Magnetic targeting and ultrasound mediated drug delivery: Benefits, limitations and combination

Abstract

This paper reviews the uses of magnetism and ultrasound in therapeutic delivery applications. Emphasis is placed upon magnetic nanoparticles and microbubble ultrasound contrast agents. The underlying physical principles, history, key developments and limitations of these techniques for drug and gene delivery in vitro and in vivo are explored. The combination of ultrasonic and magnetic techniques is also reviewed with particular focus on magnetic microbubbles as delivery agents with the potential to combine the advantages of both methods whilst addressing many of their limitations. Finally, results are presented from a study of a new magnetic microbubble formulation which shows great applicability as a therapeutic delivery vehicle.

• Drug delivery
• magnetic targeting
• ultrasound
• microbubbles
• transfection
• nanoparticles

Introduction

The rapid advances in the development of new types of pharmaceutical product have greatly increased demand for more effective delivery systems [1]. In addition, new formulations and delivery methods for existing products are increasingly being sought to mitigate the impact of patent expiration, the so called “patent cliff” [2]. In 2008 the US demand for drug delivery systems was predicted to increase by more than 10% per year reaching $132 billion by 2012 [3]. Conventional delivery methods such as injections and oral administration have significant advantages in terms of convenience and cost, but are unsuitable for several classes of therapeutic compound. For example, administration of poorly soluble drugs and large molecules such as proteins, including DNA, via intravenous or oral routes is largely ineffective [1]. Moreover, systemic administration results in the exposure of all tissues and this can lead to harmful side effects [4]. The most well known examples are perhaps those relating to cancer chemotherapy and compounds such as Paclitaxel, which inhibits the proliferation of cancer cells but is often associated with neutropenia and nausea [5].

A successful drug delivery system should provide maximum therapeutic efficacy with minimal unwanted side effects [6]. To do so it must: (i) protect the therapeutic material to prevent unwanted degradation during its passage through the body, (ii) concentrate the material at the desired location(s) and (iii) promote uptake in the target tissue. The concept of targeted drug delivery was originally put forward in the literature by Paul Erlich (1854–1915) who theorised that if malignant tissue could be selectively targeted then a toxin could also be selectively delivered in a manner akin to a “magic bullet” [7]. The central principle of drug targeting is that the distribution of drug within an organism is manipulated in such a manner that the majority interacts exclusively with the target tissue [8].

The mechanism by which the drug is released at the target site can be active or passive. Passive drug delivery involves diffusion from the drug carrier into the surrounding medium via natural metabolic processes. It takes advantage of human physiology or immune system response to maintain the drug at a given location within the body. The rate of release may be controlled, for example, through appropriate use of a biodegradable polymer to encapsulate the drug [3]. Active drug targeting involves triggering or stimulating the release of the therapeutic compound by exploiting some characteristic of the carrier. This provides much more control over drug distribution and as such is more widely applicable. The stimulus can be the overexpression of receptors on tumour cells or a physical stimulus such as temperature, pH, light, pressure/sound, electric or magnetic fields [3], [9]. Localised drug release must be followed by uptake in the target tissue for successful treatment. For example, for gene delivery to be effective, the gene must reach the cell nucleus in order to be expressed [10]. Uptake may again be achieved passively, e.g. through normal membrane transport processes, or promoted by application of the same or different physical stimuli.

Numerous strategies, from light sensitive polymeric carriers to viruses [3], [11], have been proposed for therapeutic delivery and the development of delivery systems represents a significant research field in its own right. This aim of this paper is to examine two of these strategies, namely ultrasound mediated and magnetically targeted drug delivery, and the combination of both techniques.

