration depth of near-infrared radiation is small, thereby limiting the applicability, a magnetic field has also been used to increase the temperature of the polymer. For this purpose, the thermoresponsive polymer was coated on a core of superparamagnetic iron oxide nanoparticles incorporated polymer matrix (Chapter 5) as well as on an iron rod (Chapter 6). Inductive heating of the iron rod generated substantially more heat than heating the commercial superparamagnetic iron oxide nanoparticles, resulting in enhanced drug release. However, the magnetic heating of the iron rod core depends upon the alignment with respect to the applied alternating magnetic field. To avoid this required alignment, this study focusses on spherical iron cores as the supply of thermal energy upon applying the magnetic field.

7.2 Materials & Methods

7.2.1 Materials

Spherical iron cores with a diameter of 3 mm, 4 mm, and 5 mm, intended for the use as bearing balls, were purchased from Brammer, the Netherlands. Poly(styrene- stat-butyl methacrylate) (p(S-BMA)) (MW 200,000, 50 % styrene) was purchased from Scientific Polymer Products, USA. The phosphate buffered saline (PBS) was purchased from Sigma Aldrich. The phosphate buffered saline solution contained NaCl (120 mM), KCl (2.7 mM) and phosphate buffer (10 mM) with a pH of 7.4 at 25 ◦C. Acetonitrile (HPLC - Supra gradient) and dichloromethane (DCM) (99.9+ %) were purchased from Biosolve. Trifluoracetic acid (TFA) (99.8+ %) was purchased from VWR International. Ibuprofen was purchased from Fargon.

7.2.2 Preparation of Release Coating

The thermoresponsive polymer, poly(styrene-stat-butyl methacrylate), containing ibuprofen as a model drug, was applied as a coating on macroscopic spherical cores with a diameter of 3, 4, and 5 mm. The coating was applied in a three stage process, see Figure 7.1. Firstly, the appropriate ratio of p(S-BMA) and ibuprofen was mixed in a pre-heated custom built double cone screw compounder with a volume of 5 cm3, set at 100 ◦C and 100 rpm for 15 minutes (Stage 1). Directly after compounding

119 Chapter 7

p(S-BMA) ibuprofen

Stage 1:

Compounder

Top view:

Stage 2: Side view: 3 mm press 3 mm press 4 mm press 5 mm

Press hemispheres

Polymer

Bearing core

Stage 3:

Press 2 hemispheres around core

Figure 7.1: Preparation of thermoresponsive drug delivery device of p(S-BMA) containing ibuprofen, coated around a macroscopic spherical iron core in a three stage process. small samples were cut from the strand and added to a hand press to form polymer hemispheres with a diameter of 6 mm with a hemispherical hole of either 3, 4 or 5 mm (Stage 2). The hand press was heated to 80 ◦C for approximately 5 minutes in an oven. In the third stage, one iron core was put in the opening of two hemispheres with the corresponding size hole. Subsequently, the sample was put in another hand press and heated in an oven in order to fuze the two polymer hemispheres together around the core at a temperature of 80 ◦C. As the outer diameter of all the samples was the same, the thickness of the polymer coating was 1.5, 1 and 0.5 mm for

120 Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core the samples with a core of 3, 4 and 5 mm, respectively. Supersaturated samples, containing 40 wt% ibuprofen, were stored for approximately 4 weeks before the release experiments to allow the ibuprofen to crystallize.

