Focused Ultrasound (FUS) is promoting the deposition of energy inside the human body in a non-invasive way.¹  (http://www.fusfoundation.org/for-researchers/mechanisms-of-action). Focused ultrasound energy can be deposited in tissues and lesions with a diameter as little as 1 mm providing a substantial advantage for induction of heat.^2,3  This in turn has a lot of advantages for drug targeting due to the locally increased temperature. When ultrasound is applied in biological systems it can induce local tissue heating, cavitation, and radiation force, which can be used to initiate local (focal) drug delivery, increased molecule permeation through membranes and enhanced diffusivity of drugs, only at the site of sonication therefore allowing control of local drug delivery.⁴  Delivery of certain therapeutics largely benefits from local drug delivery. Cytotoxics, immune suppressive drugs and certain biologicals would have an improved therapeutic index when administered locally with limited exposure to the healthy tissues. Smart drug delivery forms that encapsulate the drug can be designed to deliver their cargo in tissues with increased temperature. FUS (or HIFU: high intensity focused ultrasound) induced local hyperthermia can be the trigger of drug release from thermosensitive carriers.⁵ 

The ability of FUS waves to induce thermal or mechanical effects at a defined location in living tissue was first described in 1942, when Lynn et al. tested FUS in the brain.⁶  In the 1950s the Fry brothers developed a clinical device to treat patients with Parkinson disease. They used an ultrasound system in combination with X-rays (imaging) to determine the target location relative to skull and to focus the ultrasound beam through a craniotomy into the deep brain for functional neurosurgery.⁷  Later on, in the 1980’s the first FDA-approved FUS system, Sonocare CST-100, was developed to treat ocular disorders such as glaucoma and many patients were clinically treated with this system.⁸  More recently technological developments have delivered new FUS equipment coupled most of the time with an imaging device such as diagnostic ultrasound and/or MRI. Current research and development aim to design and develop novel transducer technology and array designs to achieve rapid delivery of focal sonication, to improve transducer accessibility (smaller devices) or devices to conform/fit to certain parts of the body such as a helmet of arrays for brain focal treatment of diseases.^9,10 

Several FUS devices are currently in clinical practice either for approved treatments or for research purposes. These devices are combined with either ultrasound (US) imaging or MR imaging for guidance and thermometry.¹¹  Insightec manufactures the ExAblate2000® which uses MRI for extracorporeal treatment of uterine fibroids (FDA approved) with significant success, and extensive current research focuses on investigating its application in other parts of the body.^12,13  Recently the FDA approved the transcranial MR-guided focused ultrasound for the treatment of essential tremor.¹⁴  This is a hemispheric phased-array transducer (ExAblate Neuro; InSightec Ltd, Tirat Carmel, Israel) with each element directed separately, providing individual correction of skull distortion as well as electronic guidance. The Ablatherm HIFU/US consists of a transrectal probe for prostate treatment and has CE mark approval. In this case imaging is performed with ultrasound.¹⁵  The Sonablate 500, an ultrasound guided system, uses a transrectal FUS probe to carry out prostate cancer focal ablation surgery.¹⁶  The Sonalleve HIFU/MR is an MR compatible device developed to examine applications such as fibroids and bone metastases.¹⁷  The device has been used for the treatment of neuropathic pain and essential tremor and there is also promise of possible application for brain tumours.^18–20  Essential tremor non-interventional functional neurosurgery treatment has shown an immense potential of transcranial MRgFUS application to induce lesions focally and treat patients non-surgically.^21,22  MRgFUS devices are constantly developed. Currently there are 32 manufacturers of image guided FUS worldwide that develop equipment for ablative treatments. There is an on-going interest to use these devices in combination with novel formulations for image guided targeted drug delivery.

Ultrasound propagates as mechanical vibrations that make molecules within their medium oscillate around their positions and in the direction of the waves’ propagation. As a result the molecules form compressions and rarefactions that propagate the wave. The ultrasound energy is attenuated exponentially through the tissue.²³  The rate of energy flow through a unit area, in the direction of the wave propagation, is called acoustic intensity. At 1 MHz the ultrasound wave is attenuated approximately 50% and it propagates through 7 cm of tissue. The attenuated energy is then transformed into temperature elevation; heat in the tissue.^24,25 

Ultrasound waves are transmitted from one soft tissue to another adjacent tissue. In soft tissues a small amount of waves are reflected back. However, in the soft tissue–bone interface almost one-third of the incident energy is reflected back. In addition, the amplitude attenuation coefficient of ultrasound waves is 10–20 times higher in bone than in the soft tissues. This causes the transmitted ultrasound beam to be absorbed rapidly within the bone, leading to high temperature increase.²⁶  When the ultrasound beams are focused at one point, a focal diameter of 1 mm can be achieved at 1.5 MHz. The length of the focus is 5–20 times larger than the diameter (cigarette or rice shaped focus point). If the ultrasound beam is transmitted from an applicator 2–3 cm in diameter, the ultrasound intensity at the millimetre-sized focal spot can be several hundred times higher than in the overlying tissues. The ultrasound exposure drops off rapidly across the area within the sonication path and focusing helps overcoming attenuation losses and to concentrate energy deep in the body avoiding the surrounding tissues.²⁷ 

The fact that focusing of ultrasound energy is so defined is important for the design of smart materials that can respond or transform to temperature change. Hyperthermia applied in a focused way could affect the phase transition of these materials and induce release of the therapeutic payload only at the heated site. Focussing can help the release of the drug only at the desired site.²⁸ 

Focusing of ultrasound energy is significantly improved with the use of transducer arrays that are driven by signals having the necessary phase difference to obtain a common focal point.²⁹  The advantage of phased transducer arrays is that the focal spot can be guided and controlled. In addition, the electronically focussed beam allows multiple focal points to be induced simultaneously or fast electronic scanning of the focal spot which increases the area of the focal region. This feature allows shorter treatment time.^30,31  Therefore when focused ultrasound is coupled with imaging and theranostics it could target treatments in areas inside the body independent of the size or location.