Magnetic targeting utilises an externally applied magnetic field to control the location of a magnetically responsive drug carrier in vivo, thereby avoiding many of the disadvantages of systemic administration [12], [13]. As will be explained in the next section, however, the types of magnetic particle suitable for use in vivo cannot be successfully retained in some parts of the body due to their small size [14]. Ultrasound, widely used in imaging on account of its low cost and high patient acceptability, is finding increasing application in therapy. Its ability to be focused into very small volumes provides an excellent means of localising a therapeutic effect [15]. In addition there is strong evidence that exposure to ultrasound at moderate intensities produces temporary and reversible permeabilisation of cell membranes [16]. This “sonoporation” can be significantly amplified by the presence of microbubbles [17], [18]. Originally microbubbles stabilised by a polymeric or surfactant coating were developed as contrast agents for ultrasound imaging, but they have since been intensively investigated as vehicles for drug delivery. In this respect they have a significant advantage over magnetic particles since they can carry much larger payloads. Currently, however, successful methods for concentrating microbubbles at a target site in vivo are lacking and the last part of this paper will examine recent work in the development of magnetic microbubbles and new insights into a magnetic microbubble formulation which has shown excellent transfection efficiency.

Magnetism and drug delivery

The first documented medical application of magnets dates back to the 17th century when magnets were used to remove fragments of iron from the body [19]. In the present day, in addition to their now widespread use in magnetic resonance imaging (MRI), magnetic fields are utilised in several therapeutic applications. This section examines their use in magnetic drug targeting.

Types of magnetic materials

When a magnetic material is placed within a magnetic field the individual atomic moments of the material contribute to the overall response or magnetic induction. The response of a given material to a magnetic field varies with temperature and atomic structure. This can be classified in terms of magnetic volumetric susceptibility, χ [4]. Paramagnetic materials (χ > 0) magnetise very weakly in the direction of the magnetic field whereas diamagnetic materials (χ < 0) magnetise in the opposite direction [20]. Due to the disordered structure of their magnetic moments, neither of these types of material show spontaneous magnetisation in the absence of an external field. Other materials exhibit internal ordered states which remain ordered after the removal of the magnetic field and these include ferri, ferro or antiferromagnets, where the prefix refers to the nature of the coupling interaction between the electrons in the material.

Due to remanent magnetisation, ferri and ferromagnetic materials tend to aggregate. This remanent magnetisation occurs because once a ferro or ferrimagnetic material is magnetically saturated in one direction, the moments tend to preserve the so-induced orientation to a certain extent, depending on the internal structure. This can be seen in a hysteresis curve. The resulting aggregation is undesirable in most biomedical applications.

Some of the interest in using superparamagnetic particles in biomedical applications is due to the fact that they do not exhibit remanent magnetisation. The internal structure of bulk magnetic materials is divided into magnetic domains, grouping moments with the same orientation. Below a certain size, however, formation of domains becomes energetically unfavourable, resulting in particles consisting of a single domain state where all the magnetic moments of the particle are aligned parallel to each other [21]. For single domain particles below a certain critical size – often regarded as the superparamagnetic limit – the magnetic moment of the particle is free to fluctuate in response to thermal energy in the absence of a magnetic field. Therefore these single domain nanoparticles enter into a superparamagnetic regime. Superparamagnetic materials thus provide a compromise between strong magnetisation and avoidance of particle aggregation [20].

Magnetic nanoparticles

As a consequence of the phenomena discussed above, superparamagnetic nanoparticles have become the subject of intensive research in recent years. They have three key additional features making them attractive for biomedical applications: their ability to be visualised under MRI [12], to be localised at a target site by means of a magnetic field gradient [4] and to produce localised heating when exposed to an oscillating magnetic field for hyperthermia and/or to promote drug delivery [22–24].

Ferrimagnetic iron oxides, maghemite and magnetite, which are able to form superparamagnetic iron oxide nanoparticles (SPIONs) typically below 15 nm, are currently in use as MRI contrast agents [13]. Iron oxides are relatively safe compared to other substances used as MRI contrast agents such as gadolinium [25]. The nanoscale dimensions of the particles allow them to pass not only through the narrowest blood vessel but even through cell membranes [26]. To avoid uptake by the reticuloendothelial system (RES), magnetic carriers typically have two structural configurations; a magnetic core coated by a biocompatible polymer or magnetic nanoparticles inside a porous biocompatible polymer matrix into which drugs, genes or targeting agents can also be incorporated [4], [20], [27]. According to Neuberger et al. magnetic particles smaller than 4 µm are removed by RES cells mainly in the liver and spleen [7], [28]. In general, however, the larger the particle the shorter its plasma half-life [20].