7.2.3 Release Measurement

Release experiments were performed by placing 1 sample or 2 samples in a holder in 30 mL of phosphate buffered saline. When 2 samples were used, the holder insured that the samples were kept above each other, in line with the magnetic field. Subsequently, the samples were placed in a water bath set at 25 ◦C and pretreated by moving them twice for 30 minutes to a secondary water bath set at 65 ◦C. The samples were placed in the initial water bath for 60 minutes in between. Thereafter, the PBS buffer was refreshed and the samples were placed again in the initial water bath at 25 ◦C. Triggering the on-demand drug release from the samples was done by placing the sample flask for 15 minutes in a small water bath in an alternating magnetic field setup. The setup generated a magnetic field with a field strength of 2850 A m-1 and a frequency of 745 kHz, see Chapter 2 for details. The small water bath within the setup was set at the same temperature as the initial bath. The sample flask was placed within this water bath during the experiment in order to maintain a steady buffer temperature during the experiment. The amount of release ibuprofen was measured by taking samples from the PBS solution, which were analyzed using a high performance liquid chromatograph (HPLC). An equal amount of fresh buffer was added to the flask after taking a sample to maintain a constant volume. The HPLC analysis was done on a Shimadzu VP Series HPLC using a solution of 50/50 vol% acetonitrile and water with 0.05 vol% trifluoracetic acid as the mobile phase, with a flow of 1.2 mL min-1 (LC-10AD VPA solvent pump). A reversed phase C18 column (Discovery HS, 150x4.6mm-3 µm) was used a the stationary phase and an UV-spectrophotometer (Shimadzu, SPD- 10Avp) was used to measure the ibuprofen content at a wavelength of 223 nm.

7.3 Results & Discussion

Thermoresponsive core-shell samples have been investigated for on-demand drug release using an alternating magnetic ﬁeld as external trigger. The core consists of a macroscopic spherical iron ball and the shell of poly(styrene-stat-butyl methacrylate) containing ibuprofen as a model drug. Representative examples of the macroscopic

121 Chapter 7 core-shell particles are shown in Figure 7.2. Samples in which the p(S-BMA) is supersaturated with ibuprofen (40 wt%)26 are opaque, see Figure 7.2c, and show a melting peak of ibuprofen around 80 ◦C in diﬀerential scanning calorimetry (data not shown).

(a)

(b)

(c)

Figure 7.2: Thermoresponsive core-shell samples of a 3 mm macroscopic spherical iron core and shell of p(S-BMA) containing (a) 16.7, (b) 28.6 and (c) 40 wt% ibuprofen.

7.3.1 Magnetically Triggered Drug Release from Macro- scopic Core-shell Particles

On-demand release experiments have been performed by placing a flask containing a holder with either 1 sample or 2 samples in line and 30 mL phosphate buffered saline from a water bath of 25 ◦C in an alternating magnetic field setup for 15 minutes. The buffer temperature was kept constant at 25 ◦C using a small water bath in the magnetic field setup.

122 Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core

Upon exposure of the core-shell sample to the alternating magnetic field, the macroscopic spherical iron core is heated by induction heating. Subsequently, thermal energy, generated in the core, heats the thermoresponsive polymer coating. Therefore, the temperature of the outer layer of the polymer coating exceeds the glass transition temperature, which significantly increases the release rate of ibuprofen from the polymer, see Figure 7.3. After the alternating magnetic field is removed, the heating of core is stopped and the temperature of the outer layer decreases. Hence the release rate of ibuprofen decreases again. Multiple exposures to the magnetic field during the following days result in similar release rates, showing that the samples can be used for repetitive drug release.

7.3.2 Eﬀect of Alignment: 1 Sample vs 2 Samples

Using multiple spheres should increase the release dose upon triggering compared to a single sphere, because of an increase of surface area available for release, allowing simple fine tuning of the drug dosing. In order to investigate whether the release is indeed doubled when 2 samples are placed simultaneously in the magnetic field, release from 1 sample and 2 samples of poly(styrene-stat-butyl methacrylate) containing 28.6 wt% ibuprofen around a 4 mm iron core has been triggered. In both experiments the release from the samples is significantly increased when exposed to the alternating magnetic field, see Figure 7.3, due to the induction heating of the iron core and the subsequent heating of the coating. As expected, the release rate, the slope of the cumulative release graph, for the experiment with 2 samples is larger than for the experiment with 1 sample, as the surface area for release is twice as much. In fact, normalizing the release rate to the surface area results in nearly identical release rates, see Figure 7.3b. Consequently, the on/off ratios, i.e. the ratio between the release rate upon exposure and without the magnetic field, are identical, with 36 and 34 for 1 sample and 2 samples in line, respectively. Therefore, it can be concluded that placing 2 samples in line parallel to the magnetic field lines does not influence the on-demand release from the individual thermoresponsive drug delivery core-shell particles and, hence, the temperature increase of the core of the particles by induction heating is not influenced. Consequently, using multiple samples together will increase the effective surface area for release and subsequently, the release rate of ibuprofen.

