Archive for the ‘Hyperthermia’ Category

Thermal Ablation Sutter Cancer Center Treatments & Services

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Posted 04 Sep 2011 — by James Street
Category Hyperthermia, Radio Frequency RF
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One of the most promising advances in many years, thermal ablation treats cancer tumors using heat-generating probes inserted directly into malignant tissue.

Gregory Graves, M.D. Gregory Graves, M.D. and Scott Foster, M.D.

Because cancer cells are more susceptible to destruction by heat than normal tissue cells, thermal ablation allows surgeons to treat tumors with minimal damage to surrounding tissue. Especially in cancers of the liver, lung, bone and kidney that are often difficult to treat surgically, thermal ablation offers a better alternative to eliminate or shrink tumors and reduce pain. Sutter Cancer Center, Sacramento oncologic surgeon Gregory Graves, M.D., John Lee, M.D. and interventional radiologist Scott Foster, M.D. introduced thermal ablation for lung cancer treatment to Sacramento and continue to research new techniques and applications.

What Is Thermal Ablation? | back to top
There are two types of thermal ablation: radiofrequency (RFA) and microwave. Both are minimally invasive techniques that treat cancer by applying intense heat through a small probe inserted directly into the tumor.
RFA and microwave ablation treatments aim to reduce patient symptoms, improve quality of life and increase survival rate.

How Does Thermal Ablation Work? | back to top
Surgeons insert a probe into the tumor. Alternating electrical currents pass through the tumor, heating the tissue. The heat destroys the cells and ablates (destroys) the tumor.

What are the Benefits of Thermal Ablation? | back to top
Thermal ablation provides an excellent alternative to major surgery, which can pose substantial risks even with the best and most experienced surgeons. Thermal ablation significantly reduces risks and speeds recovery. Additionally, follow-up imaging and treatment may be easier after a thermal ablation procedure.

What Are the Risks? | back to top
As with every procedure, there are risks. The risk of major complication due to thermal ablation is one to two percent. Bleeding is the most common complication. Depending on the size of the tumor, it is possible that thermal ablation will destroy only part of the tumor. If this is an issue in your case, your physician will discuss the risk during your consultation.

Who Does the Procedure? | back to top
The Sutter Cancer Center is staffed by leading specialists in thoracic and oncologic surgeries, interventional radiology, radiation therapy and pulmonary treatment. Thoracic and oncology surgeon Gregory Graves, M.D., oncology surgeon John Lee, and interventional radiologist and Scott Foster, M.D. perform the procedure. In addition, they are the specialists who introduced this new technology for lung cancer to Sacramento.

How Do I Learn More? | back to top
To find out more about radiofrequency and microwave ablation, call Jeannine Graves, R.N., CNOR, at (916) 454-6913 or email capsurg@sutterhealth.org.

Computationally Guided Photothermal Tumor Therapy Using Long-Circulating Gold Nanorod Antennas

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Posted 27 Aug 2011 — by James Street
Category gold nanorod antennas, Hyperthermia, NanoTechnology
  1. Geoffrey von Maltzahn1,
  2. Ji-Ho Park2,
  3. Amit Agrawal1,
  4. Nanda Kishor Bandaru3,
  5. Sarit K. Das3,
  6. Michael J. Sailor2, and
  7. Sangeeta N. Bhatia1,4

+ Author Affiliations


  1. 1Harvard-Massachusetts Institute of Technology Division of Health Sciences and Technology, Cambridge, Massachusetts; 2Materials Science and Engineering Program, Department of Chemistry and Biochemistry, University of California, San Diego, La Jolla, California; 3Department of Mechanical Engineering, Indian Institute of Technology Madras, Chennai, India; and 4Howard Hughes Medical Institute and Electrical Engineering and Computer Science, Massachusetts Institute of Technology/Brigham and Women’s Hospital, Boston, Massachusetts
  1. Requests for reprints:
    Sangeeta N. Bhatia, Massachusetts Institute of Technology, 77 Massachusetts Avenue, E19-502D, Cambridge, MA 02139. Phone: 617-324-0221; Fax: 617-324-0740; E-mail: sbhatia@mit.edu.

Abstract

Plasmonic nanomaterials have the opportunity to considerably improve the specificity of cancer ablation by i.v. homing to tumors and acting as antennas for accepting externally applied energy. Here, we describe an integrated approach to improved plasmonic therapy composed of multimodal nanomaterial optimization and computational irradiation protocol development. We synthesized polyethylene glycol (PEG)–protected gold nanorods (NR) that exhibit superior spectral bandwidth, photothermal heat generation per gram of gold, and circulation half-life in vivo (t1/2, ∼17 hours) compared with the prototypical tunable plasmonic particles, gold nanoshells, as well as ∼2-fold higher X-ray absorption than a clinical iodine contrast agent. After intratumoral or i.v. administration, we fuse PEG-NR biodistribution data derived via noninvasive X-ray computed tomography or ex vivo spectrometry, respectively, with four-dimensional computational heat transport modeling to predict photothermal heating during irradiation. In computationally driven pilot therapeutic studies, we show that a single i.v. injection of PEG-NRs enabled destruction of all irradiated human xenograft tumors in mice. These studies highlight the potential of integrating computational therapy design with nanotherapeutic development for ultraselective tumor ablation. [Cancer Res 2009;69(9):3892–900]

Introduction

The electromagnetic properties of plasmonic nanomaterials have been harnessed to develop ultrasensitive diagnostic ( 1, 2), spectroscopic ( 3, 4), and, recently, therapeutic technologies ( 58). In particular, tunable plasmonic nanomaterials have attracted attention for their immense optical absorption coefficients and potential as injectable nanoantennas that target tumors and locally convert otherwise benign electromagnetic energy to thermal energy for ablation. Currently, tumor ablation approaches in clinical practice, including radio frequency, laser, and focused ultrasound methods, lack intrinsic tumor specificity to energy absorption. The inability to selectively heat tumor tissues over surrounding compartments necessitates efforts to externally direct applied energy toward tumor tissues, making effective treatment of tumor margins and complex tumor geometries very challenging. By providing a tumor-specific heat source, nanoantennas could considerably broaden the clinical applicability of thermal therapies by simplifying their integration with current therapeutic practices (including improving margin clearance in surgery and synergizing with regional radiation therapies) and reducing morbidity due to off-target heating. Furthermore, by pulsing the external energy source used, tumor-targeted nanoantennas can theoretically ablate with single-cell precision, thereby providing improved accuracy over standard surgical methods and opening the possibility of precisely treating complex tumor margins in sensitive tissues.

To date, preparations of gold nanoshells and nanorods (NR) have shown considerable efficacy for tumor ablation using NIR light ( 5, 6, 9, 10), with the most recent data showing complete resorption of ∼55% and ∼25% of irradiated tumors, respectively ( 11, 12). These results highlight the clinical promise of these technologies and also motivate the further development of superior nanomaterials and improved methods for optimizing irradiation regimens, which could synergistically improve photothermal therapies. From a material perspective, the development of nanoantennas with enhanced circulation times in vivo, increased absorption coefficients per weight, and narrower absorption spectra would more efficiently permeate into tumors after i.v. administration, amplify the photothermal contrast between antennas and normal tissue, and allow improved tumor treatment at lower laser intensities or at greater depths in vivo. From a procedural perspective, despite a rich history of heat transfer modeling in the hyperthermia field, the use of quantitative modeling to predict the in vivo function of plasmonic nanomaterials has widely remained absent from their testing ( 5, 6, 9, 11, 12). Because the efficacy of photothermal therapy is driven by both the potency of nanoantenna absorption in tumors and the dose of near-IR irradiation, translation of plasmonic materials to effective clinical use will benefit from cohesive integration of biodistribution data with photothermal modeling to predict and customize the four-dimensional irradiated temperature profiles in tumors.

Recently, rod-shaped gold nanoparticles have emerged as precisely tunable plasmonic nanomaterials that may be synthesized in bulk, have narrow size distributions, optical absorption coefficients 104-fold to 106-fold higher than conventional organic fluorochromes, and theoretical per micron absorption coefficients exceeding those of NIR-absorbing gold nanoshells ( 1315). The long precedence of gold nanoparticles in clinical rheumatoid arthritis therapies make gold NRs appealing new candidates for nanoantenna-based photothermal ablation and a wide array of other biomedical applications. Already, gold NRs have been used for a diversity of biological purposes, including multiplexed in vitro detection ( 16), two-photon fluorescence imaging ( 17), and photothermal heating of tumor and bacterial cell targets ( 7, 8, 12, 1820). In addition to their plasmon resonance, the larger atomic number and high material density of gold nanomaterials (z = 79, ρ = 19.3 g/cm3) compared with clinical formulations of iodine-based reagents (z = 53) have attracted interest for X-ray computed tomography (CT) angiography and a few spherical nanoparticle reagents have been developed for in vivo use ( 21, 22).

In this report, we describe the development of polyethylene glycol (PEG)–coated gold NRs as highly absorbing nanoantennas for photothermal tumor destruction under the guidance of biodistribution-based photothermal modeling. We chose a PEG polymer coating due to the widespread clinical use of variable-length PEG polymers for extending the circulation time of protein therapeutics ( 23, 24) and for nanoparticle formulations, such as the drug-loaded liposomes Doxil. We find that PEG-NRs are highly stable, relatively noncytotoxic in vitro, superior photothermal antennas compared with a gold nanoshell preparation in vitro and are improved X-ray absorbing agents compared with clinical iodine standard. After i.v. administration, we find PEG-NRs to be among the longest circulating inorganic nanomaterials described to date (t1/2, ∼17 hours) allowing passive accumulation into xenograft tumors. Using four-dimensional biodistribution-based heat transfer simulations, we designed a therapeutic irradiation regimen that was able to fully destroy all irradiated tumors on mice injected with PEG-NRs using half the light intensity of previous nanoshell therapies.

Materials and Methods

Preparation of PEG-coated gold NRs. Highly stable, ∼13 × 47 nm ( Fig. 1A ) cetyltrimethylammonium bromide (CTAB)–coated gold NRs with longitudinal plasmon resonance at 810 nm (Nanopartz, a division of Concurrent Analytical, Inc.) were centrifuged at 16,000 rcf to concentrate and gently resuspended in 250 μmol/L 5-kDa methyl-PEG-thiol (Laysan Bio, Inc.). Thiol activity of polymers was quantified spectraphotometrically using 5,5-dithiobis(2-nitrobenzoic acid) (Sigma) against a DTT (Sigma) gradient to verify that polymer stocks were fully reduced. The solution of 5-kDa methyl-PEG-thiol and CTAB-coated gold NRs was gently mixed at room temperature for 1 h and dialyzed exhaustively against ultrapure water (18 MΩ cm−1) via cellulose ester membrane dialysis to drive PEG addition (Spectrapor). Dialyzed samples were filtered through 100-kDa filters (Millipore) to remove excess polymer and stored at 4°C. To quantify the number of polymers per particle, NRs were coated as described with an amino-PEG-thiol polymer. After dialysis and extensive filtration on 100-kDa centrifugal filters, amino NRs were harvested and an SPDP assay was performed to quantify the number of amines ( 25).