Drug delivery applications

The use of magnetic particles in drug delivery appeared in the literature in 1960 when Freeman et al. proposed that magnetic nanoparticles could be transported through the vasculature and concentrated in a specific part of the body with the aid of a magnetic field [19]. In 1963 Meyers described how magnetic iron particles in the lymphatic and vascular system of a dog could be used as a contrast and isotopic agent [29], [30]. Zimmerman and Pilwat used magnetic erythrocytes (red blood cells) for the delivery of cytotoxic drugs in 1976 [31]. In the 1970s Senyei et al. produced albumin microspheres, incorporating superparamagnetic magnetite, and successfully retained these against in vivo capillary flow conditions using a magnetic field [32], [33]. Widder et al. used magnetic albumin microspheres containing doxorubicin and showed tumour regression in rats when a localised magnetic field was applied and with no occurrences of embolism observed [34].

Novel drug delivery systems have also been developed using polymer matrices containing drugs and magnetic beads. This pulsatile delivery system opens pores to release the drug when an oscillating magnetic field is applied [1], [35]. In 1994 Hafeli et al. prepared biodegradable poly(lactic acid) microspheres that incorporated magnetite and a beta emitter for targeted radiotherapy and successfully applied it to subcutaneous tumours [36], [37]. A ferrofluid (fluid containing a magnetic suspension, in this case magnetic nanoparticles) comprising drugs was used successfully in an in vivo mouse model by Lubbe et al. [38] as an anti-tumour agent with no toxic effects observed. Another successful trial concentrated the ferrofluid Chemicell, bound to mitoxantrone, in murine tumours with the application of a magnetic field [39]. The first phase I clinical trial was performed by Lubbe et al. Each patient tolerated the ferrofluid and the nanoparticles were seen to be concentrated at tumours after the application of a magnetic field [9]. A second phase I/II trial was conducted by a start-up company, FeRx Inc, in 1997 where Carbon coated iron particles (0.5–5 µm) which had adsorbed doxorubicin were localised and retained in target sites [40]. A full review of recent magnetic drug delivery trials in patients may be found in McBain et al. [13].

Gene delivery applications

Magnetic nanoparticles also hold great potential for gene therapy. DNA delivery to cells can induce production of therapeutic proteins to treat malignancies. Use of short interfering RNA can temporarily down-regulate the expression of a gene of choice (gene silencing) and potentially any disease causing gene can be silenced [41]. If simply injected or ingested, however, DNA can degrade and is not able to pass through cellular barriers [42]. There are three primary ways to deliver genes into cells: through the use of viral vectors which is highly efficient but which can have unwanted side effects; electroporation, in which cells are electrically stimulated to promote uptake, which is again very efficient but causes a large number of cell deaths; and non-viral transfection agents which assist nucleic acid entry into cells through other mechanisms and maintain cell viability but tend to be inefficient [25].

The first study of magnetically targeted DNA delivery was performed by Mah et al. in 2000 [43]. A recombinant adeno-associated virus (rAAV), encoding for a green fluorescent protein was coated onto the surface of magnetic microspheres. Under the influence of an applied magnetic field, 1% of the vector-microsphere formulation gave the same transfection efficiency as 100% free vector. This was continued by Hughes et al. to produce “infectious, paramagnetic, retroviral vector particles” using paramagnetic streptavidin particles conjugated to antibodies which the retrovirus associated with [44].

Non-viral vectors are also being developed with an emphasis upon improving nucleic acid transfection efficiency. The association of genes with magnetic nanoparticles for magnetically guided nucleic acid delivery has been termed magnetofection [45]. The principle is similar to that of other transfection agents, such as lipofectine 2000, except plasmid DNA forms complexes with a cationic polymer – usually polyethylenimine (PEI) – which is able to bind and condense DNA coated magnetic nanoparticles to create a “magnetoplex” with magnetically controllable behaviour [46]. The mechanism by which the gene enters the cell with magnetoplexes has so far been found to be no different to that associated with normal transfection reagents [14], [45], [47].

