**5. NP-based thermal therapy using radiofrequency**

#### **5.1. Standard RF-mediated nanoablation**

An important progress has been made in improving the quality of the MNPs; therefore, for construction, high temperature crystallization or different coatings were used, such as dextran,

Several authors introduced MNPs either in the core or in between the lipid bilayer of thermosensitive liposomes and, on alternating magnetic field AMF heating, the encapsulat‐ ed drugs were released [43]. Shinkai utilized liposomes where he introduced magnetite nanoparticles (with a diameter of 10 nm). After administration, these nanoparticles in‐ creased the temperature of the tissue [57]. In another study, Ito injected magnetite cationic liposomes (MCLs) into the tumor tissue. They heated the tissue above 43°C and obtained a complete regression of mammary carcinomas in all mice [58]. Also, Jimbow [52] developed a particle with N-propionylcysteaminylphenol (NPrCAP) conjugated onto the surface of magnetite nanoparticles (NPrCAP/M). The result was the inhibition of melanoma cells growth as a result of the production of cytotoxic free radicals. In another study, a thermosensitive polymer was layered onto MNPs covalently coupled to doxorubicin with an acid-labile hydrazine bond that showed release on heating with AMF and a pH of 5.3 (the pH of endosomes) [59]. The authors combined via emulsification MNPs with a polyvinyl alcohol polymer and encapsulated hydrophobic/ hydrophilic drugs. The drugs were released after

Direct intratumoral injection was used in the first MNP HT clinical trial treating a patient with a recurrent prostatic tumor [61]. Through the use of transrectal ultrasound and fluoroscopy guidance, the authors performed a transperineal injection of the MNPs into the prostate. After the administration of MNPs, the particles were selectively heated in an externally applied alternative magnetic field. The conclusions of these trials were encouraging. Due to the low clearance of MNPs from tumors, serial heat treatments were possible after a single magnetic fluid injection. Another positive aspect was the fact that a low magnetic field was used to produce the necessary temperatures. Furthermore, this treatment does not cause discomfort or serious side effects. In these studies, the CT exam had an accuracy rate of 85% in evaluating the treatment-related parameters. The same good results were obtained later in human glioma

In 2008, Takamatsu et al. combined the intra-arterial selective HT with the transcatheter arterial embolization technique in a rabbit model for renal carcinoma [64]. For injection, they utilized a mixture of commercially available nano-sized magnetic particles (Ferucarbotran) and lipiodol as embolic material. The mixture was injected into the renal artery under fluoroscopic guidance. The intratumoral temperatures of 45ºC were obtained after the area was exposed to an external alternating-current magnetic field. Even the result was not spectacular (the treated tumor was hypovascular) the authors speculated that this method can be used only in hypervascular tumors. In another study, Huang HS injected IV MNPs (1.9 mg Fe/g tumor) in a subcutaneous squamous cell carcinoma mouse model. After the injection, they applied a field of 38 kA/m at 980 kHz; therefore, the tumors could be heated to 60°C in 2 min. The results were encouraging, showing an ablating with millimeter (mm) precision and a surrounding tissue

polyethylene glycol (PEG), dopamine, silanes and gold [43].

232 Recent Advances in Liver Diseases and Surgery

the heating with an alternative magnetic field [60].

trials [62, 63].

intact [43].

StandardRFAisaninvasiveprocedurethatrequires theinsertionofelectrodeswithinthetumor. Tumor destruction occurs as a result of vibrations of ions within tumor tissue induced by radio waves, which give rise to friction and lethal heat. Although it is possible to achieve local control in liver tumors < 2.5 cm, in larger lesions local tumor recurrence is common [67, 68].

Initially, in order to increase the efficacy of RFA, the ablation guidance methods were improved (contrast-enhanced ultrasound, fusion imaging, etc.), but this led only to a slight efficacy improvement. Because of the changes that occur after RFA (increased vascular and cellular membrane permeability), the periphery of the tumor becomes more susceptible to chemo‐ therapy. Thus, the combination of thermal ablation and chemotherapy seemed to lead to promising results. The results of these methods did increase the efficacy of RFA, but it was not enough. Therefore, new treatments that will augment cytotoxicity at the margin of the ablation zone have been developed.