During the last 15 years emphasis has been given to nanosized carriers for cancer therapy. Nanomedicine is a topic that investigates the effects of nanotechnology in healthcare and has introduced a series of novel drug delivery systems, among which are multifunctional chemical entities called nanotheranostics.^32,33  These are designed to simultaneously detect, image and treat tumours due to the fact that nanoparticles preferably accumulate in tumours compared to other tissues. If these nanoparticles can provide contrast enhancement and deliver their therapeutic cargo then these nanoparticles act as theranostics.³⁴  These systems can be engineered using biocompatible and biodegradable materials and nanomaterials, or by “labelling” previously developed nanoparticles.³⁵  The recent advances of imaging modalities enable nanotheranostics to bind onto lesions or biomarkers on specific cells opening the doors to personalised cancer therapy. The concept of “scan and treat” at the same time can provide clinicians the opportunity to adapt treatments according to imaging data in real time. Nanotheranostics can be used to detect and treat metastases.^36,37  The ability of these drug carrying nanoparticles to reach small metastatic lesions and detect them can offer a substantial advantage to image guided treatments. Nanotheranostics were introduced in the field of nanomedicine research about a decade ago, however their clinical potential is yet to be seen.³⁸  Nanotheranostics can be specially designed to respond to focused ultrasound and an imaging modality.

The combination of high-intensity focused ultrasound together with high-resolution MR guidance has created a platform that can produce focussing of the ultrasound energy deep within solid organs without invasive steps. Being non-invasive, focused ultrasound gives several advantages to modify the temperature of deep tissues and in combination with smart temperature triggered nanotheranostics to create a powerful tool for killing cancer cells. Accurate targeting and detailed accurate thermal mapping are provided by MRI and provide accurate deposition of energy in tissues that can be altered in response to near real-time thermal imaging produced by MRI, so that the variation in tissue response that is otherwise observed can be avoided.^39,40 

In recent years two clinical imaging modalities have been combined with FUS to provide guidance within the targeted area in the tissues, and this is mainly used to ablate lesions.⁴⁰  Ultrasound-guided FUS was the first image-guided system, however MR-guided HIFU or MRgFUS has been suggested as it shows several inherent advantages, such as superior anatomic detail and real-time thermometry during the thermoablation process, and it has demonstrated positive results in the ablation of both benign and malignant tumours.⁴¹  FUS has been employed in the cases of hepatocellular carcinoma, prostate cancer, uterine myomas, and breast tumours, and has shown success in palliative pain management in pancreatic cancer and bone tumours.⁴⁰  Ultrasound and MRI are widely used clinical imaging modalities that have been combined with FUS for Image guided FUS treatments. Ultrasound microbubbles for ultrasound imaging or contrast enhancing agents for MRI imaging provide the required contrast enhancement for the application of FUS, and although gadolinium based contrast enhancement agents have been found to interfere with MRI thermometry, this may be dependent on concentration and type of agent used.⁴²  On the other hand, paramagnetic liposomes have been suggested as indicators of temperature based on release of their content at 42 °C. These liposomes are loaded with Gd³⁺ contrast enhancing agents and their release at this temperature indicates a method of confirming focal hyperthermia (Figure 1.1).⁴³ 

MRI contrast enhancing agents have been suggested as a method to assess drug release from thermosensitive liposomes. In a recent study thermosensitive liposomes were doubly loaded with iron oxide nanoparticles and Gd-chelate. At low temperatures, the transverse relaxivity of the liposomes was high, allowing detection of TSLs in tissues. This is important due to the fact that these theranostics need to be imaged in tissues before intervention with FUS. After temperature increase and thermal liposomal membrane instability the longitudinal relaxivity steeply increased indicating release of the Gd-chelate contents. By choosing the appropriate MR sequences, availability and release could be assessed without interference of one contrast agent with the other.⁴⁴ 

In a separate study Peller et al. introduced the combination of therapeutic (doxorubicin thermosensitive liposomes, DOX-TSL) and imaging (contrast enhancement thermosensitive, CA-TSL) to confirm drug release in the tumour.⁴⁵ 

When using nanocarriers sensitive to mechanical forces (similar to the oscillating ultrasound pressure waves) and/or sensitive to temperature, ultrasound can induce either cavitation and/or hyperthermia and the cargo can be released locally. Although these are two different concepts and have different suggested mechanisms their interplay and combination has also been considered.⁴⁶ 

The most widely tested concept of combining ultrasound with triggered release is the use of MRgFUS with thermosensitive liposomes (Figure 1.2). Thermosensitive liposomes have been suggested for local drug release in combination with local hyperthermia more than 25 years ago.⁴⁷ 

The use of microbubbles may be designed specifically to enhance cavitation effects. Cavitation can lead to sonoporation of cell membranes and transfer of therapeutics and or their carriers.^48,49  Real-time imaging methods, such as magnetic resonance, optical and ultrasound imaging have led to novel insights and methods for ultrasound triggered drug delivery. Image guidance of ultrasound can be used for: (a) targeted tissue identification; (b) spatio-temporal guidance of energy that translates to heat to release or activate the drugs and/or permeabilise membranes; (c) imaging can assist evaluation of bio-distribution, pharmacokinetics and pharmacodynamics; and (d) physiological read-outs to evaluate the therapeutic efficacy.

Image guided focused ultrasound targeted drug delivery requires the following: (a) a clinically approved imaging modality; (b) a contrast enhancing agent for this imaging modality; (c) a drug delivery system that responds or coordinates with the energy of the focused ultrasound deposited at the site of action; (d) feedback from the targeted tissue usually provided as a temperature map or a change in the contrast enhancement; (e) the right focused ultrasound equipment for the respective targeted part of the body. If the carrier is a nanosized liposomal system that responds to temperature the attributes and mechanism of drug release should be similar to the one shown in Figure 1.2.

Ideally the drug delivery system should be non-toxic and inert to other tissues and not act by delivering the drug anywhere other than the targeted site. The drug delivery system can also provide feedback of the localisation of the pathology and response to the treatment of the focused ultrasound.