Other novel approaches include that adopted by Cai et al. who used nickel embedded carbon nanotubes coated with DNA to act as “spears” to pierce the cells resulting in high transfection efficiency [48]. Oz Biosciences has designed four different cationic nanoparticle formulations designed for specific applications in regard to gene delivery [49]. Improvements are also being looked into through the use of pulsating fields where Kamau et al. showed an increase in the number of transfected cell relative to static fields [50].

Limitations

Magnetic drug targeting is subject to two main constraints. First, there may be rapid clearance of particulate drug carriers from the circulation. Second, the strength and gradient of the magnetic field that needs to be applied, given the small size of the nanoparticles and hence their reduced magnetic strength, may be very high for successful localisation of a carrier [28] against the hydrodynamic forces produced e.g. by blood flow [4]. Hence there is a demand for magnetic particles with larger diameters but that are still capable of travelling through capillaries without posing a risk of embolism [14]. Set against this, however, is the reduced ability of larger particles to penetrate tissue and cell membranes.

Ultrasound and drug delivery

One of the first investigations into the therapeutic use of ultrasound for drug delivery was performed in the 1950s by Fellinger and Schmid who used ultrasound to enhance the permeation of ointment into inflamed tissue, successfully treating polyarthritis [51]. The first ultrasound assisted transfection in vitro was reported in 1987 when Fechheimer et al. reported low level transfection of murine fibroblasts when exposed to 20 kHz ultrasound from a sonicator [52]. This was followed by the use of ultrasound to assist drug uptake by tumours in mouse models in vivo [53].

Basic physical principles

Ultrasound can be defined as “acoustic waves propagating within a medium at frequencies exceeding the auditory band” [54]. In diagnostic imaging, it is the scattering of these waves from regions of differing density and compressibility which is used to produce the images. In addition to being scattered, however, ultrasound undergoes three other types of interaction with tissue and it is these which underlie its effect in therapeutic applications. First, a fraction of the energy in the propagating wave is also absorbed by tissue, resulting in local heating. If the intensity is sufficient, tissue ablation will occur through protein denaturation and/or boiling [15], [55], [56]. At lower intensities, effects such as increased perfusion and/or stimulated release from drug delivery vehicles may be realised [55]. Second, ultrasound may produce cavitation. Acoustic cavitation is defined as the formation and activity of gas and/or vapour filled bubbles, of natural or artificial origin, in a medium exposed to sound [57]. The motion of the bubbles can temporarily increase tissue and individual cell membrane permeability [58], termed “sonoporation,” including temporary opening of the blood brain barrier [59]. At higher intensities it can contribute to cell death through a variety of mechanisms associated with the violent collapse of the bubbles. Third, transfer of momentum from the ultrasound wave to the liquid produces streaming motions and associated shear stresses on tissue interfaces. In the presence of a bubble oscillating under ultrasound excitation, microstreaming [60] occurs through a similar mechanism and this may in fact be one of the main contributors to sonoporation [61], [62].

Microbubbles

Microbubbles are widely used in ultrasound imaging as contrast agents [63], [64]. They consist of a gas core surrounded by a coating normally of a lipid, protein or polymer. The coating reduces surface tension and inhibits gas diffusion, thereby increasing the lifetime of the microbubble [65]. The high compressibility of the microbubbles compared with the surrounding tissue gives rise to a large acoustic impedance difference and causes the bubbles to undergo volumetric oscillations producing nonlinear acoustic radiation at moderate ultrasound intensities [64]. It is these phenomena which provide the contrast enhancement between regions with and without bubbles in ultrasound images [66], [67]. Microbubbles are also finding application in therapeutic applications, however. They can be used as nuclei to increase cavitation activity [57] and as vehicles for drug delivery/gene therapy by incorporation of therapeutic material into their coatings which can subsequently be released by destroying the microbubbles with a high amplitude ultrasound pulse [68], [69].