Figure 1.

Structure and synthesis of highly absorbing, PEG-protected gold NRs. A, near-IR absorbing (810 nm longitudinal plasmon resonance peak) gold NRs were imaged via transmission electron microscopy. B, schematic of process to drive CTAB-NR conversion to PEG-NRs under dialysis with rendering and molecular schematic of PEG coating on NR surface. C, PEG-NRs show prolonged stability in biological media (>1,000 h), whereas CTAB-coated NRs precipitated over time.

Stability and cytotoxicity. Solutions of PEG-NRs or CTAB-NRs (∼60 μg Au/mL) were normalized and incubated in PBS or 10% human serum for extended periods of time. At regular intervals, samples were spectrophotometrically analyzed for plasmon resonance peak shifts, which would indicate particle aggregation. To assess material toxicity, micropatterned primary hepatocyte/human fibroblast cocultures were prepared as described previously ( 26). At 24 h after liver coculture seeding, coculture wells were exposed in triplicate to a gradient of PEG-NRs (0–280 μg Au/mL) and allowed to incubate for 24 h. At this point, PEG-NRs were removed, cells were washed repeatedly, and viability was assessed via a 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide viability assay (Sigma), a colorimetric viability assay for mitochondrial dehydrogenase enzyme activity read using a spectrophotometer (SpectraMax, Molecular Devices).

Photothermal comparison between PEG-NRs and PEG-nanoshells. Gold nanoshells (Nanocomposix) with 120 nm silica core and ∼15 nm gold shell were mixed with methyl-PEG-thiol as described ( 5, 6, 9). Both PEG-nanoshells and PEG-NRs were brought to 7 μg Au/mL, as determined by inductively coupled plasma mass spectrometry (ICP-MS). Tubes containing 200 μL of these solutions were broadly irradiated, under identical conditions, by an 810-nm diode laser (RPMC Lasers, Inc.) at 2 W/cm2. During irradiation, an IR thermographic camera (FLIR, Thermacam S60) was used to measure peak sample surface temperature. To assess photothermal cell toxicity, MDA-MB-435 human cancer cells [American Type Culture Collection (ATCC)] were cultured in a 96-well microplate and grown to 80% confluency using ATCC-recommended media. Cells were incubated with either PEG-NRs (14 μg/mL), PEG-nanoshells (14 μg/mL), or media alone. For each, triplicate wells were exposed to the diode laser light (5 min, 2 W/cm2) or no laser. After treatment, cells were incubated with Calcein AM (5 μg/mL in culture medium; 1 h incubation, Invitrogen), a fluorescent indicator of esterase activity in viable cells and imaged using phase and fluorescence microscopy.

X-ray CT of PEG-NRs. PEG-NRs were suspended in PBS after concentration via membrane centrifugation and serially diluted over a 1,000-fold concentration gradient. A clinical iodine standard was similarly diluted for comparison (Isovue-370). X-ray CT was performed in a GE eXplore Locus microCT scanner (80 kV, 450 μA, 45-μm resolution). For in vivo imaging, mice were imaged before NR injection to reveal the baseline level of soft tissue X-ray contrast. Ten microliters of PEG-NRs (∼3 pmol) were injected with a 30-gauge needle directly into the center of the tumor, and the needle was maintained in place for ∼10 s to allow the tumor to accommodate the additional fluid. After intratumoral administration, mice were imaged and irradiated (∼0.75 W/cm2, 810 nm). X-ray CT images were exported as DICOM files for exportation into modeling software (see Supporting Materials and Methods).

ICP-MS. Samples for ICP-MS (Thermo-Scientific Finnigan ELEMENT2) analysis were frozen, lyophilized, and dissolved in aqua regia, prepared by adding 100 μL of nitric acid + 300 μL of 37% hydrochloric acid for 72 h to dissolve gold particles. Then, samples were diluted to 10 mL with 9.6-mL 2% HNO3 and analyzed via ICP-MS against standards. Control saline and organ samples with exogenously added PEG-NRs were used to calibrate this method.

In vivo circulation time and biodistribution of PEG-NRs. Nude mice were bilaterally injected s.c. in the hind flank with ∼2 × 106 MDA-MB-435 cells. After 2 to 3 wk, animals were anaesthetized with isoflurane and injected through the tail vein with PEG-NRs in 0.15 mol/L NaCl, 0.1 mol/L Na phosphate buffer (pH 7.2; 20 mg Au/kg). 10 μL blood samples were withdrawn periodically from the suborbital space, diluted with PBS containing 10 mmol/L EDTA, centrifuged to remove RBCs, and read on a spectrophotometer to assess PEG-NR plasmon peak height. For biodistribution experiments, after vascular clearance of PEG-NRs (72 h), injected animals were euthanized and organs were collected, weighed, and lyophilized for ICP-MS quantification of PEG-NR biodistribution.

In vivo photothermal heating of gold NRs and photothermal therapy. For both initial modeling and growth curve assessments after photothermal treatment, nude mice were bilaterally injected in the hind flank with ∼2 × 106 MDA-MB-435 cells. After 2 to 3 wk, animals were anaesthetized with isoflurane and injected through the tail vein with PEG-NRs in PBS (20 mg Au/kg) or PBS alone. After vascular clearance of PEG-NRs (72 h), the right flank of each mouse was irradiated (2 W/cm2, 810 nm, 1 cm beam diameter). Thermographic imaging of photothermal heating was used to facilitate modeling of three-dimensional temperature distributions in tumors (n = 3 mice for each set). To explore the hematologic effects of NR administration and near-IR ablation, mice bearing bilateral MDA-MB-435 tumors were injected with saline or PEG-NRs and, 72 h later, either exposed to the therapeutic tumor irradiation protocol under anesthesia (∼2 W/cm2, 5 min, 810 nm) or anesthesia without irradiation (n = 3 each set). After exposure, blood was collected for hematology and mice were sacrificed. Therapeutic assessment of the affect of PEG-NR heating on tumor growth was conducted similarly (n = 4 mice in each treatment set). Both irradiated and unirradiated tumors of each mouse in the therapeutic assessment trial were measured at regular intervals using digital calipers. Survival studies were conducted using mice that were unilaterally injected in the hind flank with ∼2 × 106 MDA-MB-435 cells (n = 5 mice in each treatment set). Tumor sizes were measured over time, and mice were euthanized once tumors exceeded 500 mm3.

Results

Development of ultrastable, polymer-coated gold NRs. In principle, gold nanoparticles are attractive for medical applications because various formulations have been approved by the U.S. Food and Drug Administration and in clinical use for decades. However, one barrier facing the widespread biological use of gold NRs has been the need to replace the cationic CTAB surfactants used to drive their anisotropic growth with hydrophilic, biocompatible coatings. We found that NRs, with axial sizes of 12.7 nm (±3.4) and 47 nm (±9.3; Fig. 1A), coated in CTAB are cytotoxic and colloidally unstable in 0.15 mol/L NaCl or 10% human serum ( Fig. 1B and C). After PEGylation ( Fig. 1B), gold NRs contained ∼20,000 polymers per particle by the SPDP assay ( 25) and were rendered highly stable in vitro, showing minimal spectral shifting (which would indicate particle destabilization and aggregation) even after >1,000 hours in 0.15 mol/L NaCl or 10% human serum ( Fig. 1C and Supplementary Fig. S1A). Additionally, PEG-NRs could be dispersed in a variety of solvents, including acetone, acetonitrile, DMSO, DMF, ethanol, and methanol, further highlighting the stability of their coating and their amenability to future chemical processing and functionalization (Supplementary Fig. S1B). After 24 hours of incubation above primary rat hepatocyte cocultures, an in vitro liver model that was previously used to elucidate semiconductor quantum dot toxicity ( 26) and has shown utility for rank-ordering pharmacologic toxicities ( 27), PEG-NRs, displayed no significant alterations in mitochondrial activity compared with saline alone, even at doses 20 times greater than that used over cells in vitro here and equal to maximal blood concentrations during in vivo therapy experiments (Supplementary Fig. S2), highlighting their potential biocompatibility for medical applications.

Photothermal comparison of gold NRs and gold nanoshells. In light of the prior utility of gold nanoshells for photothermal tumor therapy, they were used as a benchmark here to evaluate the photothermal performance of PEG-NRs ( Fig. 2A ): PEGylated nanoshell preparations with similar composition and spectra to those used for photothermal in vivo applications (refs. 5, 6, 9; ∼120 nm silica cores and ∼15 nm gold shells, 810 nm peak plasmon resonance; Fig. 2B). PEG-NRs exhibited <1/3 of the spectral bandwidth and ∼3 times higher extinction coefficient per gram gold than PEG-nanoshells (full width at half maximum of PEG-NRs 130 nm, PEG-nanoshells 490 nm; Fig. 2B). Additionally, under identical experimental conditions, irradiated PEG-NR solutions generated heat >6 times more rapidly than PEG-nanoshells per gram gold ( Fig. 2C). These findings are consistent with theoretical calculations indicating that gold NRs of this size exhibit superior per micron absorption coefficients to gold nanoshells ( 14). Accordingly, incubation of PEG-NRs with MDA-MB-435 human tumor cells in vitro enabled their widespread destruction with levels of NIR light that were insufficient for nanoshell-mediated toxicity ( Fig. 2D). Exposure to individual stimuli (NRs, nanoshells, or laser) did not affect cell viability (Supplementary Fig. S3).

Figure 2.

Spectral and photothermal properties of highly absorbing gold NRs compared with gold nanoshells. A, schematic of photothermal heating of gold NRs. The dimensions of gold NRs are tuned to have a near-IR plasmon resonance, at which point nanoparticle electrons resonantly oscillate and dissipate energy as heat. B, spectra for PEG-gold NRs (red) and PEG-gold nanoshells (blue), a benchmark for tunable plasmonic nanomaterials, at equal gold concentrations. C, top, rate of temperature increase for triplicate PEG-NR and PEG-gold nanoshell solutions (7 μg Au/mL, 810 nm laser, 2 W/cm2, n = 3 each). Bottom, IR thermographic image of PEG-NRs versus PEG-gold nanoshells after 2 min of irradiation. Scale bar, 5 mm. D, in vitro photothermal toxicity of PEG-NRs over human cancer cells in culture (MDA-MB-435). Tumor cells were incubated with PEG-NRs (14 μg/mL; top), PEG-nanoshells (14 μg/mL; middle), or media alone (bottom) and treated with laser irradiation (2 W/cm2, 810 nm, 5 min). Calcein AM staining indicates destruction of cells with PEG-NRs, whereas cells irradiated in the presence of nanoshells or media remained viable. Phase region of calcein staining inset. Scale bar, 10 μm.