The efficiency of RFA can be significantly enhanced by administration of special thermal absorbing agents such as NPs, which are targeted into a tumor area (actively or passively) with the purpose to release locally the retained heat and thus enhance tumoral destruction.

The NPs in free form or those containing various anti-cancer agents may be administrated before, at the time, or after RFA [68, 69]. Administering CYT-6091, a TNF-labeled NP, 4 h prior to RFA yielded a significantly larger zone of central necrosis and a 23% increase in ablation volume in comparison to RFA alone [69]. Using this NP enhanced ablation, the partially ablated tissue at the periphery was replaced by completely ablated tissue [69].

The administration of NPs containing free doxorubicine at the time of RFA or after leads to an increased diameter of coagulated tumor tissue (and increased concentration of doxorubicine in the ablated tumor) [68]. The NPs accumulate in the region of ablation both in the treated tumor (as result of an increased leakage) and in the peripheral region with thermal induced inflammation. This is known as the enhanced permeability and retention (EPR) effect [70].

The liposomes were the first NPs that have been utilized in combination with RFA. The studies of Ahmed and Goldberg demonstrated that the use of lipid NPs as carriers of a drug combined with ARF was associated with an increased accumulation of doxorubicin in the tumor, while non-encapsulated free doxorubicin did not have increased tumor uptake following RFA [71]. Since then, an important number of investigators improved the lipid layer of liposomes that has contributed to enhanced tumor damage secondary to formation of lipid hydro-peroxide leading to enhanced oxidative stress. Also, the investigators demonstrated that NPs size could influence the intratumoral drug accumulation and tissue coagulation [68].

#### **5.2. Non-invasive RF nanoablation**

As a negative relationship between the frequency of the waves and the depth of penetration exists, radio waves may be used as an alternative to heat tumors that are deeply located. The heating rate of a certain tissue is described by the formula HR = SAR/69.77 CH where SAR is the specific absorption rate and CH the specific heat capacity of the tissue (kcal/kg °C). As SAR (W/kg) depends on the dielectric conductivity of the tissue, an enhanced conductivity provided by AuNPs or carbon nanotubes may increase the heat delivered to the tissue [72].

These low-frequency electromagnetic waves have the advantage to penetrate human tissues and pass through the entire body with minimal perturbations until the RF fields interact with metal. The metal particles absorb RF energy and release heat to the adjacent region. Several reports suggested that tumoral hyperthermia may be improved through the use of targeted nanomaterials, which produce an intracellular hyperthermia and act as RF-thermal transduc‐ ers, leaving the surrounding healthy tissue intact [68].

The delivery of RF generated heat in deep structures may be achieved either by RF needle inserted into the tumor (standard RFA) or by an external device that generates an RF field [68, 72].

If standard RF ablation produces a hyperthermic region of 2–4 cm diameter around the probe's tip, the nanoparticle-mediated RF field induces a hyperthermic area of approximately 100 μm. The heating mechanism of NPs in an RF field is a complex phenomenon that is still under debate [73]. Most of the RF field devices produce shortwave RF fields (13.56 MHz), allowing them to be used in the medical field. Several reports have shown that Joule heating of the background ionic suspension where the NPs are suspended can be the main source of RF heat production [74]. A relative high variety of NPs as AuNPs, carbon nanotubes (SWNTs), quantum dots (cadmium-selenide and indium-gallium-phosphide), silicon nanoparticles (Si NPs), and La0.7Sr0.3MnO3 (Dex-LSMO) have been associated with RF field [74, 75]. The use of NPs seems to improve the standard RFA by increasing the specificity of tumor destructions and affording a relative target therapy. Between these NPs are several differences, such as the SWNTs are heated faster than AuNPs unlike quantum dots that are heated in a similar manner to AuNPs [73].