Drug delivery applications

Ultrasound has been shown to increase the thrombolytic efficacy of urokinase [70]. This is most likely due to localised cavitation which can weaken the clot surface and/or improve penetration of the drug; and this process has been shown to be enhanced by the presence of microbubbles [71], [72]. Similarly, exposure to ultrasound in the presence of microbubbles has been shown to reversibly produce pores on the surfaces of cells [6], [73]. Unger et al. obtained similar results with microbubbles containing a drug filled layer of soy bean oil and the drug paclitaxel [69]. Guzmen et al. quantified molecular uptake of celcein in prostate cancer and aortic smooth muscle cells, examining the heterogeneity of cavitation and factors that influence drug uptake [74], [75]. Microbubbles have also been used with cells for uptake of Doxorubicin and colloidal particles.[76], [77]

Targeting

Albumin and lipid coated microbubbles have been found to persist in areas of inflammation. This is due to leukocytes, from an inflammatory response, attaching to the microbubble [78]. Attaching targeting agents to the surface of microbubbles should allow them to be retained at other malignant sites within the body and as such, microbubbles with monoclonal antibodies attached have been formulated [79]. The radiation force from a propagating ultrasound wave can translate a microbubble and assist in targeting [55]. It has also been shown that ultrasound radiation force is able to deposit nanoparticles bound to microbubbles against the side of a capillary tube under flow conditions [80].

Gene delivery applications

In 1996 Kim et al. used ultrasound to transfect a variety of cells with plasmid DNA. They found that the ultrasound power, duration, ambient pressure and the occurrence of cavitation were critical and all positively correlated with transfection efficiency [81]. The same authors also injected albumin microbubbles with plasmid DNA into rat knees and observed transfection under ultrasound [81]. Greenleaf et al., continuing the work by Kim et al., demonstrated an increase in the transfection rate of DNA in vitro with ultrasound in the presence of Albumin bubbles [82]. Lawrie et al. reported that exposure to ultrasound in the presence of Optison® achieved approximately 300 fold higher transgene expression in vitro than plasmid DNA alone [83]. However low transfection efficiency still remains a major limitation and the same mechanisms that cause drug uptake can also lead to cell death [84]. The physical parameters affecting gene delivery with ultrasound and microbubbles were investigated by Rahim et al. who determined optimal combinations of pulse amplitude, exposure time and bubble concentration [85]. Matsumura's group has administered gelatin coated microspheres binding a gene to rat livers with the application of ultrasound [86]. Transfection with microbubbles and ultrasound has been performed in the lungs and hearts of mice and the hearts of rats [17], [87], [88]. Comparisons between different ultrasound contrast agents have been performed in vivo to find which gave the best cellular uptake of genes [84], [89].

Conjugation strategies

Microbubbles can be used to improve cell permeation if the drug is co-administered at the target location avoiding the problems of attachment, although this limits efficacy to malignancies reached by injection [86]. As the gas core of a microbubble cannot hold organic compounds and the lipid coating is thought to be too thin, therapeutic compounds must be attached to the shell [16]. Electrostatic interactions allow charged drugs to be bound to the surface of cationic lipid coated microbubbles. DNA is a polyanion, due to the negative sugar phosphate groups on the molecule, and as such easily binds to cationic microbubbles. Drugs can also be incorporated into microbubbles in a layer of oil beneath the lipid outer layer and this can further stabilise the outer surface of microbubbles [86]. Lum et al. bound polystyrene nanoparticles to the microbubble shell through an avidin-biotin “bridge” which could be replaced by nanoparticles loaded with genes or drugs [80]. Liposomes have also been loaded onto microbubbles [90] and Cochrane et al. produced Paclitaxel and Doxurubicin loaded polymer microbubbles which cavitate under ultrasound forming polymer nanoparticles which enter cells and slowly release drugs [91].