X-ray CT and computational modeling of photothermal NR heating in vivo. To translate the photothermal efficacy of PEG-NRs to in vivo therapy, we next developed a method through which PEG-NRs could be heated under the guidance of biodistribution-based computational photothermal modeling. To model and customize patient irradiation regimens, we sought to acquire nanoparticle distributions remotely using X-ray CT, a desirable imaging modality due to its high three-dimensional anatomic resolution, rapid imaging speed, quantitative dynamic range of detection, full body penetration, and ubiquitous clinical use. To investigate the ability of PEG-NRs to act as dense X-ray absorbing agents for X-ray CT, solutions of PEG-NRs were serially diluted and imaged using a GE microCT scanner ( Fig. 3A ). The resulting relationship between PEG-NR concentration and X-ray contrast was highly linear and exhibited ∼2-fold amplified X-ray contrast compared with a clinical iodine standard per mole ( Fig. 3B and Supplementary Fig. S4), analogous to that found previously for spherical gold nanoparticle reagents ( 21, 22). In addition to providing enhanced X-ray absorption, PEG-NRs allow NIR photothermal heating whereas iodine reagents show no heating above water alone (Supplementary Fig. S4).

Figure 3.

X-ray CT, quantitative photothermal modeling, and near-IR photothermal heating of gold NRs in vivo. A, schematic of X-ray absorption by gold NRs in X-ray CT. B, X-ray CT number of PEG-NRs compared with an iodine standard (Isovue-370). C, PEG-NRs were intratumorally given to mice bearing bilateral MDA-MB-435 tumors and imaged using X-ray CT to visualize three-dimensional PEG-NR distribution in tumors (left). A three-dimensional solid model of the complete geometry was rapidly reconstructed by image processing for use with computational photothermal modeling (middle). Red, PEG-NRs. Experimental thermographic surveillance of NIR irradiation after X-ray CT (∼0.75 W/cm2, 1 min; right). D, meshed geometry of the left tumor chosen as the computational domain (left). Plot of theoretical heat flux propagation inside the tumor upon irradiation (middle left). Predicted internal temperature distribution at three different planes inside the tumor (middle right) along with surface temperature map (right) matching the left tumor in C.

To assess the utility of the high X-ray absorption of PEG-NRs for remote detection in vivo, ∼3 pmol of PEG-NRs (∼1 μmol Au) were directly injected into the tumors of mice bearing bilateral human MDA-MB-435 tumors, implanted either in the mammary fat pad or the rear flank. We found that X-ray CT rapidly detailed the three-dimensional distribution of PEG-NRs in tumors, showing clear distinction between NRs and soft tissues ( Fig. 3C and Supplementary Figs. S5 and S6). To understand the relationship between the nanoparticle distribution in tumors and the associated processes of thermal deposition and flux during irradiation, a finite element model of PEG-NR heating was developed based on the Pennes bioheat transfer equation, including (a) three-dimensional skeletal, soft tissue, and PEG-NR geometries transferred from X-ray CT imaging; (b) temperature-dependent optical absorption and scattering; (c) heat transfer, including surface thermal flux, internal diffusive flux, and temperature-dependent perfusive thermal flux in tissues; and (d) an extracorporeal NIR laser of variable intensity, beam diameter, and orientation matching used (see supplementary data). X-ray CT data was exported into ScanIP and ScanFE image processing software for skeletal, NR, and soft tissue geometry construction and subsequently into COMSOL Multiphysics modeling software for photothermal simulations. Exported geometries enabled rapid delineation PEG-NRs, along with skeletal structures and surrounding soft tissues for spatially defining model parameters ( Fig. 3C and Supplementary Figs. S5 and S6). Computational, finite element heat transfer simulations were performed using PEG-NRs and tumor geometries as computational domains to evaluate the feasibility of applying four-dimensional modeling to the process of photothermal heating under irradiation ( Fig. 3D and Supplementary Fig. S6). Simulations revealed the intratumoral vectors of thermal flux extending from regions of PEG-NRs, as well as the internal and surface temperature history expected for irradiation at varying laser intensities ( Fig. 3D and Supplementary Fig. S6). At matched computational and experimental laser intensities, simulated surface heating behavior qualitatively and quantitatively matched the observed surface temperature values and distributions acquired using IR thermographic observation ( Fig. 3D and Supplementary Fig. S6). We believe this establishes the potential of fusing quantitative imaging with computational modeling to provide insight into the unintuitive, highly complex phenomena of in vivo photothermal heating. Next, we proceeded to explore the power of this modeling to quantitatively predict in vivo heating after i.v. tumor targeting.

Long circulation time and photothermal tumor heating after i.v. NR administration. We next investigated the behavior of PEG-NRs after i.v. administration in mice to systemically target tumors through the enhanced permeability and retention effect ( 28). For PEG-NRs to passively target tumors via the enhanced permeability and retention effect and act as nanoantennas for photothermal therapy, they must be able to traverse the systemic circulation, deter protein opsonization and reticuloendothelial system (RES) clearance, permeate through transendothelial pores in tumor blood vessels, and be retained in the tumor interstitium. After i.v. administration to tumor-bearing mice (20 mg Au/kg), our PEG-NRs were found to exhibit blood half-lives of ∼17 hours ( Fig. 4A ) and to maintain their 810-nm longitudinal plasmon resonance throughout this time, allowing spectrophotometric detection in serum over time (Supplementary Fig. S7).

Figure 4.

Long circulation time, passive tumor targeting, and photothermal heating of passively targeted gold NR antennas in tumors. A, PEG-NRs were i.v. given (20 mg/kg) to three mice bearing MDA-MB-435 tumors, and blood was withdrawn over time to monitor clearance from circulation. B, PEG-NR biodistribution and targeting to MDA-MB-435 tumors 72 h after i.v. administration, quantified via ICP-MS (three mice). T, tumor; Br, brain; Bl, bladder; M, muscle; H, heart; Lu, lung; K, kidney; Li, liver; SP, spleen. Data are tabulated in Supplementary Table S1. C, PEG-NRs or saline were i.v. given (20 mg/kg) to mice bearing MDA-MB-435 tumors on opposing flanks. After NRs had cleared from circulation (72 h after injection), the right flank was irradiated using an 810-nm diode laser (2 W/cm2; beam size indicated by dotted circle). D, thermographic surveillance of photothermal heating in PEG-NR–injected (top) and saline-injected (bottom) mice.

To quantitatively assess tumor accumulation of PEG-NRs, nude mice bearing MDA-MB-435 human tumors were given i.v. PEG-NRs, and once NRs had cleared from circulation (72 hours), organs were removed for gold quantitation. Here, ICP-MS was used to quantify the accumulation of exogenously given gold in tissues. ICP-MS Au NR detection was found to be highly sensitive and linear across a relevant range for gold NR detection (Supplementary Fig. S8). As expected for nanomaterials above the renal filtration limit, PEG-NRs were found to clear via RES uptake with splenic clearance dominating hepatic ( Fig. 4B; Supplementary Table S1), a pattern analogous to that observed previously for PEG-protected 10-nm diameter spherical gold nanoparticles ( 21). Importantly, passive tumor accumulation of PEG-NRs after injection was found to be highly efficient, even at 72 hours after injection, accumulating at ∼7% ID/g ( Fig. 4B; Supplementary Table S1), allowing PEG-NRs to amplify the tumor absorption coefficient of 810 nm light by ∼7 fold (see supplementary data). Based on empirical studies of passive tumor targeting, the enhancement of the PEG-NR circulation time over reported PEG-nanoshell formulations should amplify their passive tumor accumulation in tumors, which, in concert with their enhanced photothermal heat generation in vitro, would be expected to provide overall enhanced photothermal contrast between tumors and normal tissue. However, as PEG-nanoshells are not commercially available, a side-by-side in vivo comparison could not be pursued. Separately, to study the long-term clearance of PEG-NRs from RES organs, tumor-free mice were injected with PEG-NRs and sacrificed 2 months after injection. During these 2 months, injected mice showed no signs over NR toxicity, such as weakness, belabored breathing, or failure to thrive. Organ analysis showed that the %ID/g values for tissues decreased by >50% in all organs and by >80% in muscle, heart, lungs, and kidneys (Supplementary Fig. S9), indicating the existence of routes for gold NR clearance after uptake in tissues.

Having observed their efficient passive homing to human tumors via the enhanced permeability and retention effect, the ability of PEG-NRs to remain active as optical nanoantennas for photothermal tumor heating after clearance from the systemic circulation (∼72 hours) was subsequently investigated. Either PEG-NRs (20 mg Au/kg) or a saline solution was given i.v. into mice bearing two MDA-MB-435 tumors on opposite flanks. Once PEG-NRs had cleared from circulation (72 hours), irradiated tumors on PEG-NR–injected mice rapidly heated to temperatures of over 70°C ( Fig. 4C), whereas saline-injected mice displayed less focal temperature increases with maximum surface temperatures of ∼40°C ( Fig. 4D). To inform the development of near-IR radiation doses that would fully destroy tumors, photothermal heating simulations were again conducted using the i.v. biodistribution data and computational domains containing ellipsoidal tumor geometries matching those used in experiments (Supplementary Fig. S10). Photothermal heating simulations closely matched experimental surface temperature data ( Figs. 4D and 5A and B ), suggesting that tumor-targeted nanoantennas substantially retained their photothermal efficacy during the 72 hours after injection in vivo. Furthermore, the simulations predicted that the entire tumor volume, including the deepest tumor/tissue interface, would be heated to ablative temperatures (>60°C) by 5 minutes after the onset of laser irradiation ( Fig. 5C). Thus, a 5-minute, 2-W/cm2 irradiation regimen was used for subsequent therapeutic experiments in an effort to provide fully destructive photothermal tumor treatment.

Figure 5.

Quantitative photothermal modeling of gold NR tumor heating. A, three-dimensional finite element modeling of PEG-NR heating in vivo. Simulated three-dimensional temperature distributions matching the four-dimensional thermographic time points for PEG-NR (top) and control tumor irradiation (bottom). B, thermographically measured and simulated tumor surface temperatures over time for irradiation of PEG-NR or saline mice. C, simulated temperature increases various depths for PEG-NR–injected and saline-injected mice. By 5 min after the onset of irradiation, the entire PEG-NR tumor is predicted to have reached ablative temperatures of >60°C, motivating the choice of this irradiation regimen for subsequent therapeutic trials.

Photothermal tumor destruction using a computationally designed irradiation regimen. To test the prediction that a single dose of PEG-NRs could destroy tumors under the computationally designed protocol of NIR irradiation, nude mice bearing bilateral human MDA-MB-435 tumors were injected with either PEG-NRs or saline. After i.v. clearance of PEG-NRs, the right flank of each mouse was irradiated for 5 min (810 nm, 2 W/cm2) and all tumors were measured at regular intervals over time. Within 10 days all the irradiated, PEG-NR–targeted tumors completely disappeared by external observation whereas all other tumors, including those exposed to laser after saline injection, continued to grow uninhibited ( Fig. 6A ). To assess the survival benefit of PEG-NR–directed tumor ablation, mice bearing a single MDA-MB-435 tumor were divided between four cohorts (PEG-NRs + laser, PEG-NRs − laser, saline + laser, saline − laser) and all tumors were measured over time ( Fig. 6B). By 20 days after treatment, all irradiated, PEG-NR–injected mice displayed only a minor scar with no evidence of tumor regrowth, whereas all other surviving mice harbored thriving tumors ( Fig. 6B and C). Over the course of >50 days of observation, no irradiated, PEG-NR–injected mice showed evidence of recurrence whereas all mice in the control had to be euthanized by day 33. Body weights of PEG-NR–treated mice were monitored over time and showed no obvious losses due to tumor ablation and resorption (Supplementary Fig. S11). In a separate experiment to assess the acute hematologic effects of NR-directed tumor ablation, the only statistically significant change observed in response to NR-mediated tumor ablation was a slight increase in the percentage of band neutrophils in NR + laser sets (P < 0.05 for NR+ laser versus NR, saline + laser, and saline; Supplementary Fig. S12), likely due to an acute inflammatory response to tumor ablation.