Thermosensitive liposomal carriers were first developed almost four decades ago. The concept of a temperature modified liposomal membrane had triggered scientists to use liposomes as models to understand lipid membranes behaviour at different temperatures. Yatvin et al. first described the effect of increased temperature (hyperthermia) on liposomes in 1978.⁵⁰  However, development of temperature sensitive liposomes (TSL) for therapeutic reasons, such as cancer treatment, was first introduced in 1999 by Needham’s group, who suggested that lipid phase transition enhanced lipid membrane permeability.⁵¹  The team performed in vivo data using cancer mice models. The authors described a new advanced lipid formulation containing the drug doxorubicin with thermosensitive properties. The formulation was tuned to be stable at 37 °C, but became porous at mild hyperthermic temperatures (39 °C to 40 °C) achievable in the clinic, leading to very rapid and sharp doxorubicin release. This new TSL formulation, in combination with hyperthermia (42 °C), was found to be significantly more efficient in reducing tumour growth in a human squamous cell carcinoma xenograft than free drug or previous liposome formulations, indicating that thermally triggered release was advantageous.⁵²  These low temperature-sensitive liposomes (LTSL) were later tested in dogs having canine tumours and showed a tumour growth efficiency indicating that the concept of thermally triggered release is an attractive mechanism for targeted drug delivery of chemotherapeutics.^53,54  A formulation made by these thermosensitive liposomes took the brand name Thermodox® and was later developed by Celsion corporation, an organisation that currently leads the field of FUS thermally triggered doxorubicin drug release. Thermodox® or LTSL liposomes have been used by several groups in combination with induced hyperthermia. They can be triggered to release their pharmaceutical payload by any heat-based treatment such as radiofrequency thermal ablation (RFA), microwave hyperthermia (MWH), and HIFU. Results from a Phase I study using Thermodox® were recently published.⁵⁵  Researchers used escalating dose of Thermodox® and concluded that Thermodox® can be safely administered at 50 mg m⁻² in combination with RFA. Thermodox® in combination with RFA has been tested in a large Phase I study to treat hepatocellular carcinoma.⁵⁶  Analysis of the results from this study indicated the imaging features following RFA with LTLD were different from those after standard RFA, the increased size of the LTLD-treated ablation zone after RFA indicated that the doxorubicin drug-induced effects in the treated tissues were ongoing.⁵⁷ 

There is a clear shift in Celsion’s research pipeline to use HIFU as a more efficient and targeted approach of locally increasing the temperature.

The concept of using liposomes and FUS was introduced recently, in 2006 when Frenkel et al. used liposomal doxorubicin (Doxil®) in combination with pulsed HIFU exposures in a murine breast cancer tumour model. Doxil® is a stable liposomal preparation that has no response to increased temperature.⁵⁸  This was developed to minimise doxorubicin’s cardiotoxicity, by encapsulating doxorubicin within stealth liposomes. Although Doxil® achieves a long circulation of doxorubicin with minimum cardiotoxicity it does not rapidly release the drug within the tumour, instead it facilitates the drug’s accumulation in tumours due to EPR effect and allows the drug to be slowly diffused from the liposomes in the tumour. In another study, pulsed-HIFU exposures were not found to enhance the therapeutic delivery of doxorubicin and did not induce tumour regression. However, a fluorescent dextran showed blood vessels to be dilated as a result of the exposures.⁵⁹  This indicates that hyperthermia induces vasodilation and helps extravasation of nanoparticles and drugs. In that study experiments with polystyrene nanoparticles of similar size to the liposomes showed a greater abundance to be present in the treated tumours.⁵⁹  Although this study did not achieve or prove a therapeutic advantage of the use of FUS with chemotherapeutic liposomes, it showed clearly that pulsed HIFU has the potential of improving drug distribution in the tumour and induces a substantial increase of permeation of macromolecules (fluorescent dextran) and nanoparticles (fluorescent microspheres) through the tumour blood vessels.

In 2007 Dromi et al. presented the first study on thermosensitive liposomes LTSL and HIFU hyperthermia. The authors investigated pulsed-HIFU as a source of heat in combination with thermosensitive liposomes to enhance delivery and therapeutic efficacy of doxorubicin in murine adenocarcinoma tumours. In vitro treatments simulating the pulsed-HIFU thermal dose (42 °C for 2 min) triggered drug release of 50% of doxorubicin from the thermosensitive liposomes; however, no detectable release from the non-temperature sensitive liposomes (similar to Doxil®) was observed, as expected. Similarly, in vivo experiments showed that pulsed-HIFU exposures combined with the LTSL-Dox resulted in a more rapid release of doxorubicin, as well as significantly higher concentration within the tumour when compared with LTSL-Dox alone or the non-thermosensitive liposomes (doxil type), with or without exposures.⁶⁰  This study showed that FUS can induced focal hyperthermia that affects the phase transition of lipid bilayers of thermosensitive liposomes.

The same team developed MR imageable thermosensitive liposomes (iLTSL), with the objective to characterise drug release in phantoms and in vivo. An MRI contrast agent (ProHance® Gd-HP-DO3A) and doxorubicin were loaded in liposomes and drug release was quantified by spectroscopic and fluorescence techniques, respectively. The drug release after application of FUS under MR guidance was examined in tissue-mimicking phantoms containing iLTSL and in a VX2 rabbit tumour model usually used in interventional radiology, iLTSLs demonstrated consistent size and doxorubicin release kinetics. Release of doxorubicin and ProHance® from iLTSL was minimal at 37 °C, but faster when heated to 41.3 °C. The MRI relaxivity of iLTSL increased significantly from 1.95 +/− 0.05 to 4.01 +/− 0.1 mMs⁻¹ when liposomes were heated above the phase transition temperature, indicating the release of ProHance® from liposomes and its exposure to the aqueous surroundings (enhanced MRI contrast). Importantly, the signal increase corresponded spatially and temporally to MRgFUS-heated locations in the phantoms used in this experiment. In vivo, the investigators confirmed MRI signal after administration of iLTSL injection and after each 10 min heating, with the greatest increase in the heated tumour region. The authors concluded that MRgFUS combined with iLTSL may enable real-time monitoring and spatial control of drug release from liposomes.⁶¹  This is important as it provides control of the administration and offers the opportunity for the development of personalised treatments.