Limitations

Close proximity between microbubbles and target cells is required for therapeutic effects to be realised at moderate ultrasound intensities and without unwanted damage [85]. The concentration of microbubbles at a target site needs to be optimised as adequate numbers are needed both for sonoporation and to achieve the required dose of a drug; but increased microbubble concentration poses a potential risk of embolism and may also produce excessive ultrasound scattering or “shielding” preventing exposure of the target tissue. Currently, repeatable, efficient bio-chemical targeting of microbubbles is still proving to be a challenge in vivo [92]. Surface architecture designed to maximise ligand exposure to the target tissue also runs the risk of boosting presentation of an immunogenic compound that leads to early particle clearance or worse a hypersensitivity response [16].

Ultrasound and magnetism

As described above, magnetic and ultrasonic drug delivery have both been examined and show great potential but both suffer from certain limitations. These limitations can to some extent be addressed by combining the two techniques and the remainder of this review will examine developments in this field.

High intensity focused ultrasound surgery

High intensity focused ultrasound (HIFU) and MRI, were combined in 1991 with the latter used for the purposes of treatment monitoring. Since then, both animal experiments and human trials have been performed. A commercial device has been created and FDA approved, with further clinical systems currently undergoing trials [93]. In 1993, a lipid coated microbubble incorporating paramagnetic particles was patented for use as an ultrasound and MRI contrast agent. It was also proposed as an agent for mediating therapeutic effects through cavitation nucleation [94]. More recently Smith et al. showed that magnetite nanoparticle agglomerates could be used to promote cavitation in HIFU treatments which could be used for drug release [95].

Stimulated release

Ultrasound has also been used to improve drug delivery from magnetic particles. Saravan et al. found that the application of ultrasound enhanced the release kinetics of gelatin magnetic microspheres [96]. Kaminski et al. (2008) showed that ultrasound increased enzyme release from magnetic microcarriers, but temperature rises within the carrier could denature the enzyme [97].

Magnetic microbubbles

Magnetic nanoparticles can be incorporated into microbubbles in order to form magnetic microbubbles. These have been synthesised for use in ultrasound and magnetically mediated drug delivery and as dual purpose MRI and ultrasound contrast agents.

Magnetic microbubbles as dual contrast agents

Even in the absence of magnetic nanoparticles, the gas-liquid interface in microbubbles induces large local magnetic susceptibility differences. This was demonstrated experimentally in vivo with Optison® microbubbles in rat liver [98]. This then prompted in 2008 the use of emulsion stabilised particles of varying size with inclusions of superparamagnetic iron oxide nanoparticles as dual modality (MRI and Ultrasound) contrast agents. The SPION loaded bubbles were shown to improve the contrast of both ultrasound and MR images [99]. Yang et al. improved upon this by preparing a larger microbubble construct loaded with both gas and SPIONs [100]. Park et al. also produced stable microbubbles to which could be attached numerous nanoparticles, magnetic and non-magnetic, for use in image contrast enhancement [101].

Magnetic microbubbles for drug delivery

In 2000, Soetanto et al. attached magnetic micro-particles to microbubbles to enable remote manipulation of the microbubbles for both imaging and therapeutic applications. The microbubbles were formed from charged stearates and the microparticles were also coated in charged stearate and a calcium ion was used to bind them to the microbubble. The magnetic microbubbles responded to a magnet with a 2 kg equivalent force [102], [103]. This technique has not undergone any further testing according to the literature, however and it is unknown what effect ions in the blood would have upon the electrostatic interaction between the microbubbles and the microparticles. The large size of the microbubble-microparticle aggregates would also be of concern for in vivo applications.