Figure 6.

Photothermal destruction of human tumors in mice using long-circulating gold NRs. A, mice harboring two MDA-MB-435 human tumors on opposite flanks were injected with either saline or PEG-NRs. After PEG-NRs had cleared from circulation (72 h after injection), the right flank of each mouse was exposed to the computationally designed irradiation regimen (810 nm, 2 W/cm2, 5 min). Volumetric changes in tumor sizes are plotted over time after irradiation. B, mice harboring one MDA-MB-435 human tumor were injected with either saline or PEG-NRs and irradiated as in A. Survival of mice after irradiation is plotted versus time after irradiation. C, at 20 d after irradiation, NIR-irradiated, all PEG-NR–injected mice showed only a minor scar and no evidence of tumor regrowth whereas all other treatment groups harbored thriving tumors.

Discussion

Here, we present the development of an integrated system for nanoantenna-based photothermal tumor therapy involving the synthesis of long-circulating gold NRs as efficient NIR-nanoantennas, biodistribution data acquisition via X-ray CT nanomaterial imaging or ex vivo spectrometry, and photothermal computational modeling to guide surgical irradiation planning. Broadly, the efficacy of a nanoantenna for photothermal therapy depends on both intrinsic (optical absorption coefficient and material cytotoxicity) and extrinsic (polymer coating, macrophage affinity, and circulation time) material characteristics, as well as external parameters, such as the use of optimized dosing and irradiation protocols for effective treatment.

We show that PEG-NRs exhibit superior intrinsic absorption and photothermal efficacy compared with gold nanoshells (∼6 times greater heat generation per weight gold), as well as substantially improved circulation times in vivo (∼17 hours versus ∼4 hours), extrinsically imparted by their polymer coating ( 5, 6). Surveying literature on inorganic nanoparticle circulation times in vivo, the circulation half-life of PEG-NRs is among the longest achieved to date. Previously, polymer-stabilized inorganic nanomaterials have been described with circulation half-lives of a few hours in vivo ( 29, 30), including a variety of other gold nanoparticle preparations ( 22, 3133), and on occasion with circulation times of ∼10 to 15 hours in mice ( 21, 28, 34). Elsewhere, another PEG-NR formulation was developed for in vivo applications, but showed a 30-minute half-life without investigation into their ability to passively target tumors or mediate in vivo photothermal heating ( 32). Because nanoparticle circulation time has been shown to determine the efficiency of nanoparticle accumulation in tumors via the enhanced permeability and retention effect, in mouse cancer models and clinical cancer treatment ( 28), the long circulation time reported here has the potential to directly translate to improved clinical tumor accumulation over previous nanoantennas.

Beyond the material determinants of nanoantenna efficacy, the irradiation protocol used (i.e., beam intensity, shape, cross-section, duration, direction, etc.) and nanoantenna dosing regimen directly control the rates of energy capture and dissipation to surrounding tissues in vivo. Whereas nanoantennas have the potential to increase the selectivity of tumor ablation, unoptimized irradiation of tissues carries the risks of either unnecessary morbidity due to collateral damage or ineffective therapy due to incomplete treatment of tumor margins. Here, we show that quantitative biodistribution data incorporated into computational modeling can help anticipate the photothermal heating in tumors and surrounding tissues during irradiation. Future developments of the quantitative model presented here could enable rapid quantitative modeling of photothermal temperature gradients in arbitrarily complex three-dimensional tissues and provide a route toward a priori personalization of irradiation regimens. As a proof of principle, we establish a means of integrating whole-subject X-ray CT data with quantitative heat transfer modeling, offering a new route toward merging the clinical paradigms of imaging and therapy for personalized four-dimensional radiation planning and optimization. Furthermore, using a computationally planned therapeutic method, we show that i.v. administration of PEG-NR nanoantennas enabled complete destruction of all irradiated tumors under otherwise benign near-IR light.

We believe our findings motivate future investigation into the long-term biodistribution of PEG-NRs, more extensive analysis of their potential toxicity in vivo, and the development of methods for detecting low concentrations of PEG-NRs in whole animals to remotely quantify i.v. tumor targeting. Methods for actively targeting NRs to tumors, particularly to vascular epitopes, could potentially enhance their specificity for tumors or direct their additional accumulation in premalignant lesions and metastatic lymphatics. Finally, we provide clear evidence that the application of quantitative biodistribution-based modeling to the in vivo testing of nanomaterials can provide insight into their function and direct procedural optimization.

Disclosure of Potential Conflicts of Interest

G. von Maltzahn: Consultant and ownership interest, Concurrent Analytical. The other authors declared no potential conflicts of interest.

Acknowledgments

Grant support: NIH grant BRP R01CA124427-01, NIH/National Cancer Institute grants U54 CA119349 and U54 CA119335, Packard Fellowship (1999-1453), and Whitaker Foundation and National Science Foundation (G. von Maltzahn).

The costs of publication of this article were defrayed in part by the payment of page charges. This article must therefore be hereby marked advertisement in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.

We thank Dr. Shelley J. Coldiron and Dr. Christian Schoen at Nanopartz for developing the CTAB nanorods used in this work, Dr. Yoel Fink for generously lending the FLIR IR thermographic camera, and Dr. Eugene Zubarev and Bishnu Khanal at Rice University for their synthesis and characterization of CTAB-coated gold nanorods through Nanopartz.

Footnotes

  • Received November 5, 2008.
  • Revision received February 5, 2009.
  • Accepted February 5, 2009.

References

  1. Elghanian R, Storhoff JJ, Mucic RC, Letsinger RL, Mirkin CA. Selective colorimetric detection of polynucleotides based on the distance-dependent optical properties of gold nanoparticles. Science 1997; 277: 1078–81.
  2. Grubisha DS, Lipert RJ, Park HY, Driskell J, Porter MD. Femtomolar detection of prostate-specific antigen: an immunoassay based on surface-enhanced Raman scattering and immunogold labels. Anal Chem 2003; 75: 5936–43.
  3. Jackson JB, Westcott SL, Hirsch LR, West JL, Halas NJ. Controlling the surface enhanced Raman effect via the nanoshell geometry. Appl Phys Lett 2003; 82: 257–9.
  4. Qian XM, Peng XH, Ansari DO, et al. In vivo tumor targeting and spectroscopic detection with surface-enhanced Raman nanoparticle tags. Nat Biotechnol 2008; 26: 83–90.
  5. Hirsch LR, Stafford RJ, Bankson JA, et al. Nanoshell-mediated near-infrared thermal therapy of tumors under magnetic resonance guidance. P Natl Acad Sci USA 2003; 100: 13549–54.
  6. O’Neal DP, Hirsch LR, Halas NJ, Payne JD, West JL. Photo-thermal tumor ablation in mice using near infrared-absorbing nanoparticles. Cancer Lett 2004; 209: 171–6.
  7. Norman RS, Stone JW, Gole A, Murphy CJ, Sabo-Attwood TL. Targeted photothermal lysis of the pathogenic bacteria, Pseudomonas aeruginosa, with gold nanorods. Nano Lett 2008; 8: 302–6.
  8. Huang X, El-Sayed IH, Qian W, El-Sayed MA. Cancer cell imaging and photothermal therapy in the near-infrared region by using gold nanorods. J Am Chem Soc 2006; 128: 2115–20.
  9. Gobin AM, Lee MH, Halas NJ, James WD, Drezek RA, West JL. Near-infrared resonant nanoshells for combined optical imaging and photothermal cancer therapy. Nano Lett 2007; 7: 1929–34.
  10. Xie H, Gill-Sharp KL, O’Neal P. Quantitative estimation of gold nanoshell concentrations in whole blood using dynamic light scattering. Nanomed Nanotechnol 2007; 3: 89–94.
  11. James WD, Hirsch LR, West JL, O’Neal PD, Payne JD. Application of INAA to the build-up and clearance of gold nanoshells in clinical studies in mice. J Radioanal Nucl Ch 2007; 271: 455–9.
  12. Dickerson EB, Dreaden EC, Huang X, et al. Gold nanorod assisted near-infrared plasmonic photothermal therapy (PPTT) of squamous cell carcinoma in mice. Cancer Lett 2008; 269: 57–66.
  13. Murphy CJ, San TK, Gole AM, et al. Anisotropic metal nanoparticles: synthesis, assembly, and optical applications. J Phys Chem B 2005; 109: 13857–70.
  14. Jain PK, Lee KS, El-Sayed IH, El-Sayed MA. Calculated absorption and scattering properties of gold nanoparticles of different size, shape, and composition: applications in biological imaging and biomedicine. J Phys Chem B 2006; 110: 7238–48.
  15. Hu M, Chen JY, Li ZY, et al. Gold nanostructures: engineering their plasmonic properties for biomedical applications. Chem Soc Rev 2006; 35: 1084–94.
  16. Yu C, Nakshatri H, Irudayaraj J. Identity profiling of cell surface markers by multiplex gold nanorod probes. Nano Lett 2007; 7: 2300–6.
  17. Wang H, Huff TB, Zweifel DA, et al. In vitro and in vivo two-photon luminescence imaging of single gold nanorods. Proc Natl Acad Sci U S A 2005; 102: 15752–6.
  18. Skirtach AG, Karageorgiev P, De Geest BG, Pazos-Perez N, Braun D, Sukhorukov GB. Nanorods as wavelength-selective absorption centers in the visible and near-infrared regions of the electromagnetic spectrum. Adv Mater 2008; 20: 506–10.
  19. Tong L, Zhao Y, Huff TB, Hansen MN, Wei A, Cheng JX. Gold nanorods mediate tumor cell death by compromising membrane integrity. Adv Mater 2007; 19: 3136–41.
  20. Huff TB, Tong L, Zhao Y, Hansen MN, Cheng JX, Wei A. Hyperthermic effects of gold nanorods on tumor cells. Nanomedicine-UK 2007; 2: 125–32.
  21. Cai QY, Kim SH, Choi KS, et al. Colloidal gold nanoparticles as a blood-pool contrast agent for X-ray computed tomography in mice. Invest Radiol 2007; 42: 797–806.
  22. Kim D, Park S, Lee JH, Jeong YY, Jon S. Antibiofouling polymer-coated gold nanoparticles as a contrast agent for in vivo X-ray computed tomography imaging. J Am Chem Soc 2007; 129: 7661–5.
  23. Harris JM, Chess RB. Effect of pegylation on pharmaceuticals. Nat Rev Drug Discov 2003; 2: 214–21.
  24. Duncan R. The dawning era of polymer therapeutics. Nat Rev Drug Discov 2003; 2: 347–60.
  25. Josephson L, Tung CH, Moore A, Weissleder R. High-efficiency intracellular magnetic labeling with novel superparamagnetic-tat peptide conjugates. Bioconjug Chem 1999; 10: 186–91.
  26. Derfus AM, Chan WCW, Bhatia SN. Probing the cytotoxicity of semiconductor quantum dots. Nano Lett 2004; 4: 11–8.
  27. Khetani SR, Bhatia SN. Microscale culture of human liver cells for drug development. Nat Biotechnol 2008; 26: 120–6.
  28. Moghimi SM, Hunter AC, Murray JC. Long-circulating and target-specific nanoparticles: theory to practice. Pharmacol Rev 2001; 53: 283–318.
  29. Ballou B, Lagerholm BC, Ernst LA, Bruchez MP, Waggoner AS. Noninvasive imaging of quantum dots in mice. Bioconjug Chem 2004; 15: 79–86.
  30. Rabin O, Perez JM, Grimm J, Wojtkiewicz G, Weissleder R. An X-ray computed tomography imaging agent based on long-circulating bismuth sulphide nanoparticles. Nat Mater 2006; 5: 118–22.
  31. Paciotti GF, Myer L, Weinreich D, et al. Colloidal gold: a novel nanoparticle vector for tumor directed drug delivery. Drug Deliv 2004; 11: 169–83.
  32. Niidome T, Yamagata M, Okamoto Y, et al. PEG-modified gold nanorods with a stealth character for in vivo applications. J Control Release 2006; 114: 343–7.
  33. Qian X, Peng XH, Ansari DO, et al. In vivo tumor targeting and spectroscopic detection with surface-enhanced Raman nanoparticle tags. Nat Biotechnol 2008; 26: 83–90.
  34. Weissleder R, Bogdanov A, Neuwelt EA, Papisov M. Long-circulating iron-oxides for MR-imaging. Adv Drug Deliver Rev 1995; 16: 321–34.