In a follow up study the same authors investigated the effect of iLTSL in rabbits bearing VX2 tumours. In that study image-guided non-invasive hyperthermia was applied for a total of 30 min, completed within 1 hour after LTSL infusion and quantified doxorubicin release in tumours with HPLC and fluorescence microscopy. Sonication of VX2 tumours resulted in accurate and spatially homogenous temperature control in the target region. LTSL + MR-HIFU resulted in significantly higher tumour doxorubicin concentrations (3.4-fold greater compared with LTSL respectively, p < 0.05, Newman-Keuls). The authors observed that both free doxorubicin and LTSL treatments appeared to deliver more drug in the tumour periphery as compared to the tumour core. This indicated that FUS induced hyperthermia and LTSL increases the permeability of doxorubicin as doxorubicin was found in both the tumour periphery and core.⁶²  The group further developed a heating algorithm using the same rabbit tumour model proving that the use of the binary feedback algorithm results in accurate and homogenous heating within the targeted area.⁶³  A computational model that simulated the tissue heating with FUS treatments and the resulting hyperthermia in tissues (similar to the one that leads to drug release) were developed by Haemmerich et al.⁶⁴  In this model a spatio-temporal multicompartmental pharmacokinetic model simulated the drug release in the blood vessels and its transport into the interstitial space, as well as cell uptake. Two heating schedules were simulated each lasting 30 min, the first corresponding to hyperthermia (HT; 43 °C) and the second corresponding to HT followed by a high temperature (50 °C) for a 20s pulse, (HT+). Using the computational model (validated in rabbit VX2 tumours) the authors found that the cellular drug uptake is directly related to hyperthermia duration. However, HT+ enhanced drug delivery by 40% compared to HT.⁶⁴  The study indicated the importance of simulations in the application of drug delivery mechanisms to tumours and that understanding hyperthermia effects can support the design of MRgFUS treatments for patients.

In addition to the progress in the understanding of the physical mechanism of drug delivery from well validated thermosensitive liposomes carrying doxorubicin, researchers further investigated the chemical composition of such liposomes in response to MRgFUS induced hyperthermia.

De Smet et al. compared thermosensitive liposomes carrying doxorubicin and ProHance® (gadolinium based contrast enhancing agent). Two temperature-sensitive systems composed of the following lipids DPPC : MPPC : DPPE-PEG²⁰⁰⁰ (LTSL) and DPPC : HSPC : cholesterol : DPPE-PEG²⁰⁰⁰ (traditional temperature-sensitive liposomes, TTSL) were investigated for their stability and release profile at 37 °C and 42 °C in phantoms using MRI [lipids; 1,2-dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), 1-palmitoyl-sn-glycero-3-phosphocholine (MPPC), 1,2-dipalmitoyl-sn-glycero-3-phosphoethanolamine-N-[methoxy(polyethyleneglycol)-2000](DPPE-PEG²⁰⁰⁰), hydrogenated-l-α-phosphatidylcholine (HSPC)]. The LTSL system showed a higher leakage of doxorubicin at 37 °C, but a faster release of doxorubicin at 42 °C compared to the TTSL system, indicating that lipid composition plays an important role in the stability and release profile.⁶⁵  The authors further investigated the more stable traditional temperature sensitive liposomes carrying doxorubicin and ProHance® in vivo in rats bearing 9L gliosarcoma tumours. A clinical MRgFUS system was applied in a proof-of-concept study to induce local hyperthermia for 30 min. The local temperature-triggered release of ProHance® was monitored with interleaved T₁ mapping of the tumour. A good correlation between the ΔR₁, (change in longitudinal relaxation rate ΔR₁ = Δ(1/T₁)), and the intratumour doxorubicin and gadolinium concentration was found, implying that the in vivo release of doxorubicin from the thermosensitive liposomes in the tumours can be probed (imaged) in situ with the longitudinal relaxation time of the co-released MRI contrast agent (dose painting).⁶⁵ 

Temperature sensitive liposomes release their encapsulated drugs at the melting phase transition temperature (T_m) of the lipid bilayer. At this T_m the lipid membrane changes its structure as it transfers from a gel to the liquid crystalline phase.⁶⁶  When the liposomal membranes are in the gel phase they show less permeability to molecules and water compared to the liquid crystalline phase.

The liposomes transition to the liquid crystalline phase can be achieved with the incorporation of a lyso-phospholipid such as MSPC (R = −C₁₇H₃₅). This lipid is also used in the Thermodox® formulation.⁴⁷  A potential disadvantage of MSPC containing liposomal formulations is the their rapid doxorubicin leakage at 37 °C.⁶⁴  Tagami et al. prepared temperature sensitive liposomes using nonionic surfactants Brij which are PEG-ylated lysolipids resembling the chemical structures of MSPC and DSPE-PEG²⁰⁰⁰. Results indicated that the optimal acyl chain length of the surfactant was between C(16) and C(18) with a saturated carbon chain and a PEG repeating unit ranging between 10 and 100 with a molecule weight above 600 Da. In the panel of surfactants tested, Brij78 was optimal and could be incorporated into the liposomes by the thin film hydration or the post-insertion method with an optimal range of 1 to 8 mol%.⁶⁷  The authors continued with in vivo experiments in mice bearing mammary carcinoma cells EMT-6, investigating Gd³⁺DTPA (diethylene triamine pentaacetic acid) release with NMR relaxometry. The authors observed a good correlation between relaxation enhancement in the heated tumour and the inhibition of tumour growth at day 21 post treatment.⁶⁸ 

Kono et al. investigated the effect of poly[2-ethoxy(ethoxyethyl)vinyl ether] chains (having a lower critical solution temperature) and polyamidoamine G3 dendron-based lipids with Gd³⁺ chelate residues into PEGylated liposomes. These theranostic liposomes exhibited excellent ability to shorten the longitudinal proton relaxation time (MRI). When administered intravenously into tumour-bearing mice, accumulated liposomes in tumours increased with time, reaching a constant level 8 hours after administration by following T₁-weighted MRI signal intensity. Liposome size affected the liposome accumulation efficiency in tumours; liposomes of about 100 nm diameter were accumulated more efficiently than those with about 50 nm diameter. Tumour growth was strongly suppressed when liposomes loaded with doxorubicin were administered intravenously into tumour-bearing mice and the tumour heated mildly at 44 °C for 10 min, 8 hours after administration.⁶⁹ 

Ultrasound sonication using microbubbles has been suggested to enhance the delivery of nanoparticles into the treated tumours.⁷⁰  Tumour-bearing mice were injected with microbubbles intravenously and sonicated before being injected with the PEGylated liposomal doxorubicin. It was demonstrated that FUS sonication with microbubbles can significantly enhance drug accumulation in the sonicated tumour at 24 hours after treatment. A significant hindrance to tumour growth was additionally achieved indicating that cavitation can enhance extravasation and drug tumour distribution.⁷⁰ 