SWNTs showed that they can be activated from a distance by RF field to produce thermal cytotoxicity [75]. The SWNTs have been injected in Vx2 tumors and induced the necrosis of all tumors within 5 min of RF field exposure. Regions of necrosis were identified with 2–5 mm borders. It is important to highlight that SWNTs alone or RF field exposure alone did not induce any measurable tumor necrosis or liver injury. In another study, the authors demonstrated that SWNTs injected into malignant cells may allow noninvasive RF field treatments to produce lethal thermal injury to the malignant cells. In a similar study conducted by Raoof, Hep3B and HepG2 cells were injected to kentera modified SWNT and were exposed to an 800 W RF field. Significant thermal cytotoxicity was demonstrated with 2 min of RF exposure in a concentration-dependent manner [75]. Also the group conducted by Cardinal obtained similar results after they exposed a rat model (with HepG2 cells) into an RFA field following the administration of AuNPs [76]. In a study conducted by Glazer ES, AuNPs utilized cetuximabconjugated AuNPs in nonionizing RF radiation to investigate human pancreatic xenograft destruction in a murine model [73]. The result showed an increased apoptosis with decreased viability of tumoral cells after treatment with cetuximab-conjugated AuNPs and RF field exposure. Another important observation was the lack of injury to other organs.

It becomes a reality the fact that nanotechnologies will play a major role in new antitumoral therapies. In the last years, the thermal approach using nanoparticles, nanoemulsion, pH responsive nanoparticles, nanoparticles combined with radiation, and nanovectors for drug delivery have been the most evaluated nanoparticle-based cancer treatment methods. The ability of SWNTs to convert NIR laser radiation into heat, due to the photon–phonon and electron interactions, provides the opportunity to create a new generation of immunoconju‐ gates for cancer phototherapy. In 2011, Iancu et al. demonstrated that the HepG2 cells treated with multi-walled carbon nanotubes (HSA–MWCNTs) following laser irradiation had a higher necrotic rate compared with normal cells [77].

#### **5.3. Thermosensitive liposomes currently in advanced clinical trials**

The liposomes were the first NPs that have been utilized in combination with RFA. The studies of Ahmed and Goldberg demonstrated that the use of lipid NPs as carriers of a drug combined with ARF was associated with an increased accumulation of doxorubicin in the tumor, while non-encapsulated free doxorubicin did not have increased tumor uptake following RFA [71]. Since then, an important number of investigators improved the lipid layer of liposomes that has contributed to enhanced tumor damage secondary to formation of lipid hydro-peroxide leading to enhanced oxidative stress. Also, the investigators demonstrated that NPs size could

As a negative relationship between the frequency of the waves and the depth of penetration exists, radio waves may be used as an alternative to heat tumors that are deeply located. The heating rate of a certain tissue is described by the formula HR = SAR/69.77 CH where SAR is the specific absorption rate and CH the specific heat capacity of the tissue (kcal/kg °C). As SAR (W/kg) depends on the dielectric conductivity of the tissue, an enhanced conductivity provided

These low-frequency electromagnetic waves have the advantage to penetrate human tissues and pass through the entire body with minimal perturbations until the RF fields interact with metal. The metal particles absorb RF energy and release heat to the adjacent region. Several reports suggested that tumoral hyperthermia may be improved through the use of targeted nanomaterials, which produce an intracellular hyperthermia and act as RF-thermal transduc‐

The delivery of RF generated heat in deep structures may be achieved either by RF needle inserted into the tumor (standard RFA) or by an external device that generates an RF field

If standard RF ablation produces a hyperthermic region of 2–4 cm diameter around the probe's tip, the nanoparticle-mediated RF field induces a hyperthermic area of approximately 100 μm. The heating mechanism of NPs in an RF field is a complex phenomenon that is still under debate [73]. Most of the RF field devices produce shortwave RF fields (13.56 MHz), allowing them to be used in the medical field. Several reports have shown that Joule heating of the background ionic suspension where the NPs are suspended can be the main source of RF heat production [74]. A relative high variety of NPs as AuNPs, carbon nanotubes (SWNTs), quantum dots (cadmium-selenide and indium-gallium-phosphide), silicon nanoparticles (Si NPs), and La0.7Sr0.3MnO3 (Dex-LSMO) have been associated with RF field [74, 75]. The use of NPs seems to improve the standard RFA by increasing the specificity of tumor destructions and affording a relative target therapy. Between these NPs are several differences, such as the SWNTs are heated faster than AuNPs unlike quantum dots that are heated in a similar manner

SWNTs showed that they can be activated from a distance by RF field to produce thermal cytotoxicity [75]. The SWNTs have been injected in Vx2 tumors and induced the necrosis of all tumors within 5 min of RF field exposure. Regions of necrosis were identified with 2–5 mm

by AuNPs or carbon nanotubes may increase the heat delivered to the tissue [72].

influence the intratumoral drug accumulation and tissue coagulation [68].