 

Celsion Announces Publication of Clinical and Scientific Review of ThermoDox(R) as Treatment for Primary Liver Cancer

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Posted 17 Aug 2011 — by James Street
Category doxorubicin, Drug Delivery, Hyperthermia, Liposomes, Liver

press release

Aug. 17, 2011, 8:00 a.m. EDT

COLUMBIA, MD, Aug 17, 2011 (MARKETWIRE via COMTEX) — Celsion Corporation CLSN +1.64% , a leading oncology drug development company, announced today the publication of a clinical and scientific review of ThermoDox(R), the Company’s proprietary heat-activated liposomal encapsulation of doxorubicin, as a treatment for Hepatocellular Carcinoma (HCC or primary liver cancer) in the August 2011 issue of Future Oncology (Volume 7, Number 8). The article, titled “Lyso-Thermosensitive Liposomal Doxorubicin: An Adjuvant to Increase the Cure Rate of Radiofrequency Ablation (RFA) in Liver Cancer,” provides an overview of current standards and investigational approaches to the treatment of HCC, focusing on the curative and synergistic potential of combining lyso-thermosensitive liposomal doxorubicin (LTLD or “ThermoDox(R)”) and RFA as front-line therapy. Ronnie T.P. Poon, MD (QMH), MS, PhD, FRCS (Edin), FACS, Professor of Surgery at the University of Hong Kong and a Lead Asia Pacific Principal Investigator in Celsion’s pivotal Phase III HEAT Study of ThermoDox(R) in primary liver cancer, and Nicholas Borys, MD, Celsion’s Chief Medical Officer, were co-authors of the article, which is available online ( http://www.futuremedicine.com/doi/abs/10.2217/fon.11.73 ).

“The clinical potential of ThermoDox(R) in HCC stems from a number of properties, including the known efficacy and tolerability of doxorubicin in HCC, the enhancement of cell killing when doxorubicin is combined with hyperthermia, localization of ThermoDox(R) in tumors and tumor vasculature and rapid release of drug in the treatment area only when triggered by heat,” said Dr. Poon. “These properties are designed to extend the therapeutic benefit of RFA, a treatment whose efficacy is significantly influenced by size, to larger tumors. If this curative and synergistic potential is borne out in the Phase III HEAT Study, a rational future strategy for larger HCC legions is to employ RFA and ThermoDox(R) as front-line therapy.”

Dr. Borys added: “No more than 30 percent of HCC patients are considered suitable for curative treatment because of tumor size, severity of liver impairment and other factors, leading to a high rate of mortality for this globally epidemic disease. We believe that ThermoDox(R), as an adjuvant therapy that interacts synergistically with RFA, may represent one of the most important new treatment advances for primary liver cancer. Having met the enrollment objective in our Phase III HEAT Study, we remain diligent in our efforts to support the highest level of study execution ahead of a planned interim analysis by an independent Data Monitoring Committee and final data readout.”

The article in Future Oncology details the clinical activity, safety and tolerability, mechanisms of action and pharmacokinetic properties of ThermoDox(R), as well as strategies to improve RFA treatment for HCC. Among the properties highlighted in the article are:

        
        --  ThermoDox(R), as a liposome, rapidly concentrates in the liver and
            spleen. As tumors have much higher microvascular permeability than
            normal tissue, ThermoDox(R) further accumulates in liver tumors;
        --  Hyperthermia has a biological effect of increasing the pore size in
            tumor blood vessels and therefore enhancing the extravasation of
            liposomes into the tumor interstitium;
        --  ThermoDox(R) is over 1,000 times less permeable across normal blood
            vessels than free doxorubicin, offering less potential for systemic
            toxicity;
        --  Hyperthermia has been shown to preferentially increase liposomal
            permeability within the microvasculature in tumor versus normal
            tissue;
        --  The optimal liposome size for heat-induced extravasation was found to
            be 100 nm, the mean diameter of ThermoDox(R).

The article also describes results from a Phase I study of ThermoDox(R) in 24 patients with HCC or liver tumors metastatic from other primary sites. The trial established a statistically significant dose-response effect and a maximum tolerated dose (MTD), 50mg/m(2), for further study. Mean time to treatment failure for patients receiving at least the maximum tolerated dose was 374 days, while that for patients receiving less than 50 mg/m(2) was 80 days. Drug-related adverse events were consistent with the adverse event profile of systemic doxorubicin.

About Primary Liver Cancer

Primary liver cancer is one of the most deadly forms of cancer and ranks as the fifth most common solid tumor cancer. The incidence of primary liver cancer is approximately 20,000 cases per year in the United States, approximately 40,000 cases per year in Europe and is rapidly growing worldwide, now at approximately 700,000 cases per year, due to the high prevalence of Hepatitis B and C in developing countries. The standard first-line treatment for liver cancer is surgical resection of the tumor; however, 90% of patients are ineligible for surgery. Radio frequency ablation (RFA) has increasingly become the standard of care for non-resectable liver tumors, but the treatment becomes less effective for larger tumors. There are few non-surgical therapeutic treatment options available as radiation therapy and chemotherapy are largely ineffective in the treatment of primary liver cancer.

About ThermoDox(R) and the Phase III HEAT Study

ThermoDox(R) is a proprietary heat-activated liposomal encapsulation of doxorubicin, an approved and frequently used oncology drug for the treatment of a wide range of cancers. In the HEAT Study, ThermoDox(R) is administered intravenously in combination with RFA. Localized mild hyperthermia (39.5 – 42 degrees Celsius) created by the RFA releases the entrapped doxorubicin from the liposome. This delivery technology enables high concentrations of doxorubicin to be deposited preferentially in a targeted tumor.

For primary liver cancer, ThermoDox(R) is being evaluated in a 600 patient global Phase III study under an FDA Special Protocol Assessment. The study is designed to evaluate the efficacy of ThermoDox(R) in combination with Radio Frequency Ablation (RFA) when compared to patients who receive RFA alone as the control. The primary endpoint for the study is progression-free survival (PFS) with a secondary confirmatory endpoint of overall survival. A pre-planned, unblinded interim efficacy analysis will be performed by the independent Data Monitoring Committee when 190 PFS events are realized in the study population. Additional information on the Company’s ThermoDox(R) clinical studies may be found at www.clinicaltrials.gov .

About Celsion Corporation

Celsion is a leading oncology company dedicated to the development and commercialization of innovative cancer drugs including tumor-targeting treatments using focused heat energy in combination with heat-activated drug delivery systems. Celsion has research, license, or commercialization agreements with leading institutions such as the National Institutes of Health, Duke University Medical Center, University of Hong Kong, the University of Pisa, and the North Shore Long Island Jewish Health System.

For more information on Celsion, visit our website: http://www.celsion.com .

Celsion wishes to inform readers that forward-looking statements in this release are made pursuant to the “safe harbor” provisions of the Private Securities Litigation Reform Act of 1995. Readers are cautioned that such forward-looking statements involve risks and uncertainties including, without limitation, unforeseen changes in the course of research and development activities and in clinical trials by others; possible acquisitions of other technologies, assets or businesses; possible actions by customers, suppliers, competitors, regulatory authorities; and other risks detailed from time to time in the Company’s periodic reports filed with the Securities and Exchange Commission.

        
        Investor Contact

        David Pitts
        Argot Partners
        212-600-1902
        Email Contact

SOURCE: Celsion Corporation

Tumor Ablation with Irreversible Electroporation

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Posted 09 Aug 2011 — by James Street
Category electromagnetic, General Cancer Research, Hyperthermia, Radiation

Bassim Al-Sakere1,2, Franck André1,2, Claire Bernat1,2, Elisabeth Connault1,2, Paule Opolon1,2, Rafael V. Davalos3, Boris Rubinsky4,5,6, Lluis M. Mir1,2*

1 CNRS UMR 8121, Institut Gustave-Roussy, Villejuif, France, 2 University Paris-Sud, UMR 8121, Villejuif, France, 3 School of Biomedical Engineering and Sciences, Virginia Tech-Wake Forest University, Blacksburg, Virginia, United States of America, 4 Department of Bioengineering, University of California at Berkeley, Berkeley, California, United States of America, 5 Department of Mechanical Engineering and Graduate Program in Biophysics, University of California at Berkeley, Berkeley, California, United States of America, 6 Center for Bioengineering in the Service of Humanity and Society, School of Computer Science and Engineering, Hebrew University of Jerusalem, Givat Ram, Jerusalem, Israel

Abstract Top

We report the first successful use of irreversible electroporation for the minimally invasive treatment of aggressive cutaneous tumors implanted in mice. Irreversible electroporation is a newly developed non-thermal tissue ablation technique in which certain short duration electrical fields are used to permanently permeabilize the cell membrane, presumably through the formation of nanoscale defects in the cell membrane. Mathematical models of the electrical and thermal fields that develop during the application of the pulses were used to design an efficient treatment protocol with minimal heating of the tissue. Tumor regression was confirmed by histological studies which also revealed that it occurred as a direct result of irreversible cell membrane permeabilization. Parametric studies show that the successful outcome of the procedure is related to the applied electric field strength, the total pulse duration as well as the temporal mode of delivery of the pulses. Our best results were obtained using plate electrodes to deliver across the tumor 80 pulses of 100 µs at 0.3 Hz with an electrical field magnitude of 2500 V/cm. These conditions induced complete regression in 12 out of 13 treated tumors, (92%), in the absence of tissue heating. Irreversible electroporation is thus a new effective modality for non-thermal tumor ablation.

