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
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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
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
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 nano-particles 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 N´eel and Brown relaxation. N´eel 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.
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 N´eel 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), N´eel relaxation is the only relaxation process that contributes to the heating of the particles in ferrofluid, because of the significantly shorter relaxation time for N´eel 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
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 ex-ternally triggered using an alternating magnetic field. Due to their biocompatibility and the absence of an alignment effect, superparamagnetic iron oxide nanoparticles
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.
Extern aangestuurde medicijnafgifte uit een implantaat kan de effici¨entie van de behandeling sterk verbeteren omdat de pati¨ent of arts daarmee direct de medicijntoediening aan kan passen aan de behoefte van de pati¨ent en het medicijn alleen vrijkomt op de gewenste locatie in het lichaam. Daardoor zal de pati¨ent 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 tem-peratuurgevoelig 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 diffusieco¨effici¨ent 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 diffusieco¨effici¨ent 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 ge-bruik 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 N´eel en Brown relaxatie. Bij N´eel relaxatie wordt
het magnetisch moment van het deeltje gericht naar het magnetisch veld waardoor warmte wordt gegenereerd door het overschrijden van een anisotropie-barri`ere. 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 N´eel 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 N´eel 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 N´eel 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
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 commerci¨ele 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 methacry-laat) (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 diffusieco¨effici¨ent 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 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.
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 Preprogrammed Drug Delivery . . . 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 Magnetism in Medicine . . . 7
1.2.1 Magnetic Imaging & Spectroscopy . . . 7
1.2.2 Hyperthermia . . . 11
1.2.3 Magnetic Separation . . . 14
1.2.4 Magnetic Drug Targeting . . . 16
1.2.5 Magnetic Drug Release . . . 18
1.2.6 Conclusion . . . 21
1.3 Thesis Outline . . . 21
References . . . 23
2 Characterization and magnetic heating of commercial superpara-magnetic iron oxide nanoparticles 35 2.1 Introduction . . . 36
2.2 Materials & Methods . . . 36
2.2.1 Materials . . . 36
2.2.2 Magnetic Field . . . 36
2.2.4 Temperature Measurements . . . 39
2.3 Results & Discussion . . . 39
2.3.1 Characterization of the Magnetic Field . . . 39
2.3.2 Characterization of the Particles . . . 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 Temperature Measurements . . . 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 Temperature Measurements . . . 72
4.3 Results & Discussion . . . 73
4.3.1 Characterization . . . 73
4.3.2 Temperature Measurements . . . 74
4.4 Conclusion . . . 76
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 Release Measurements . . . 82
5.3 Results & Discussion . . . 83
5.3.1 Effect of Iron Oxide Concentration . . . 83
5.3.2 Effect of Ibuprofen Concentration . . . 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 Diffusion Measurements . . . 100
6.2.3 Preparation of Release Coating . . . 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
7.2.2 Preparation of Release Coating . . . 119
7.2.3 Release Measurement . . . 121
7.3 Results & Discussion . . . 121
7.3.1 Magnetically Triggered Drug Release from Macroscopic Core-shell Particles . . . 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 Repetitive On-demand Drug Release . . . 143
8.3.1 Temperature Control in Magnetically Triggered Thermo-responsive Drug Delivery Systems . . . 143
8.3.2 Design Criteria . . . 153
8.3.3 In Vivo Application . . . 156
References . . . 160
Dankwoord 166
List of Publications 169
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 adminis-tration is a very patient friendly, easy, relatively inexpensive adminisadminis-tration 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.
Time
Drug plasma level
Minimum effective level Toxic level
Dose Dose Dose
Dose of
sustained delivery
Conventional dosing Sustained delivery dosing
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
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.16,17By 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
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,
Temperature
Specific Volume
Core volume
Interstitial volume Hole free volume
T g V g V c
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
Time Magnetic field ON OFF Time Drug level
Minimum effective level Toxic level
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), N´eel relaxation and Brown relaxation.59–61N´eel relaxation
is the reorientation of the magnetic moment within the particle and Brown relaxation the reorientation of the magnetic particle itself. Currently, N´eel and Brown relaxation
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
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.70,71As 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 y x z y x Mz (a) z y x z y x Mxy Mxy (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
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
T1-(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.87–89Therefore,
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
a
b
c
.
NAA -1 -2 -3 -4 -5 -6 -7 PPMFigure 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
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, N´eel 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 N´eel and Brown relaxation. N´eel 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
should be minimized. Therefore, the amount of dissipated heat per gram of iron
oxide, the specific absorption rate (SAR, [W g-1
iron 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
are reported having a large SAR around 600 W g-1
iron oxide (11 kA m-1, 410 kHz).61
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
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.123,124The 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.
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 X Flow direction F drag (-x) F gravity (-y) Field/gradient direction F magnetic (+x) F drag (+y) + F buoyant (+y) (a) Y X gradient Sample inlet Sheath 2 Sheath 1 direction collection bins (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
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.138These
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
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
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.146Furthermore,
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
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
Figure 1.9: Schematic representation of on-demand drug release using a drug incor-porated 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.
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