The use of a different type of magnetic microbubble for ultrasound mediated delivery was published by Stride et al. in 2009. Various magnetic bubble preparations, the main formulation is shown in Figure 1, were synthesised and their efficacy as transfection reagents under the influence of a magnetic field and/or ultrasound was tested in vitro in chinese hamster ovary cells. Each bubble formulation was co-injected with naked plasmid DNA encoding for luciferase. Exposure of the bubble formulations to either ultrasound or a magnetic field alone had very little effect but a combination of ultrasound and a magnetic field showed a significant increase in transfection as measured using bioluminescence imaging as shown in Figure 2. A mixture of non-magnetic microbubbles and magnetic micelles (containing magnetic nanoparticles but no gas) in fact showed slightly higher transfection efficiency upon exposure to both ultrasound and a magnetic field. It was hypothesised that this could be due to aggregation of the micelles and microbubbles or a sensitisation of the cell surface by the magnetic particles perhaps related to the mechanisms responsible for magnetofection. These hypotheses are discussed further later [104]. A similar experiment was subsequently conducted in a preliminary in vivo study in a mouse model with magnetic microbubbles and plasmid DNA again co-injected. A small magnet was positioned above right lung and both in vivo and ex vivo bioluminescence imaging demonstrated that the transfection was confined to this organ and only observed with exposure to both ultrasound and the magnetic field [105].

Figure 1. Theoretical structure of a magnetic microbubble based upon the formulation by Stride et al. The gas core is surrounded by a ferrofluid and stabilized by an outer coating of L-α-phosphatidylcholine. The average size is about 2 µm.

Figure 2. Bioluminescence imaging data for transfection enhancement of Chinese hamster ovary cells by magnetic microbubbles under different exposure conditions (101). A = non-magnetic microbubbles; B = magnetic micelles; C(i) = magnetic microbubbles. US = ultrasound; MF = magnetic field. Up/down refers to the surface on of the exposure chamber on which the cells were located and hence whether cell/bubble contact would be promoted by buoyancy.

In 2010, Vlaskou et al. published a study on magnetically and acoustically active lipospheres [42], [106]. Lipospheres were synthesised based on a technique by Unger et al. who produced liposheres with soy bean oil incorporating paclitaxel [69]. The lipospheres are synthesised by shaking a mixture of soybean oil, a cationic lipid, magnetic nanoparticles, DNA, aqueous buffer and a perfluoropropane gas in a sealed vial, optical and fluorescence images can be seen in Figure 3. The lipospheres showed lower contrast to tissue ratios than for ultrasound contrast agents that have been approved for clinical use unlike those developed by Stride et al.

Figure 3. (A) Phase contrast and (B) fluorescence (measured at a wavelength of 490/509 nm) microscopy images of the Magnetic and acoustically active liposomes composed of fluidMAG-Tween-60 magnetic nanoparticles and the luciferase plasmid pBLuc, which was fluorescently labelled with the intercalated dye YOYO-1. Reproduced with kind permission from John Wiley & Sons from Figure 2 of reference [106] © 2010 WILEY-VDH Verlag GmbH & Co. KGaA, Weinheim.

As therapeutic delivery agents the lipospheres did achieve nucleic acid delivery comparable to commercially available transfection reagents when a magnetic field was applied. When ultrasound and a magnetic field were applied, however, there was no improvement in transfection efficiency in vitro. The reason was thought to be due to the soybean oil reducing the acoustic responsiveness of the microbubbles. Using a skin fold chamber on a mouse, injections of lipospheres with ultrasound and a magnetic field did achieve a large deposition of fluorescent material, but no direct observation or quantitative data were obtained for the transfection [106].

The lipospheres were then substituted by Vlaskou et al. for lipid shelled microbubbles conjugated with positively charged magnetic nanoparticles through electrostatic interactions. These microbubbles had a much greater response to ultrasound and the ultrasound contrast to tissue ratio of the magnetic microbubbles was found to be higher than Sonovue and a combination of ultrasound and magnetic exposure improved transfection efficiency in vitro [107].