Citation: Al-Sakere B, André F, Bernat C, Connault E, Opolon P, et al. (2007) Tumor Ablation with Irreversible Electroporation. PLoS ONE 2(11): e1135. doi:10.1371/journal.pone.0001135

Academic Editor: Mark Isalan, Center for Genomic Regulation, Spain

Received: September 4, 2007; Accepted: October 9, 2007; Published: November 7, 2007

Copyright: © 2007 Al-Sakere et al. This is an open-access article distributed under the terms of the Creative Commons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original author and source are credited.

Funding: This work was supported by grants of CNRS and IGR. BR was supported in part by the U.S. National Institutes of Health (NIH) under Grant NIH R01 RR018961. The sponsors had no role in study design, data collection and analysis, decision to publish, or preparation of the manuscript.

Competing interests: R. Davalos and B. Rubinsky have potential interests in the revenues of pending patents held by Univ. of California at Berkeley (e.g. No 10/571,162, Tissue Ablation with Irreversible Electroporation). B. Rubinsky also has a financial interest in Excellin Life Sciences and in Oncobionic, companies in the field of electroporation. L. M. Mir was a consultant for Oncobionics.

* To whom correspondence should be addressed. E-mail: luismir@igr.fr

Introduction Top

Minimally invasive tissue ablation has become of central importance in the modern surgery armamentarium. In the treatment of benign or malignant tumors it is important to achieve ablation of the undesirable tissue in a well-controlled and precise way without affecting the surrounding healthy tissue. As an alternative to surgical resection, a number of minimally invasive methods have been developed to destroy specific areas of undesirable tissues. Most of these techniques are thermal using cold, e.g. cryosurgery [1][3] or heat, e.g. radiofrequency [4], [5].

Electroporation, also known as electropermeabilization, is a term used to describe the permeabilization of the cell membrane as a consequence of the application of certain short and intense electric fields across the cell membrane, the cells or the tissues. The permeabilization can be temporary (reversible electroporation) or permanent (irreversible electroporation) as a function of the electrical field magnitude and duration, and the number of pulses [6]. Reversible electroporation is commonly used in vitro to facilitate the penetration of various otherwise non-permeable macromolecules across the cell membrane [7][9]. Irreversible electroporation, the ability of certain electrical pulses to permanently permeabilize the cell membrane, has been known for over three decades. For most of this period irreversible electroporation (IRE) was used primarily for ablation of microorganisms and cells in vitro and studied only as an upper limit of electrical parameters for reversible tissue electroporation applications. Our group has pursued the understanding of the electrical fields and processes that produce IRE with single cell micro-electroporation technology [10], [11].

The study of Davalos, Mir and Rubinsky, which showed that IRE can ablate substantial volumes of tissue without inducing a thermal effect and therefore serve as an independent and new tissue ablation modality, opened the way to the use of IRE in surgery [12]. Subsequently, Edd et al. demonstrated tissue ablation with IRE in vivo in the normal liver of rats. [13]. Complete ablation of the targeted liver tissue was achieved by exposing the tissue to electrical parameters that do not induce thermal damage [13]. Massive blood vessel congestion was observed in the sinusoids of the treated volume, which should significantly enhance the treatment. The study concluded that IRE produces precisely delineated ablation zones with cell scale resolution between ablated and non-ablated areas and the ability of mathematical modelling to precisely predict the ablated area. A more recent study was performed to evaluate the long term effects of IRE in a large animal model [14]. The results demonstrated the ability of electroporation to ablate large volumes of tissue using electrical parameters that while substantially above those conventionally used in reversible electroporation do not induce substantial thermal effects. The histology has reconfirmed the results in Davalos et al. [12] and Edd et al. [13] showing that mathematical modeling of electrical and thermal fields are a powerful tool in designing IRE ablation protocols, that IRE can be used to ablate tissue with cell scale resolution and that indeed IRE affects only the cell membrane and therefore spares connective tissue. Another important finding is that IRE can ablate tissue to the margin of large blood vessels and, unlike thermal ablation, is not affected by blood flow in these vessels. This implies that IRE could become an important modality for treatment of tumors near blood vessels.

IRE is a member of a family of non-thermal methods to ablate tissue with electrical pulses, which includes electrochemotherapy (ECT) [7], [15][18] and supra-poration [19][21]. ECT is a relatively new minimally invasive tissue ablation technique that employs reversible electroporation pulses (typically a sequence of eight 100 µs pulses of approximately 1000 V/cm) to reversibly permeabilize the cell membrane and thereby facilitate the penetration into cells of small amounts of non-permeant or low-permeant anti-cancer drugs, such as bleomycin or cisplatin. A major advantage of ECT is that the technique selectively kills, through the use of bleomycin, only the dividing tumour cells and spares the normal non-dividing cells. However, by its nature, ECT requires of the use of chemical agents, which IRE does not. In tissue ablation, ECT is proven to be a safe and highly efficient method to introduce non-permeable cancer drugs into malignant cells and is currently used to treat cutaneous and subcutaneous tumors in humans [15], [16], [22][24]. The standard operating procedures [17] have been established after a clinical multicenter study [18] and several previous single center clinical trials, as reviewed by G. Sersa [25].

Supra-poration, another non-thermal method to kill tissue, is achieved by means of nanosecond electrical pulses in the tens of nanoseconds range and 40–80 kV/cm of field strength [19], [20]. R. Nuccitelli et al. [26] described antitumor effects in mice when nanosecond pulses were delivered in two sets consisting of three consecutive days of treatment separated by two to three weeks. In supra-poration, which also does not employ chemical agents, the pulses delivered are much shorter and the magnitude of the field is higher by an order of magnitude than in IRE. In supra-poration cell death is not a consequence of the irreversible cell membrane permeabilization as in IRE, but the probable result of Ca2+ ions released inside the cells from internal Ca2+ storage vesicles permeabilized by the nanosecond pulses [20]. Each of these new techniques based on the non thermal delivery of electric pulses, namely ECT, IRE and supra-poration, has inherent advantages and disadvantages for tissue ablation and it is quite likely that each will find appropriate uses in modern medicine, separately or in combination.

The present study is the first report of an attempt to evaluate the effectiveness of IRE in treatment of tumors in vivo using preclinical mouse models. Our goal was to determine whether IRE alone, with minimal thermal effects, could actually produce substantial tumor ablation. We also analyzed the cell death and the changes in the vasculature in the treated tissue. The results presented in this study provide further evidence that IRE may become an important minimally invasive modality for treatment of cancer.

Methods Top

Tumour cells culture and tumour production

Cells from a LPB cell line, a methylcholanthrene-induced C57Bl/6 mouse sarcoma cell line [27], were cultured using standard procedures in MEM (Gibco BRL, Cer-gy-Pontoise, France) supplemented with 100 U.ml−1 penicillin, 100 mg.ml−1 streptomycin (Sarbach, France) and 8% foetal calf serum (Gibco). C57Bl/6 female mice, 6–8 weeks old, were inoculated subcutaneously in the left flank with 1×106 cells, producing in 9 days tumors of 4 to 5mm in diameter. Animals were housed at IGR, handled according to the recommended guidelines [28] and protocols approved at IGR.

Tumour treatment

At the start of the procedure mice were anaesthetised using a mixture of xylazine 12.5 mg.kg−1 (Bayer Pharma, Puteaux, France) and ketamine 125 mg.kg−1 (Parke Davis, Courbevoie, France). An incision was performed on the skin near the tumor and the skin flap containing the tumor was lifted, taking particular care to avoid cutting the main blood vessels nourishing the tumor. Stainless-steel plate electrodes were placed in direct contact with both sides of the cutaneous tumor, with the tumor sandwiched between the parallel plates. Good contact of the electrodes with the tumor tissue was produced using electrocardiography paste (Eko-gel, Camina, Egna, Italy). The distance between the electrodes ranged from 3 to 5 mm and was adjusted to tumour size. The spacing between the electrodes was measured and the information was used to set the voltage delivered by the pulse generator, so as to produce the design field magnitude across the tumor, planned for the experiment. The electrical parameters used in the various experiments are listed in Table 1. The square-wave electric pulses (EP) were generated by an electroporation power supply, a Cliniporator™ (Igea, Carpi, Italy) or, to deliver pulses with a frequency of 0.3 Hz, a GHT 1287 (Jouan, St Herblain, France ). To obtain a pulse application frequency of 0.03 Hz, pulses were manually delivered by the operator every 33 s. After EP delivery, the skin incisions were closed with metallic clips, the mice were returned to their cages and the evolution of the treated tumors was followed with measurements of tumor size every second day. Alternatively, mice were kept for different periods of time (between 1 and 72 h) and then humanely sacrificed by CO2 inhalation before the tumors were removed and processed for histological or immunohistochemical analysis.

thumbnailTable 1. Details of the experimental conditions tested and summary of the antitumor effects achieved.

doi:10.1371/journal.pone.0001135.t001

Numerical model of temperature distribution and thermal dose assessment

Electrical fields produce thermal effects due to Joule type electrical energy dissipation. Thermal damage occurs when tissue is exposed to a temperature higher than the physiological temperature for an extended period of time. In electroporation the tissue temperature fluctuates during the application of the pulses and between the pulses. For situations involving fluctuations of temperature, an analytical procedure was developed in which thermal damage is assessed by treating the tissue as if it was at a constant temperature, typically 43°C, for a duration that would result in an equivalent thermal effect to the fluctuating temperature [29]. The following expression is used to calculate the duration necessary to hold the tissue at 43°C to result in an equivalent thermal dose to a fluctuating temperature:
(1)
where Tt is the average temperature during Δt with R = 0.25 when Tt≤43°C and R = 0.5 when Tt>43°C [30], [31] In order to calculate the equivalent thermal dose for each experimental condi-tion, the temperature within the tissue as a function of time was estimated from the Pennes bioheat transfer equation [29] with the addition of a Joule heating term as described in detail in [32].

A one-dimensional explicit transient finite difference model was created to estimate the heating associated with each electroporation procedure [33]. The model geometry was representative of the experimental conditions of a tumour in contact with electrodes, in which the distance between the parallel electrodes was typically 4 mm. The problem was solved under the assumption that the electroporation electrode side facing away from the tumor is cooled by convection heat transfer with the surrounding air and is treated as an infinite fin as described in [32]. The free convection heat transfer coefficient was taken as 15 W m2K−1, and the ambient air, as 25°C [34]. It is assumed that the tissue and the electrode are initially at 37°C and 25°C, respectively.

The values for the tissue heat capacity (4 kJ kg−1K−1), electrical conductivity (0.2 S m−1), thermal conductivity (0.5 W m−1K−1) and density (1000 kg m−3) are taken from the literature for mouse fibrosarcoma tissue 7 days after inoculation of the tumor cells [35], [36]. The values for the electrode heat capacity (477 J kg−1K−1), electrical conductivity (2,222,222 S m−1), thermal conductivity (14 W m−1K−1) and density (7900 kg m−3) are for 304 stainless steel [37]. The metabolism was assumed to be 33,800 W m−3 [35] and blood flow was neglected because of the results presented by Edd et al. which show that perfusion stops during such a procedure [13]. For the experiments with immune-competent mice that had a delay between pulse trains (see Table 1), a 45 second delay was used. For the scenarios with immune-deficient mice, no delay between pulse trains was assumed.