**5.2. Non-invasive RF nanoablation**

234 Recent Advances in Liver Diseases and Surgery

[68, 72].

to AuNPs [73].

ers, leaving the surrounding healthy tissue intact [68].

Discovered in 1964 by Alec Bangham, liposomes are self-assembling, biocompatible, biode‐ gradable, and nonimmunogenic nanovesicles consisting of a lipid bilayer enclosing an aqueous phase [78]. The features of liposomes allow for a wide range of drug delivery; consequently, hydrophilic drugs can be trapped in the liposome's aqueous compartments while the lipid bilayer can be utilized to incorporate hydrophobic drugs. Due to the discontinuous endothelial lining and the lack of efficient lymphatic drainage of the tumor, the extravasations of liposomes into the interstitial space is increased and the liposomes can accumulate in the tumoral tissue; therefore, they will function as a sustained drug-release formula [79]. Immordino mentioned for the first time this process and named it as EPR effect [80]. Moreover, the combination (liposome–chemotherapy) changes drug pharmacokinetic properties and minimizes its systemic toxicity. Furthermore, the drug prevents the entrapped drug from premature inactivation in the circulation. The main issue of liposomes is that they are rapidly phagocy‐ tized by the mononuclear phagocyte (MP) and removed from the blood circulation after intravenous injection. To avoid this inconvenience, the authors developed a grafting poly- (ethylene glycol) (PEG) or oligoglycerol-moieties on the surface of the liposomal carrier. By reducing MP system uptake [80], long-circulating PEGylated liposomes can passively accu‐ mulate into solid tumors undergoing angiogenesis. Another improvement was the incorpo‐ ration of additional lipid compounds that further enhance membrane permeability at the phase transition temperature of the lipid membrane (lysolipid or oligoglycerol-polyglycol) [79, 81– 84]. The result was a long blood circulation time *in vivo*. These types of low temperature thermosensitive liposomes (LTSLs)[79] are injected just prior to or during the HT treatment, with immediate release of their contents upon arrival in the heated tumor area.

The main limit of this type of therapy remains the intimate relation between the biodisponi‐ bility of liposomes and the vascular permeability. It is important to underline that vascular permeability between different tumor types and even within tumors can be highly variable, resulting in unpredictable liposome extravasation into the tumor tissue [85, 86]. Due to the combination of sub-optimal drug release kinetics and unpredictable vascular permeability, only modest results in the therapeutic index of chemotherapy have been obtained using liposomes for target drug delivery [87].

An important progress in the use of liposomes was the invention of small, 100 nm-long circulating liposomes that have a long blood-residence time as their main characteristic. These favorable circulation properties resulted in an enhanced accumulation of liposomal drugs in the tumor area.

To date, several liposomal products have been approved for clinical use: liposomes with doxorubicin (Doxil/Caelyx, Myocet, and Lipo-Dox) for treatment of Kaposi's sarcoma, ovarian cancer, breast cancer, and multiple myeloma; liposomes with daunorubicin (DaunoXome) for treatment of Kaposi's sarcoma; and liposomes encapsulating vincristine (Marqibo) for acute lymphoblastic leukemia [88].

Hyperthermia represents the heating of tumors to temperatures of up to 43°C. The main effect consists of an increased tissue perfusion, oxygenation and blood flow velocity, and microvessel permeability contributing to increased antibodies, drug, or nanoparticles levels in tumors at clinically tolerated temperatures [89–92]. Nowadays, hyperthermia for triggering TSLs is applied locally and in a noninvasive way from an external source to a targeted area using focused ultrasound technology (FUS) and high-intensity focused ultrasound (HIFU), or invasively using ARF or MWA [93, 94]. For superficial tumors, the authors used regional HT and external antennas or applicators that emit microwaves or radio waves. Localized HT is used to destroy deeply located tumors. The antennas (microwave antennas, radiofrequency electrodes) are inserted directly within the tumor. The major limit of this heating method is the tumor diameter (less than 5 cm). Focused ultrasound is used to heat small lesions (mm). In a recent study Dromi et al. combined LTSLs with hyperthermia from FUS [95]. They obtained an increased drug discharge at the tumoral area and the most important tumor had a delayed growth.