HIFU has also been tested in combination with ultrasound imaging and microbubbles (ultrasound contrast agent/cavitation agent) and chemotherapy drugs. The effect of HIFU (1.0 MHz; 12 999 W cm⁻² in continuous waves), in the presence of hematoporphyrin and/or microbubbles, on the anticancer potency (in cells) of 5-fluorouracil, cisplatin, paclitaxel, mitomycin C or adriamycin, was investigated. Insonated adriamycin resulted in a lower death rate of human cancer cells HO-8910 (45.85 ± 2.65% vs. 34.84 ± 1.21%, p < 0.05), which was exacerbated when employing hematoporphyrin (34.84 ± 1.21% vs. 23.09 ± 7.82%, p < 0.05) or hematoporphyrin combined with microbubbles (34.84 ± 1.21% vs. 8.79 ± 3.69%, p < 0.05); the therapeutic activity was not affected when adding microbubbles alone. Overall the authors suggested to avoid the use of microbubbles as the enhanced ultrasonic cavitation produces a large amount of free radicals that affect the integrity of the anticancer agents.⁷¹  This suggests that a thorough study of the mechanisms involved in image guided FUS is required before selecting the imaging modality as well as the theranostic system. However, the selection of the chemotherapy appears to be most important.

Micelles have been also suggested as drug delivery carriers for FUS induced hyperthermia. Polymeric micelles (pluronics; P-105) were used to encapsulate doxorubicin.⁷²  These polymers were suggested as sensitisers of multidrug resistant (MDR) cells to chemotherapeutic drugs. The authors suggested that upon the accumulation of drug-loaded micelles at the tumour site, ultrasound treatment released the drug from micelles. This enhanced the intracellular uptake of both the released and encapsulated drug, possibly as a result of the induced extravasation and drug release from the micelles.⁷²  This work suggested that formulations other than thermosensitive liposomes could be used in combination with the ultrasound induced hyperthermia. Nevertheless such formulations have been less explored during recent years.

Targeted liposomes have also been investigated in combination with FUS. This was attempted to achieve therapeutically effective drug concentrations in tumours while avoiding healthy tissue damage (targeted drug delivery). In a recent work, a novel tumour-targeting peptide iRGD (CCRGDKGPDC) was used to modify the drug-loaded LTSL (iRGD-LTSL-DOX) and to investigate the anti-tumour effects when HIFU was applied to iRGD-LTSL-DOX specifically targeting α_νβ₃-positive cells and release the encapsulated doxorubicin (DOX) after triggering. In vivo results showed that DOX from iRGD-LTSL-DOX was intravascularly released and rapidly penetrated into the tumour interstitial space after FUS-triggered treatment, overcoming the limited tumour penetration of anticancer drugs. Strong anti-tumour efficacy further supported the effective combination of iRGD-LTSL-DOX and FUS-induced hyperthermia.⁷³ 

In our research project we aim to design theranostic liposomes. We investigated the potential of a labelled phospholipid/lysolipid containing liposomes to accumulate in tumours and release the drug under conditions of mild hyperthermia. We have prepared liposomal nanoparticles and we have investigated the potential of labelling for imaging. We developed optically labelled TSLs for image guidance drug delivery (iTSLs). These iTSLs were used to encapsulate the drug topotecan and iTSL-topotecan liposomes were administered in tumour bearing mice. These iTSLs allow the simultaneous, real-time diagnostic imaging of nanoparticle biodistribution using a near-infrared (NIR; 750–950 nm) fluorophore coupled to the lipidic component of the lipid bilayer. When combined with multispectral fluorescence analysis, this allowed for specific and high sensitivity tracking of the nanoparticles in vivo.

Application of hyperthermia indicated higher accumulation in tumours and concomitant drug release in the heated tumour.⁷⁴ 

Near Infrared Fluorescence (NIRF) labelling of the liposomes proved to be an excellent tool of monitoring the liposome accumulation in tumours and indicated that the liposomes signal increases with time (Figure 1.3).

When these liposomes were administered intravenously in mice with tumours and mild hyperthermia was applied for a short period of time, drug release was observed as assessed by fluorescence microscopy (Figure 1.4). Tumour sections that were isolated from tumours treated with hyperthermia clearly showed improved drug distribution within the tumour, compared to non-heated tumours (Figure 1.4). This study indicated that when formulations can be imaged (theranostics) modulation by means of triggered release is possible.

MRI labelling of the liposome nanoparticles can be done with a lipid that consists of a DOTA [1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid] headgroup (Figure 1.2).^75,76  Introducing the imaging lipid in the lipid bilayer provides better and clearer monitoring of liposomal particle kinetics and a better knowledge of the time required for maximum nanoparticles accumulation in tumours (monitored by MRI in the MRgFUS).

Combination of macromolecules, such as polymer drug conjugates with hyperthermia could also be a method of increased drug accumulation in the tumours that are treated with FUS. In a recent study researchers have introduced the use of heat shock (HS)-targeting towards tumour tissues as a method to enhance the drug accumulation and the retention, and improve therapeutic outcomes. HIFU was applied to generate hyperthermia in prostate tumour tissue in mice. When hyperthermia is applied an upregulation of the cell surface HS receptor glucose regulated protein 78 kDa (GRP78) was observed. This receptor was the further targeted by specific HS-targeting peptides attached to polymer drug conjugates. It was shown that HIFU-mediated HS-targeted copolymer–docetaxel conjugates improved treatment efficacy in a murine prostate tumour xenograft model. The study showed that the use of HIFU hyperthermia in combination with polymer therapeutics has potential to improve therapeutic outcomes in prostate cancer treatment.⁷⁷ 

In a similar study MRgFUS was found to enhance the delivery of both Evans blue dye (EBD) and gadolinium-chelated N-(2-hydroxypropyl)methacrylamide (HPMA) copolymers (polymer therapeutic). The EBD accumulation in the heated tumours increased by nearly two-fold compared to unheated tumours, an observation consistent with similar FUS treatments. The gadolinium-chelated HPMA copolymers also showed significant enhancement in accumulation over control as evaluated through MRI T1-mapping measurements.⁷⁸  These labelled drug conjugates indicate that macromolecules can also function as theranostics, of which delivery can be modulated.

Our recent results using NIRF labelling of an antibody indicates that the concept of theranostics can find application on established anticancer therapies. We have recently labelled trastuzumab with a near infrared label and monitored its distribution within tumours with or without repeated applications of FUS hyperthermia.⁷⁹  We have used a custom made FUS (Philips Research) that mimics the transducers used in the clinic (Figure 1.5).