New magnetic micelle/microbubbles formulation

To understand why the physical mixture of magnetic micelles and non-magnetic microbubbles produced the highest transfection efficiency in the previous in vitro study, a series of experiments was devised. First, the microbubble/micelle mixture was injected into a 200 µm capillary tube at a flow rate of ∼1 mm/s and observed both optically and acoustically under pulse echo imaging. The majority of the particles were seen to move towards a 1.5 T FeNdB permanent magnet positioned below the capillary. This implied there had been some association between the micelles and the bubbles to produce magnetic microbubbles. To study the interaction further, magnetic micelles incorporating the fluorophore boron-dipyrromethene (BoDIPY) C₁₀ were synthesised and mixed with non-fluorescent non-magnetic microbubbles. This was performed by half filling a syringe (1 ml) with magnetic micelles, then filling the rest with microbubbles. As the microbubbles float, they pass through the micelle mixture. The syringe was then inverted three times and the mixture used directly. The microbubbles were then observed to fluoresce as if surrounded by the micelles as shown in Figure 4. The microbubbles still fluoresced after being centrifuged, indicating a strong conjugation or aggregation and apparently producing a higher yield of magnetically responsive microbubbles than the direct fabrication method. It is this effective increase in concentration that is thought to have been responsible for the improved transfection performance in the previous experiments.

Figure 4. Magnetic microbubbles prepared by mixing non-magnetic microbubbles with fluorescently labelled magnetic micelles. (A) retention against flow in a capillary tube following application of a magnetic field (B) fluorescent microbubble obtained by mixing non-fluorescent microbubbles with fluorescent magnetic micelles.

Conclusions and outlook

Magnetic nanoparticles have great potential for drug delivery showing tolerance and efficacy in human trials. Magnetofection also has been shown to be a safe efficient method of gene therapy. As the nanoparticles can be controlled with an external magnetic field they present excellent targeting capabilities. However retention against blood flow at a target site within the body is a significant challenge on account of their small size.

Microbubbles, already used in ultrasound contrast enhancement, have been shown to be effective drug delivery vehicles. Microbubbles under ultrasound excitation can promote sonoporation reversibly increasing the permeability of nearby cells and tissue membranes and thereby, increasing drug uptake. Microbubbles have also been formulated to carry therapeutic agents attached to the shell or in an inner solvent layer. However, efficient targeting of microbubbles in vivo is still limited.

The advantages of magnetic and ultrasound drug delivery have been combined in a number of ways. In particular, combining magnetic nanoparticles and microbubbles to form magnetic microbubbles has been shown to produce a controllable, versatile delivery vehicle that can be visualised by MRI and ultrasound. They have been successfully used for localised transfection in vitro and promising results from in vivo studies have recently been obtained.

In terms of application, magnetic microbubbles could, in principle, be used across the range of applications for which both non-magnetic microbubbles and magnetic microparticles have been proposed as described above. The most immediate application would perhaps be for targeted delivery where magnetic localisation could be used to increase the dwell time of microbubbles in a target volume, whilst specificity could be provided by biochemical targeting. Passive extravasation and/or tissue penetration would be limited for bubbles in the micrometre size range. As has been discussed earlier, however, exposure to both ultrasound and magnetic fields can temporarily increase tissue permeability and uptake. It would be possible, moreover, both to load microbubbles with smaller therapeutic particles which could be released following microbubble disruption and to produce magnetically and acoustically responsive particles that were large compared with conventional nanoparticles but smaller than 1 µm. Perfluorocarbon droplets, which are initially a few hundred nanometres in diameter but which expand following ultrasound exposure have for example have been utilised for contrast enhancement [55].

The maximum depth at which magnetic localisation can be achieved remains an open question and depends on numerous factors including the nature of the particles being confined, the generation of the magnetic field and potential safety considerations and the physiological environment e.g. shear stresses experienced within a vessel. As described above, successful localisation of magnetic microparticles has been demonstrated in vivo including at tissue depths greater than 10 cm [108] and work is currently underway by the authors to quantify the conditions under which localisation of magnetic microbubbles can be achieved.

Acknowledgements

The authors would like to thank the University of Oxford, University College London and the ESPRC for supporting this research and The Royal Institution of Great Britain for use of the live cell imager. They would also like to thank Dr Daniel Ortega and Cristina Blanco-Andujar for technical support.

Declaration of interest: The authors report no conflicts of interest. The authors alone are responsible for the content and writing of the paper.