The newest heating method is magnetic resonance guided focused ultrasound technology (MRgFUS). These combinations allow simultaneous treatments, imaging to guide the treat‐ ment and MR thermometry to noninvasively monitor temperature changes and assure feedback in real-time [87]. In two recent studies, the authors used MRgFUS and drug-loaded liposome in rat [96] and rabbit [97] models. The results showed that the combination MRgFUS with drug loaded liposome assured the greatest uptake of the drug when compared to controls (liposome only and/or free drug). Several studies have analyzed the combination of RFA and the non-thermally sensitive liposomal doxorubicin, showing larger ablation zones compared with RFA alone, both at the preclinical and clinical levels. The suggested mechanisms for the synergistic effect of liposomal doxorubicine and RFA are as follows: increased markers of DNA breakage, oxidative stress and apoptosis, increased heat-shock protein 70 in the areas sur‐ rounding the ablation zone after combination treatment [98, 99]. In addition, Ahmed and colleagues observed that after combining RFA with Doxil, the intratumoral drug uptake increased, while the dose of doxorubicin necessary for tumor destruction decreased [100].

transition temperature of the lipid membrane (lysolipid or oligoglycerol-polyglycol) [79, 81– 84]. The result was a long blood circulation time *in vivo*. These types of low temperature thermosensitive liposomes (LTSLs)[79] are injected just prior to or during the HT treatment,

The main limit of this type of therapy remains the intimate relation between the biodisponi‐ bility of liposomes and the vascular permeability. It is important to underline that vascular permeability between different tumor types and even within tumors can be highly variable, resulting in unpredictable liposome extravasation into the tumor tissue [85, 86]. Due to the combination of sub-optimal drug release kinetics and unpredictable vascular permeability, only modest results in the therapeutic index of chemotherapy have been obtained using

An important progress in the use of liposomes was the invention of small, 100 nm-long circulating liposomes that have a long blood-residence time as their main characteristic. These favorable circulation properties resulted in an enhanced accumulation of liposomal drugs in

To date, several liposomal products have been approved for clinical use: liposomes with doxorubicin (Doxil/Caelyx, Myocet, and Lipo-Dox) for treatment of Kaposi's sarcoma, ovarian cancer, breast cancer, and multiple myeloma; liposomes with daunorubicin (DaunoXome) for treatment of Kaposi's sarcoma; and liposomes encapsulating vincristine (Marqibo) for acute

Hyperthermia represents the heating of tumors to temperatures of up to 43°C. The main effect consists of an increased tissue perfusion, oxygenation and blood flow velocity, and microvessel permeability contributing to increased antibodies, drug, or nanoparticles levels in tumors at clinically tolerated temperatures [89–92]. Nowadays, hyperthermia for triggering TSLs is applied locally and in a noninvasive way from an external source to a targeted area using focused ultrasound technology (FUS) and high-intensity focused ultrasound (HIFU), or invasively using ARF or MWA [93, 94]. For superficial tumors, the authors used regional HT and external antennas or applicators that emit microwaves or radio waves. Localized HT is used to destroy deeply located tumors. The antennas (microwave antennas, radiofrequency electrodes) are inserted directly within the tumor. The major limit of this heating method is the tumor diameter (less than 5 cm). Focused ultrasound is used to heat small lesions (mm). In a recent study Dromi et al. combined LTSLs with hyperthermia from FUS [95]. They obtained an increased drug discharge at the tumoral area and the most important tumor had

The newest heating method is magnetic resonance guided focused ultrasound technology (MRgFUS). These combinations allow simultaneous treatments, imaging to guide the treat‐ ment and MR thermometry to noninvasively monitor temperature changes and assure feedback in real-time [87]. In two recent studies, the authors used MRgFUS and drug-loaded liposome in rat [96] and rabbit [97] models. The results showed that the combination MRgFUS with drug loaded liposome assured the greatest uptake of the drug when compared to controls (liposome only and/or free drug). Several studies have analyzed the combination of RFA and

with immediate release of their contents upon arrival in the heated tumor area.

liposomes for target drug delivery [87].