Using this application of FUS as a non-invasive method of hyperthermia we increased the tumour temperature to 42 °C for a short period of time (3–5 min) and we observed an increased accumulation of labelled antibody. Repetition of focused ultrasound induced hyperthermic treatment further increased the accumulation of the antibodies in the tumour (Figure 1.6). This treatment also augmented the accumulation of other macromolecules non-specific to the tumour, such as IgG and albumin.⁷⁹  Image guided FUS could be used to enhance the therapeutic efficiency of antibodies and/or targeted nanoparticles.

Assessment of the current research in this field shows that the number of applications of image guided drug delivery that employ the focused ultrasound as a method of drug modulation is expanding. New methods and formulations are suggested that can maximise the effect of drug delivery. Combination of nanobubbles with TSLs proved to be an efficient method of delivering doxorubicin in mouse tumour model avoiding drug delivery in other tissues.⁸⁰ 

Triggered drug delivery using an external physical force provides the required control of drug deposition in certain tissues avoiding exposure of healthy tissues to toxic concentrations. The triggered induced delivery should be immediate and the effect induced on non-targeted tissues non-damaging. Hyperthermia induced by a means such as ultrasound can be exploited as an external trigger in drug delivery.^4,81 

Mild hyperthermia can be induced by pulsed FUS that can reduce extreme tissue heating by allowing the tissue to cool down between US exposures.⁸²  The increase in temperature can be 3–5 °C (hyperthermia) despite the high energy deposited in the tissue. Hyperthermia applied in tumours can increase blood flow and enhances vascular permeability. Studies with canine soft tissue sarcoma and human tumour clinical studies have also demonstrated that hyperthermia improves tumour oxygenation, and enhances response of such tumours to radiotherapy or chemoradiotherapy. The increased blood flow and vascular permeability caused by temperatures such as 42 °C may also improve the delivery of chemotherapy drugs, immunotherapeutic agents and genes to tumour cells.⁸³  FUS exposures in pulsed mode lowers the rates of energy deposition and generate primarily mechanical effects for enhancing tissue permeability to improve local drug delivery. These pulsed exposures can be modified for low-level hyperthermia as an enhancement of drug delivery that would lead to better drug deposition and better therapeutic effect.⁸⁴  Mild hyperthermia of 42 °C can improve the degree of nanocarrier extravasation as shown by Kong et al.⁸⁵  The reason that this leads to increased extravasation maybe due to down regulation of VE-cadherin that contributes to vascular integrity as it was shown in HUVEC endothelial cells.⁸⁶  It is clear that hyperthermia can provide a boost to extravasation and drug deposition in tumours. This should provide an adjuvant effect when nanocarriers are used and accumulate in tumours due to enhanced permeation and retention effect. It would be interesting to investigate the effect of hyperthermia on tumour/tissue drug clearance.

FUS can also induce non-thermal effects on tissues. Acoustic cavitation can be induced using microbubbles exposed to US.⁸⁷  Acoustic cavitation can be defined as the growth, oscillation and collapse of gas containing bubbles under the influence of the varying pressure field of sound waves in a medium and can have an effect on the permeability of a biological tissue.^87–89  There are two types of acoustic cavitation; non-inertial and inertial cavitation.⁹⁰  The non-inertial (stable) cavitation occurs when bubbles persist for a number of acoustic cycles. In this case the bubble’s size increases and decreases (the bubble expands and contracts) according to the applied ultrasound frequency. Inertial (transient cavitation) occurs when bubbles grow very fast and expand two or three-fold their resonant size, oscillate in an unstable manner and collapse in one compression half cycle.⁸⁸  It has been considered that the primary mechanism to affect the structure of intact cells is inertial cavitation that can induce irreversible damage as well as increase cell membrane permeability.^91,92  This has been tested as a hypothesis of drug permeation improvement.^93–96 

An important application of FUS and bubbles lies in the area of altering the permeability of the blood brain barrier.⁹⁷ 

In 2002, Mesiwala et al. observed that HIFU could alter BBB permeability. HIFU induced reversible, non-destructive, BBB disruption in a targeted area and this opening reversed after 72 hours. The authors showed with microscopy that HIFU either preserved brain architecture while affecting (opening) the BBB, or generated tissue damage in a small volume within the affected region of BBB. Further electron microscopy suggested that HIFU disrupted the BBB by widening capillary endothelial cell tight junctions, a mechanism used to open BBB.⁹⁸ 

The effect of FUS on the integrity of tight junctions was later confirmed in a study investigating the rat brain microvessels after this BBB disruption. The authors used immune-electron microscopy, to identify tight junctional proteins such as occludin, claudin-1, claudin-5, and sub-membranous ZO-1 junctional protein after sonication. They found substantial redistribution and loss of occludin, claudin-5 and ZO-1 indicating that the tight junction protein complex is affected. However, claudin-1 seemed less involved. Monitoring the leakage of horse radish peroxidase (large protein, permeability marker; MW 40 kDa) the authors observed that the BBB disruption appears to last up to 4 hours after sonication.⁹⁹  In a later study the role of caveolin in the mechanism of FUS BBB enhanced permeation was suggested. In a study investigating caveolae density it was found that caveolae and caveolin-1 were primarily localised in the brain microvascular endothelial cells of all the animals tested (rats) regardless of treatment, and that caveolin-1 expression was highest in the rats treated with both FUS and microbubbles. The authors in that study concluded that caveolin-1-mediated transcellular transport pathway may cooperate with other transport pathways (e.g. tight junctional disruption) to induce the opening of the BBB and the increased permeation of drug molecules.¹⁰⁰ 

Hynynen and colleagues investigated the BBB FUS enhanced permeability in rabbits. Rabbit brains were exposed to pulsed focused ultrasound while micro-bubbles were intravenously administered. The BBB opening was measured by an MRI contrast agent evaluating the local enhancement of permeation in the brain using MRI. The authors found that low ultrasound powers and pressure amplitudes were found to cause focal enhancement of BBB permeability to this contrast enhancing agent. Trypan blue injected before animals were sacrificed indicated blue spots in the areas of the sonicated locations.¹⁰¹  The authors concluded that HIFU disruption of BBB could be used to enhance drug delivery to the brain.¹⁰² 