123 Chapter 7

350 1 sample 2 samples in line 300

250

200

150

100

50 Cumulative Release Ibuprofen [µg] 0 40 60 80 100 120 140 Time [h]

(a)

] 40

−2 1 sample 35 2 samples in line cm −1 30

Release Rate Ibuprofen [µg h 0 40 60 80 100 120 140 Time [h]

(b)

Figure 7.3: On-demand release from 1 and 2 core-shell samples of a p(S-BMA) shell containing 28.6 wt% ibuprofen with a 4 mm macroscopic spherical iron core, with (a) the cumulative release of ibuprofen and (b) the release rate normalized to the release surface.

7.3.3 Eﬀect of Ibuprofen Concentration & Core Size

A variation of the concentration of ibuprofen in the thermoresponsive coating will influence the driving force for diffusion of ibuprofen and, consequently, the release from the system. Therefore, samples have been prepared with different ibuprofen concentrations of 16.7, 28.6 and 40 wt% ibuprofen around cores of 3 and 4 mm. The

124 Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core on-demand release has been triggered for 1 sample and 2 samples in line. As shown previously, the drug release rate is signiﬁcantly increased when the samples are exposed to the alternating magnetic ﬁeld, see Figure 7.4. As expected, increasing

] 100 ] 100

−2 1 sample −2 1 sample 2 samples in line 2 samples in line cm cm

−1 80 3 mm −1 80 4 mm 16.7 wt% ibuprofen 16.7 wt% ibuprofen

60 60

40 40

20 20

Release Rate Ibuprofen [µg h 0 Release Rate Ibuprofen [µg h 0 0 50 100 150 200 250 300 350 0 50 100 150 200 250 300 350 Time [h] Time [h]

(a) (b)

] 100 ] 100

−2 1 sample −2 1 sample 2 samples in line 2 samples in line cm cm

−1 80 3 mm −1 80 4 mm 28.6 wt% ibuprofen 28.6 wt% ibuprofen

60 60

40 40

20 20

Release Rate Ibuprofen [µg h 0 Release Rate Ibuprofen [µg h 0 0 50 100 150 200 250 300 350 0 50 100 150 200 250 300 350 Time [h] Time [h]

] 100 ] 100

−2 1 sample −2 1 sample 2 samples in line 2 samples in line cm cm

−1 80 3 mm −1 80 4 mm 40 wt% ibuprofen 40 wt% ibuprofen

60 60

40 40

20 20

Release Rate Ibuprofen [µg h 0 Release Rate Ibuprofen [µg h 0 0 50 100 150 200 250 300 350 0 50 100 150 200 250 300 350 Time [h] Time [h]

(e) (f)

Figure 7.4: On-demand release from 1 and 2 core-shell samples of a p(S-BMA) shell containing diﬀerent concentrations of ibuprofen ((a) and (b): 16.7 wt%, (c) and (d): 28.6 wt% and (e) and (f): 40 wt%) and diﬀerent sizes of iron cores ((a), (c) and (e): 3 mm and (b), (d) and (f): 4 mm).

125 Chapter 7 the concentration of ibuprofen in the p(S-BMA) coating increases the release rate of ibuprofen in both the on and off state. However, in case of the undersaturated samples, i.e. 16.7 and 28.6 wt% ibuprofen, the release rate during exposure to the magnetic field appears to increase more with concentration compared to the release rate in the off situation, i.e. without the magnetic field. The on/off ratio increases from 7 to 9 and from 15 to 35 for the 3 and 4 mm cores, respectively, when increasing the ibuprofen concentration. Due to the plasticizing effect of the ibuprofen, the bulk Tg of the p(S-BMA) coating containing 28.6 wt% ibuprofen is lower than that containing 16.7 wt% ibuprofen. In the off situation, the temperature of the outer layer of the thermoresponsive coating is below the Tg of the pure polymer. Since the ibuprofen concentration at the outer layer is negligible, the release rate is almost independent of the ibuprofen concentration in the bulk of the polymer. However, in the on situation, the temperature of the outer layer of the coating is above the