236 Recent Advances in Liver Diseases and Surgery

the tumor area.

a delayed growth.

lymphoblastic leukemia [88].

In order to optimize the effects of liposomes, the use of TSLs that trigger the release of the drug at the edge of the heated zone was suggested [101–103]. These TSLs contain thermosensitive lipids in their bilayer, undergoing a gel-to-liquid phase transition at the desired temperature (usually between 41°C and 43°C), after which the drug enters tumor cells in free form. This conversion is the consequence of a conformational change in the alkyl chains of the lipids, which leads to an increase in the volume occupied by the hydrocarbon chains in the membrane and thus an increase in the permeability of the lipid bilayer [79]. Common TSLs have been composed from 1, 2-dipalimitoyl-sn -glycero-3-phosphocholine (DPPC) as the primary lipid, because its phase transition temperature (Tm) occurs at 41.5°C.

In 2009, TSLs containing Dox known as ThermoDox®, became the first heat-triggered release formula of the anthracycline doxorubicin that reached pharmaceutical development (Celsion Corporation, Columbia, Maryland, USA) and clinical application [104–105]. Thermodox® is composed of DPPC:MSPC:DSPE-PEG2000 (86:10:4 molar ratio) and in combination with mild was used in the Phase III clinical trial to treat hepatocellular carcinoma and the Phase II trial in combination with local mild for patients with recurrent breast cancer of the chest wall and colorectal liver. After intravenous administration, Thermodox® concentrates in the liver where it rapidly permeates HCC lesions and their vasculature. Regarding safety and tolerability, in Phase I ThermoDox® was associated with low side effects and the maximum tolerated dose was established at 50 mg/m2. According to the Phase I trial, RFA and ThermoDox® may be used as a front-line therapy for HCC > 3 cm [106]. Unfortunately, in 2013 Celsion Corp. was unable to demonstrate the effectiveness of ThermoDox® in the improvement of free survival [79]. It seemed that the temperature of drug release is different between *in vivo* and *in vitro*. In a study conducted by Hossann, about 90–100% Dox release from LTSLs in plasma or serum at 39–40°C resulted in 2°C below the theoretical temperature [107]. Therefore, it might be that all drug content is released from the LTSLs below 41–42°C, which means that the drug is discharged in blood circulation before the accumulation of LTLS in the target heated tumoral area [79]. In a recent study, after the incorporation of lysophosphatidylcholines (lyso-PC, e.g. 1-stearoyl-2-hydroxy-sn-glycero-3-phosphocholine, MSPC) into the liposomal membrane, it was possible to further accelerate the encapsulated drug at Tm [108].

Fine tunings in drug release kinetics of LTSLs was demanded to assure an improved dug release [109]. In 2014, Chen J evaluated [79] high temperature triggered TSLs (HTSLs) com‐ posed of DPPC and hydrogenated soy phosphatidylcholine (HSPC). For these types of liposomes, the theoretical temperature of discharge of HTSLs was set at 44°C; thus, the body temperature had less influence on the drug release from the vesicles. The result of this study was encouraging. Compared to conventional LTSLs, the new formula of HTSLs was associated with higher stability and less content discharge to the heated tumor area.

Several authors recommended the attaching of targeting ligands to the nanoparticles to assure a more specific localization and retention of the liposomal drug in tumors. Another reason to utilize these ligands is the capacity of promoting active cellular uptake of the drug-containing nanoparticles through binding to targeted internalizing receptors [110-112].

The cationic TSLs, called CTSLs (cationic thermosensitive liposome) is a new class of LTSL that contains a cationic lipid in its membrane. The CTSLs are absorbed by vascular endothelium and tumor cells; afterwards, they release their contents upon applying a temperature trigger [113]. It seems that, once accumulated, rapid drug release by intracellular cationic liposomes may achieve high intracellular concentrations of drug, thereby maximizing damage to both the endothelial cell and tumor cell compartments [113]. To evaluate tumoral accumulation of liposomes, radionuclides and nuclear imaging may be used. Even if the authors have obtained good results, in the future these types of treatment will have to demonstrate their therapeutic potential in clinical practice.