McDannold et al. tested the safety of this method by searching for ischemia and apoptosis in areas with BBB disruption induced by pulsed ultrasound in the presence of gas bubbles and by looking for post treatment effects up to one month after sonication. Pulsed ultrasound exposures (sonications) were performed in the brains of rabbits under monitoring by MRI. BBB disruption was confirmed with contrast-enhanced MR images. Whole brain histologic examination was performed using staining for ischemic neurons and TUNEL staining for apoptosis. Tiny regions of extravasated red blood cells scattered around the sonicated locations, indicated capillaries. Despite these vasculature effects, only a few cells in some of the sonicated areas showed evidence of apoptosis or ischemia. The authors found that ultrasound-induced BBB disruption is possible without inducing substantial vascular damage that would result in ischemic or apoptotic death to neurons.¹⁰³ 

The method could find application in the delivery of large therapeutic molecules that do not normally permeate the BBB. Herceptin (trastuzumab), a humanised anti-human epidermal growth factor receptor 2 (HER2/c-erbB2) monoclonal antibody, was delivered locally and noninvasively into the mouse central nervous system through the BBB under image guidance by using an MRI-guided FUS. The amount of Herceptin delivered to the target tissue was correlated with the extent of the MRI-monitored barrier opening, making it possible to estimate indirectly the amount of Herceptin delivered. The method could be used to treat breast cancer metastases to the brain.¹⁰⁴  It was further shown that dopamine D(4) receptor-targeting antibody could also be delivered using the same technique in the brain.^105,106 

Delivery of small molecules can also be enhanced with the use of FUS and disruption of the BBB through cavitation. Treat et al. demonstrated relatively high concentrations of doxorubicin in the brain with minimal healthy tissue damage effects and effect that would benefit the delivery of small drugs that do not cross the BBB. The authors observed that doxorubicin accumulation in the non-targeted contralateral brain tissue remained significantly lower showing the efficiency of the method. MRI signal enhancement in the sonicated region correlated strongly with tissue doxorubicin concentration, suggesting that contrast-enhanced MRI could perhaps indicate drug penetration during image-guided interventions.¹⁰⁷ 

Konofagou and co-workers assessed the spatial permeability of the BBB-opened region using dynamic contrast-enhanced MRI (DCE-MRI) in mice. The authors generated permeability maps and K_trans (the transfer rate constant from the intravascular system to the extracellular extravascular space) values were estimated for a predefined volume of interest in the sonicated and the control area for each mouse. The results demonstrated that K_trans in the BBB-opened region was at least two orders of magnitude higher when compared to the contralateral (control) side confirming the hypothesis that FUS can transiently open the BBB.¹⁰⁸ 

There are several parameters that affect the level of BBB enhanced permeability and the endothelial tight junctions disruption; the pulse sequence comprising short bursts, the spacing between bursts or the rate of infusion of the microbubbles, and the size of microbubbles were found to affect the effect on BBB disruption.^109,110 

The method could be applied for a number of therapeutic applications. The brain-derived neurotrophic factor (BDNF) was delivered to the left hippocampus in mice through the noninvasively disrupted BBB using FUS. The BDNF bioactivity was found to be preserved following delivery as assessed quantitatively by immunohistochemical detection of the pTrkB receptor and activated pAkt, pMAPK, and pCREB in the hippocampal neurons. It was shown that BDNF delivered this way induced signalling effects in a highly localised region in the brain.¹¹¹ 

FUS with the presence of microbubbles has been shown to induce transient and local opening of the BBB for the delivery of therapeutic molecules which normally cannot penetrate into the brain. The success of FUS brain-drug delivery relies on its integration with in vivo imaging to monitor the kinetic change of therapeutic molecules (image guided drug delivery).

However, it is the area of targeting brain tumours that has attracted most interest in the FUS disrupted BBB.¹¹²  Mei and colleagues investigated the effects of targeted and reversible disruption of the BBB by MRI-guided FUS and delivery of methotrexate to the rabbit brain. The authors recorded that the methotrexate concentration in the sonicated group was notably higher than that in both the control group (intravenous administration) and the internal carotid artery administered group. They observed a greater than 10-fold increase in the drug level compared to internal carotid administration without FUS.¹¹³ 

Liu et al. investigated the delivery of 1,3-bis(2-chloroethyl)-1-nitrosourea (BCNU) to glioblastomas in rats with induced tumours with the help of FUS. The authors found that FUS significantly enhanced the penetration of BCNU through the BBB in normal and tumour-implanted brains without causing bleeding. Surprisingly, treatment of tumour-implanted rats with FUS alone had no beneficial effect on tumour progression. However, treatment with FUS before BCNU administration controlled tumour progression and improved animal survival relative to untreated controls.¹¹⁴ 

The blood–brain/tumor barrier inhibits the uptake and accumulation of chemotherapeutic drugs, Liu and colleagues recently assessed FUS-mediated delivery of an iron oxide magnetic nanoparticles (MNPs) conjugated to an antineoplastic agent, epirubicin. They used MNPs because of the favourable MRI characteristics, which could facilitate imaging. They demonstrated a substantial accumulation of MNPs, as well as epirubicin, up to 15 times the therapeutic range in the brain when delivered with FUS. They further showed decreased tumour progression in animals with brain tumours that received MNP with epirubicin via FUS.¹¹⁵ 

Receptor targeting liposomal nanocarriers have been combined with image guided FUS to treat brain tumours. In a recently presented study it was shown that pulsed HIFU and human atherosclerotic plaque-specific peptide-1 (AP-1)-conjugated liposomes containing doxorubicin (AP-1 Lipo-Dox) acted synergistically in an experimental brain tumour model. Prior to each sonication, AP-1 Lipo-Dox or unconjugated Lipo-Dox were administered intravenously, and the concentration in the brain was quantified. Drug injection with sonication increased the tumour : normal brain doxorubicin ratio of the target tumours by about two fold compared with the control tumours. Moreover, the tumour : normal brain ratio was highest after the injection of AP-1 Lipo-Dox with sonication. The results of this study indicate that combining targeting strategies can substantially enhance delivery of chemotherapy in the brain.¹¹⁶  In a separate study the authors investigated the pharmacokinetics of ¹¹¹I-labeled AP1-Lipo-dox using microSPECT. The authors confirmed that sonication increased liposomal doxorubicin concentrations in tumour areas (murine glioblastoma) and that molecular targeting acts synergistically with FUS.¹¹⁷ 