Histology, immunohistochemistry and DNA break detection

For haematoxylin–eosin-saffron (HES) staining.

Tumors were fixed in Finefix (Milestone, Italy) and embedded in paraffin. Sections of 4 µm were prepared for routine HES staining.

For immunohistochemistry of microvessels CD31.

Paraffin sections (4 µm-thick) were dewaxed and rehydrated. Endogenous peroxidase activity was quenched by 3% H2O2 for 10 min. Sections were placed in cover-plates (Shandon, Life Sciences Technology, Cergy-Pontoise, France) and incubated with blocking serum Power Block 1:10 (BioGenex, San Ramon, CA, USA) for 10 min. The slides were then incubated for 1 h with purified rat-anti-mouse monoclonal anti-platelet endothelial cell adhesion molecule (PECAM-1 also called CD31), dilution 1:300, (PharMingen, Heidelberg, Germany) followed by rabbit anti-rat immunoglobulins (Dako Denmark) dilution 1:200, for 30 min, followed by PowerVision poly-HRP anti-Rabbit IgG (ImmunoVision Technologies, Brisbane, CA) for 20 min. Finally, slides were exposed to diaminobenzidine chromogenic substrate (DAB PowerVision Histostaining Kit; ImmunoVision Technologies) for 10 min, washed with distilled water, counterstained with Mayer’s hematoxylin, and mounted in permanent medium (Pertex). All slides were immunolabelled the same day to ensure standardized intensities of immunochemical signals and counterstaining.

TUNEL (Terminal deoxynucleotidyl transferase (TdT)-mediated dUTP Nick End-Labeling)

Double Strand DNA breaks, which are often associated with cell apoptosis, were detected using the “In Situ Cell Death” Detection kit (Roche; Mannheim, Germany) (TUNEL method) performed according to the manufacturer’s instructions. De-paraffinized sections were incubated with Citrate buffer, pH 6 and placed in a water-bath, 98°C for 40 min and all sections were treated with TUNEL reagents (TUNEL mixture: 1 hour at 37°C under a coverslip) except for one where the enzyme was omitted (negative control). After washings with Rince Buffer Biogenex, slides were incubated with the secondary anti-fluorescein-AP conjugate, and the signal was revealed with Fast Red substrate solution for 20 min. Slides were lightly counterstained with hematoxylin prior to aqueous mounting by Aqua-Perm (Shandon Aqua-Perm™ Thermo Electron IVDD Compliant, Waltham, MA).

Results Top

Antitumour effects of irreversible electroporation

Effects on tumors transplanted in immunocompetent mice.

Several electrical pulse parameters (Table 1) were tested to determine those providing a good tumour regression index. The initial parameters tested were derived from electrical parameters of reversible electroporation and employed an electric field (voltage to electrodes distance ratio) of 2000 V/cm and 800 µs total duration of the electric pulses (EP). These conditions did not result in any complete regression but, for some of the treated tumors, a slow-down of growth was recorded (Table 1 and Fig. 1B and 1C).

thumbnailFigure 1. Determination of appropriate electrical parameters for an effective tumor treatment by IRE – experiments with immunocompetent mice.

Panels are identified by letters corresponding to the parameters described in table 1, except for panel A (tumor growth in untreated mice followed as controls of treatment conditions B and C), panel D (control of E and F), panel G (control of H and I) and panel J (control of K and L).

doi:10.1371/journal.pone.0001135.g001

We tested additional modes of application of electrical pulses. In the following set of tests the EP were delivered in two perpendicular directions for a more complete coverage of the tumor. In some tests the number of EP was increased. These changes were only partly efficient as a slow down in tumour growth was again noticed but no complete regression (CR) was achieved, except for one single tumor (Table 1 and Fig. 1E and 1F). The major increase in efficacy on tumor regression resulted from the increase of the ratio of the applied voltage to electrodes distance. The voltage-to-distance ratio is used to characterize the “average” electric field strength in the tissue. The actual electric field in each point of the tissue depends on the geometry of the electrodes (not only the distance between the electrodes but also their shape and dimensions) and on the applied voltage. At 2500 V/cm, between 33% and 67% of complete regressions were achieved in the four experimental conditions tested (Table 1, Fig. 1H, 1I, 1K, and 1L). The efficacy does not seem to be related to the frequency, either 1Hz or 5000 Hz (Figs. 1I and 1K), at which the electric pulses were delivered.

Effects on tumors transplanted in immunodeficient mice.

Similar results to those discussed above were achieved with tumors transplanted in nude mice to assess whether the host immune system has a low or moderate participation in the antitumor effects observed. A supplementary study confirmed that the immune response is not instrumental for IRE ablation which broadens the potential application space for IRE treatment to immunodepressed patients [38]. Another interesting point is that two different regimes were applied (80 EP of 100 µs versus 8 EP of 1000 µs; Table 1 and Fig. 2B and 2E, and 2C and 2F) which were characterized by the same EP total duration and by the same total treatment duration (26.7 s, 2B and 2C, or 267 s, 2E and 2F). The most effective electroporation parameters proved to be condition 2F that led to 12 CR out of 13 mice treated.

thumbnailFigure 2. Determination of appropriate electrical parameters for an effective tumor treatment by IRE – experiments with immunodeficient (nude) mice.

Panels are identified by letters corresponding to the parameters described in table 1, except for panel A (tumor growth in untreated mice followed as controls of treatment conditions B and C) and panel D (control of treatments E and F).

doi:10.1371/journal.pone.0001135.g002

Analysis of thermal effects in the treated tumors.

Table 1 contains two columns of results from the analysis of the thermal effects. One of the columns gives the maximum temperature associated with each procedure, and the second contains the duration required to maintain the tissue at 43°C in order to achieve an equivalent thermal dose as the procedure. The results are for the centerline, which is where the maximal temperature develops and can be considered as an upper limit. The analysis shows that the time delay between pulse trains allowed for significant heat dissipation through the electrodes. The results show that the heating was insignificant for cases 1B, 1C, 1E, 1F, 1I, 1K, 2C, 2E, 2F and most likely did not play a role during cases 1H, 1L and 2B. Since the model does not account for convective heat dissipation out of the sides and top of the tumor or of conduction though the animal itself, it is conservative and provides upper limits of the temperature increase in the tissue.

Histological and immunohistochemical analysis

Evolution of the tumor cell morphology and tumor structure.

In the untreated control tumors, cells display a large nucleus surrounded by a well marked cytoplasm and a well defined cell membrane (Fig 3A). The slides stained with the classical HES staining revealed that 5 min after the EP, no change in tumour cell morphology was detectable. However, the congestion of blood in the tumour vessels was evident. This effect is similar to the irreversible electroporation consequences in the liver described by Edd et al. [13]. Few changes in tissue architecture and cell morphology were detected 1 h and 2 h later (Fig 3B and 3C). However, 6 h after the EP, dramatic changes were observed (Fig 3D). Cytoplasmic limits between the cells were barely distinguishable in several parts of the tumour. The nuclei were still visible, but in a sort of syncytium in these parts of the tumour. At 24 h, almost the entire tumor presented this image (Fig 3F). The nuclei, extremely pycnotic, appeared to be all in the same “cytoplasm” as no limit was detectable between the original cells. At 48 h (Fig 3F) and 72 h, the nuclei were even smaller and tissue necrosis still more evident.

thumbnailFigure 3. Analysis of tumor evolution by HES histological staining after IRE.

A: control; B, C, D, E and F: respectively 1, 2, 6, 24 and 48 h after IRE.

doi:10.1371/journal.pone.0001135.g003

Analysis of cell death in the treated tumors.

TUNEL staining revealed that in the control slides, very few cells were stained positive. The staining was strictly localized at the level of the cell nucleus (Fig 4A). Background was very clear (Fig 4A). However, as early as 5 min after the treatment, changes began, albeit slightly noticeable: a larger number of nuclei were positive, and in many of these nuclei, staining was not contained inside the nucleus, but was clearly spreading in the cytoplasm (Fig 4B). Background was still very clear (Fig 4B). The overall staining of the slices increased 1 h after EP delivery. There was a clear increase in the number of red-dye leaking cells and there was also the presence of a red background (Fig 4C). Thus, two types of staining can be clearly identified in these slices and at later times. At 6 h, a large number of cells still display the diffuse staining (not shown). However, at longer times (24 h, Fig 4D), there was only a heavily red-stained background, which completely diffused and continued to spread throughout the entire tumor section. At 24 h, no more red-dye leaking cells were found, in agreement with the HES images that showed the complete disintegration of the cells.

thumbnailFigure 4. TUNEL analysis at different times after the pulses delivery.

A: control; B, C and D: respectively 5 min and 1 and 24 h after IRE.

doi:10.1371/journal.pone.0001135.g004

Evolution of the tumor vasculature.

Since congestion was detectable immediately after treatment in the HES slides, evolution of the tumour vasculature was also analysed using antibodies against the CD31 endothelial cells specific marker (Fig. 5). The changes in vasculature, which were observed 5 min after EP delivery were slight with respect to the controls (Fig. 5A), displayed highly branched and tortuous vasculature typical to this well vascularized fibrosarcoma tumor. However, 2 h later, well established vascular congestion with dilated vessels was detectable, long vessels were not visible and a slight diffuse staining began to spread in some parts of the slices. At 6 h, blood vessel walls, when still present, were either hyper- or hypo-pigmented. The diffusion of the CD31 marker in the extra-cellular interstitial spaces was much more intense and micro-vascular occlusions were also detectable. These signs indicate advanced tumour vasculature lesions. At 24 h, very advanced vascular lesions were seen, with faint CD31 marker staining of the endothelial cells indicating damaged blood vessel walls, and intense diffusion of the staining in the whole tissue. What remained were blood vessel skeletons over a necrotic background.

thumbnailFigure 5. Immunohistochemical analysis of the tumour vasculature evolution by means of CD 31 staining in LPB tumors after treatment.

A: control; B, C and D: respectively 2, 6 and 24 h after IRE.

doi:10.1371/journal.pone.0001135.g005

Discussion Top

Our results provide evidence that IRE can be used to ablate tumors in vivo. The results achieved with the different sets of electrical parameters indicate that the main parameter affecting the results is the electric field strength. Trains of a large number of short pulses resulted in the best antitumor effects (up to 92% of tumor ablation).

From previous studies on the electric field distribution during electroporation in vivo [13], [39][41], the electrodes type (plates, different types of needle arrays) and the adequacy with which the IRE electrical field affects the totality of the tumor volume will be crucial parameters in achieving successful tumor ablation. In our study the electric pulses were delivered through parallel plates in direct contact with the tumor tissue, which is a configuration that ensures a rather homogeneous electrical field throughout the tumor [39][41]. Therefore, the values used in this study for the voltage to distance ratio are actually a good measure for the field that has developed in the treated tumors.