Tg of the pure polymer. Therefore, the release of ibuprofen from the polymer is not only limited by the diffusion at the outer layer, but in addition, the diffusion in the bulk is important. Due to the lower Tg of the bulk at 28.6 wt% ibuprofen, the bulk diffusion is higher, resulting in a higher on release and on/off ratio. Increasing the ibuprofen concentration to supersaturation (40 wt%) results in ibuprofen dissolved in the p(S-BMA) matrix and the presence of ibuprofen crystals in the matrix, which are visible by eye (Figure 7.2c), as well as an ibuprofen melting peak in differential scanning calorimetry. Even though the release rate from these supersaturated samples is higher than that of the lower concentrations in both the on and off situation, it is surprising that the on/off ratio is lower, with 6 and 20 for the 3 and 4 mm cores, respectively. This may be caused by the fact that by pretreatment of the samples using two cycles of release triggered in a water bath, a boundary layer is created with a negligible concentration of ibuprofen. Hence, the glass transition of the polymer matrix in this boundary layer is equal to that of the pure polymer. Consequently, the diffusion coefficient of ibuprofen in this layer is rate limiting in the off state, as is the case with lower concentration of ibuprofen in the overall coating. The initial higher off release might be ascribed to the presence of ibuprofen crystals close to the boundary, which redissolve and provide a high driving force for diffusion. This hypothesis is also supported by the fact that the off release decreases after each triggering, because of depletion of the crystals. The experiments show a significant increase in on/off ratio with increasing size of the iron core. As expected, the size of the core does not influence the release rate without exposure to the magnetic field since the outer surface area of the

126 Repetitive on-demand drug release from polymeric matrices containing a macroscopic spherical iron core sample is equal. However, increasing the core size significantly increases the release rate upon exposure to the field. Although the induction heating of the 5 mm core generates most thermal energy, the main contribution to the increase in the on/off ratio is thought to originate from the considerably thinner coating layer, since the temperature of the outer layer of the coating is determining the on-demand release. When the distance of this layer to the heating source, i.e. the iron core, is smaller because of an increasing core size, the outer layer will become warmer, resulting in a higher release rate. In fact, increasing the core size to 5 mm with 40 wt% ibuprofen increases the on/off ratio to 32 compared to 6 and 20 for the cores of 3 and 4 mm, respectively, see Figure 7.5.

] 160

−2 3 mm 140 4 mm cm 5 mm −1 120

100

Release Rate Ibuprofen [µg h 0 0 20 40 60 80 100 120 140 Time [h]

Figure 7.5: On-demand release from 2 core-shell samples of a p(S-BMA) shell containing 40 wt% of ibuprofen and iron cores of 3, 4 and 5 mm.

7.4 Conclusion

A drug delivery system has been prepared which can be triggered multiple times for on-demand drug release using an alternating magnetic field as external trigger. The system exists of a thermoresponsive polymer, poly(styrene-stat-butyl methacrylate), containing ibuprofen as a model drug, coated on a macroscopic spherical iron core. When a sample is exposed to the alternating magnetic field, the iron core is heated due to induction heating and, subsequently, the polymer coating around the core is heated. As the diffusion coefficient of ibuprofen in the thermoresponsive layer increases with temperature, particularly around the glass transition of the polymer,

127 Chapter 7 the release rate from the system is significantly increased during exposure to the magnetic field. The main advantage of this system over previously reported devices with iron rods is that no alignment of the implant in the magnetic field is required. Using 1 sample or 2 samples in line with the magnetic field does not influence the release rate of ibuprofen, normalized to the surface area for release and, therefore, the on/off ratios of both situations are equal. As the surface area is increased when using multiple samples, this can be used to control the drug dose. Increasing the concentration of ibuprofen increases the release rate of ibuprofen during exposure as well as without the magnetic field. However, the on/off ratio is the largest at a concentration approaching the saturation limit of ibuprofen in the polymer (28.6 wt% on/off = 35). As the outer diameter of the samples has been held constant, using a larger size iron core decreases the thickness of the polymer coating. Subsequently, the distance of the outer layer of the coating to the heated core is decreased, resulting in a higher temperature of the outer layer and, therefore, a higher release rate during exposure. Since the distance from the core does not influence the release rate without the magnetic field, this results directly in a higher on/off ratio.

128 References