Targeted gene transfer into central nervous system was investigated using MRI-guided FUS-induced BBB disruption. The results of this study showed that MRI-guided FUS achieved plasmid DNA transfer across the opened BBB furthermore plasmid were endocytosed by neuronal cells presenting heterogeneous distribution. BDNF (and BDNF-EGFP) expressions were markedly enhanced by the combination of ultrasound and plasmid BDNF-EGFP-loaded microbubbles about 20-fold compared to that of the control group. This indicates that this strategy can be used to deliver functional biologicals and gene therapy. The method of using MRI-guided FUS to induce the local BBB disruption could accomplish effective targeted gene therapy in the CNS. In this type of experiment the microbubbles are used as the plasmid carrier. The investigators conjugated plasmids onto the surface of microbubbles and they coated these carriers using polymers in a layer by layer technique.¹¹⁸ 

An exciting application is the delivery of therapeutic stem cells to the brain using FUS to potentially treat neurodegenerative diseases, traumatic brain injury, and stroke. MRI guidance was used to target the ultrasound beam thereby delivering iron-labelled, green fluorescent protein (GFP)-expressing neural stem cells specifically to the striatum and the hippocampus of the rat brain. Immunohistochemical analysis confirmed the presence of GFP-positive cells in the targeted brain regions suggesting that MRIgFUS may be an effective alternative to invasive intracranial surgery for stem cell transplantation.¹¹⁹ 

Contrast-enhanced magnetic resonance imaging (CE-MRI) is used to monitor contrast agent leakage to verify BBB-opening and infer drug deposition. However, despite being administered concurrently, microbubbles, therapeutic agent, and contrast agent have distinct pharmacokinetics and pharmacodynamics, thus complicating the quantification and optimisation of BBB-opening and drug delivery. Multifunctional microbubbles (MB) were loaded with therapeutic agent (doxorubicin; DOX) and conjugated with superparamagnetic iron oxide (SPIO) nanoparticles providing a theranostic tool to treat brain tumours. These DOX-SPIO-MBs were designed to concurrently open the BBB and perform drug delivery upon FUS exposure, act as bimodal MRI and ultrasound contrast agent, and allow magnetic targeting (MT) to achieve enhanced drug delivery (theranostic). Burst-tone FUS was applied after injection of DOX-SPIO-MBs, followed by magnetic targeting with an external magnet attached to the scalp in a rat glioma model. The authors found that BBB-opening and drug delivery were achieved concurrently during the FUS treatment. In addition, MT increased local SPIO deposition in tumour regions by 22.4%. Our findings suggest that DOX-SPIO-MBs with FUS could be an excellent theranostic tool for future image-guided drug delivery to brain tumours.¹²⁰ 

Magnetic nanoparticles have been suggested in combination with FUS and microbubbles for magnetic hyperthermia. The superparamagnetic properties of the MNPs provide the opportunity to guide them by an externally positioned magnet and also provide contrast for MRI (theranostics). However, their therapeutic effect in treating gliomas, is limited by insufficient local accumulation and retention due to their inability to traverse biological barriers. To overcome this the use of FUS in combination with magnetic targeting was suggested as a mechanism to deliver therapeutic MNPs across the BBB and to enter the brain both passively and actively. MRI was used to monitor and quantify their distribution in vivo. Synergistic targeting and image monitoring are powerful techniques for the delivery of macromolecular chemotherapeutic agents into the CNS under the guidance of MRI.¹¹⁵  This study shows that theranostics and their guided delivery can overcome difficult biological barriers and deliver the therapy at the right site.

During recent years there has been an expansion in research in the design of MRgFUS drug delivery. The main dosage forms tested in MRgFUS drug delivery strategy are the thermosensitive liposomes and the lipid based microbubbles that can incorporate drugs or other liposomes on their surface.^118,121 

There is limited research in the area of using other responsive materials or nanocarriers. Rapoport recently discussed the potential of using micelles and FUS¹²²  for enhanced tissue permeation. Micelles are nanosized carriers able to carry hydrophobic drugs; their combination with FUS could substantially enhance their delivery in tissues. Kostarelos and colleagues suggested the incorporation of thermosensitive peptides onto liposome bilayers to enhance thermo-responsiveness,¹²³  and the group of Lammers designed polymer based micro-bubbles for US drug release.¹²⁴ 

A novel type of HIFU triggered active tumour-targeting polymeric micelle was recently prepared and investigated for controlled drug release and enhanced cellular uptake. Amphiphilic hyaluronic acid (HA) conjugates were synthesised to form docetaxel loaded micelles in aqueous conditions with high encapsulation efficiencies of over 80%. It was shown that HIFU enhanced the cellular uptake behaviour by altering the permeability of the cell membrane. It was also able to aid with the extravasation of micelles into the interior of tumours. These micelles can emerge as promising nanocarriers of chemotherapeutic agents for image guided FUS controlled drug release and tumour targeting.¹²⁵ 

Another recent study suggested the use of a multifunctional hollow mesoporous Prussian blue (HMPBs) theranostic nanoplatform, the hollow structure of which is capable of encapsulating doxorubicin (DOX) and perfluorohexane (HMPBs-DOX/PFH) to be used in combination with FUS. In vitro and in vivo studies validated that HMPBs-DOX/PFH can be used as an amplifiable dual-mode imaging contrast agent, which can simultaneously enhance US and photoacoustic imaging for guiding and monitoring tumour therapy. When exposed to HIFU, this versatile HMPBs-DOX/PFH agent could increase the cavitation effect and use lower HIFU intensity to achieve coagulative necrosis.¹²⁶ 

From the above it is evident that already established delivery systems such as different structurally nanocarriers have not been investigated in combination with image guided FUS. It would be interesting to see the effect of FUS on the enhanced permeability of different micelles, polymers (dendrimers) or metal nanoparticles (gold-iron) to tissues. Thermosensitive materials have been hardly explored in this field. Polymers or proteins that respond to small change of temperature could form suitable image guided FUS triggered platforms.

Overall MRgFUS drug delivery is a novel and valuable tool to increase drug targeting and tissue specific drug delivery. It is expected that future studies will prove the clinical efficacy of MRgFUS drug delivery applications.