This study has produced several observations concerning the mechanisms of cell ablation in the IRE method as well as evidence of the efficacy of IRE to completely destroy aggressive tumors. As shown in normal liver tissue [13], classical HES staining revealed that the IRE pulses induce vascular congestion, which should also cause tissue hypoxia and may further contribute to tumor cell death. Twenty-four hours after the application of the pulses, all treated tissue was necrotic. Interestingly, the treatment does not produce massive apoptosis. The number of nuclei stained by the TUNEL reaction slightly increases 1 h after the treatment. The original and remarkable observation reported is the detection of diffused TUNEL staining first in the cytoplasm around the cell nucleus, and later, around the cells. In fact, this diffused staining, that constitutes almost the only TUNEL staining a few hours after the EP, is consistent with the expected effects of the applied IRE pulses. Indeed if irreversible electroporation is actually achieved, the plasma membrane should not reseal and there is also a possibility of affecting the nuclear envelope. Thus, DNA becomes accessible to extracellular nucleases (the DNA double strand breaks (DSB) are the seeding point for the TUNEL staining). Then DNA can spread out of the nucleus and out of the cell (facilitating the generation of DSB, the DSB themselves facilitating DNA spreading out of the nucleus). The observation of a diffused TUNEL staining corroborates that cells are no longer limited by their natural membrane barrier, which is indeed the hallmark of irreversible electroporation.

The analysis of the CD31 staining confirmed that membrane alteration due to IRE is the basis for the efficacy of the treatment. Initially, this staining was meant for the study of the evolution of the tumor vasculature disruption. Our results show rapid and severe lesions of the vasculature. They also reveal that the disaggregation of the membranes starts a couple of hours after the pulse delivery and becomes very intense 6 h later, and is complete at 24 h. The localization of the membrane antigen CD 31 at the cell membrane is progressive and becomes complete. We find that the antigen diffuses throughout the treated tumor volume, indicating the complete disruption of the membranes as well as cell necrosis.

The goal of our study was to evaluate the ability of the IRE electrical pulses to produce cell ablation through their effect on the cell membrane. However, electrical pulses can also have a thermal effect through electrical Joule heating. Since heating is also a well known mechanism for cell ablation, we designed our studies in such a way that the IRE electrical pulses we used produced the desired effect on the cell membrane without having a substantial thermal effect. Therefore, in treatment planning we designed for a mild and benign increase in temperature and used a very conservative mathematical model in the analysis, which does not account for convective heat dissipation through the air or conduction through the animal. With regards to the effects of temperature, it is interesting to compare the experiments 2B and 2C (Table 1 and Fig 2). According to our calculations, eight 1000 µs pulses produce more heating than eighty 100 µs pulses. A larger percentage of CR was achieved with the eighty 100 µs pulses than with the eight 1000 µs pulses. Thus, we can infer that heating cannot be the cause of tumor regression in IRE. Moreover, the higher tumor regression efficacy achieved by applying the same energy in a sequence of shorter pulses to induce less of a thermal dose is consistent with results achieved in vitro by Miller et al. [42]. In their study, 10 pulses of 0.3 ms (at 1500 V/cm) were more efficient in cell killing in vitro than 1 pulse of 3 ms (identical total duration) at the same voltage-to-distance ratio. Miller et al. also showed that complete cell killing was achieved using 3 sets of 10 pulses of 0.3 ms [42]. This results in a total pulse duration of 9 ms which is very close to the 8 ms used in vivo in our experiments.

The electrical pulses used in the experiments described in (2E and 2F) were chosen to yield an almost negligible thermal effect. The two experiments with 80 and 8 pulses yield an equivalent thermal dose of less than 1 s at 43°C. Even though it takes several tens of minutes of exposure at 43°C to retard the growth of murine fibrosarcoma [43], [44], condition 2F resulted in an impressive 92% of CR. Therefore our results, taken together, verify that the antitumor effects of IRE are not thermal. For an identical total energy delivery at the same electrical field, a large number of short pulses, which produce an overall lower tissue temperature, seem more effective than the use of a lower number of longer pulses. Furthermore, the use of long delays between pulses, which allows the tissue to cool, yields an overall lower tissue temperature thereby producing a very effective treatment.

In conclusion, our study has produced the first evidence that IRE can be used to effectively ablate tumors in vivo. Even though our results are encouraging, future studies can be conducted to optimize the results and explore the entire parameter space for this treatment. Nevertheless, the results achieved with different sets of electrical parameters suggest that the main parameters affecting the success of the treatment are not only the electric field strength, but also the number of pulses and the total pulse duration. The histological and immunohistochemical findings reported, as well as the theoretical arguments linked to the use of a high number of short pulses at a very low repetition frequency, demonstrate that tumor ablation is actually due to irreversible permeabilization of the tumor cells and not to excessive heating. It is worthy to note that TUNEL analysis of the cell death and CD31 analysis of the tumor vasculature evolution revealed not only the necrotic pathway of the IRE-caused cell death, but also confirms the mechanisms of the method, that is the actual irreversible electroporation of the cell membrane and its consecutive disintegration. This manuscript thus reports all the parameters necessary to bring IRE towards clinical trials that should help in defining the indications of this new cancer treatment.

Acknowledgments Top

We acknowledge the staff of the Service Commun d’Expérimentation Animale (headed by Dr. P. Gonin).

Author Contributions Top

Conceived and designed the experiments: BR LM RD. Performed the experiments: PO BA FA CB EC. Analyzed the data: PO BR LM BA FA RD. Contributed reagents/materials/analysis tools: PO. Wrote the paper: BR LM RD.

References Top

  1. Onik G, Rubinsky B, Zemel R, Weaver L, Diamond D, et al. (1991) Ultrasound-guided hepatic cryosurgery in the treatment of metastatic colon carcinoma. Preliminary results. Cancer 67: 901–907. Find this article online
  2. Onik GM, Cohen JK, Reyes GD, Rubinsky B, Chang Z, et al. (1993) Transrectal ultrasound-guided percutaneous radical cryosurgical ablation of the prostate. Cancer 72: 1291–1299. Find this article online
  3. Mouraviev V, Polascik TJ (2006) Update on cryotherapy for prostate cancer in 2006. Curr Opin Urol 16: 152–156. Find this article online
  4. de Baere T, Rehim MA, Teriitheau C, Deschamps F, Lapeyre M, et al. (2006) Usefulness of guiding needles for radiofrequency ablative treatment of liver tumors. Cardiovasc Intervent Radiol 29: 650–654. Find this article online
  5. Martin RC (2006) Hepatic tumor ablation: cryo versus radiofrequency, which is better? Am Surg 72: 391–392. Find this article online
  6. Orlowski S, Mir LM (1993) Cell electropermeabilization: a new tool for biochemical and pharmacological studies. Biochim Biophys Acta 1154: 51–63. Find this article online
  7. Mir LM (2001) Therapeutic perspectives of in vivo cell electropermeabilization. Bioelectrochemistry 53: 1–10. Find this article online
  8. Andre F, Mir LM (2004) DNA electrotransfer: its principles and an updated review of its therapeutic applications. Gene Ther 11: Suppl 1S33–42. Find this article online
  9. Mir LM, Moller PH, Andre F, Gehl J (2005) Electric pulse-mediated gene delivery to various animal tissues. Adv Genet 54: 83–114. Find this article online
  10. Huang Y, Rubinsky B (1999) Micro-Electroporation: Improving the efficiency and understanding of electrical permeabilization of cells. Biomedical Microdevices 2: 145–150. Find this article online
  11. Davalos R, Huang Y, Rubinsky B (2000) Electroporation: Bio-electrochemical mass transfer at the nano scale. Microscale Thermophysical Engineering 4: 147–159. Find this article online
  12. Davalos RV, Mir LM, Rubinsky B (2005) Tissue ablation with irreversible electroporation. Ann Biomed Eng 33: 223–231. Find this article online
  13. Edd JF, Horowitz L, Davalos RV, Mir LM, Rubinsky B (2006) In vivo results of a new focal tissue ablation technique: irreversible electroporation. IEEE Trans Biomed Eng 53: 1409–1415. Find this article online
  14. Rubinsky B, Onik G, Mikus P (2007) Irreversible electroporation: a new ablation modality–clinical implications. Technol Cancer Res Treat 6: 37–48. Find this article online
  15. Mir LM, Orlowski S, Belehradek J Jr, Paoletti C (1991) Electrochemotherapy potentiation of antitumour effect of bleomycin by local electric pulses. Eur J Cancer 27: 68–72. Find this article online
  16. Mir LM, Glass LF, Sersa G, Teissie J, Domenge C, et al. (1998) Effective treatment of cutaneous and subcutaneous malignant tumours by electrochemotherapy. Br J Cancer 77: 2336–2342. Find this article online
  17. Mir LM, Gehl J, Sersa G, Collins CG, Garbay JR, et al. (2006) Standard Operating Procedures of the Electrochemotherapy. Eur J Cancer Supplements 4: 14–25. Find this article online
  18. Marty M, Sersa G, Garbay JR, Gehl J, Collins CG, et al. (2006) Electrochemotherapy – an easy, highly effective and safe treatment of cutaneous and subcutaneous metastases: results of the ESOPE (European Standard Operating Procedures of Electrochemotherapy) study. Eur J Cancer Supplements 4: 3–13. Find this article online
  19. Deng J, Schoenbach KH, Buescher ES, Hair PS, Fox PM, et al. (2003) The effects of intense submicrosecond electrical pulses on cells. Biophys J 84: 2709–2714. Find this article online
  20. Beebe SJ, White J, Blackmore PF, Deng Y, Somers K, et al. (2003) Diverse effects of nanosecond pulsed electric fields on cells and tissues. DNA Cell Biol 22: 785–796. Find this article online
  21. Gowrishankar TR, Weaver JC (2006) Electrical behavior and pore accumulation in a multicellular model for conventional and supra-electroporation. Biochem Biophys Res Commun 349: 643–653. Find this article online
  22. Belehradek M, Domenge C, Luboinski B, Orlowski S, Belehradek J Jr, et al. (1993) Electrochemotherapy, a new antitumor treatment. First clinical phase I–II trial. Cancer 72: 3694–3700. Find this article online
  23. Gothelf A, Mir LM, Gehl J (2003) Electrochemotherapy: results of cancer treatment using enhanced delivery of bleomycin by electroporation. Cancer Treat Rev 29: 371–387. Find this article online
  24. Sersa G, Cemazar M, Rudolf Z (2003) Electrochemotherapy: advantages and drawbacks in treatment of cancer patients. Cancer Therapy 1: 133–142. Find this article online
  25. Sersa G (2006) The State-of-the-art of electrochemotherapy before the ESOPE study; advantages and clinical uses. Eur J Cancer Supplements 4: 52–59. Find this article online
  26. Nuccitelli R, Pliquett U, Chen X, Ford W, James Swanson R, et al. (2006) Nanosecond pulsed electric fields cause melanomas to self-destruct. Biochem Biophys Res Commun 343: 351–360. Find this article online
  27. Belehradek J Jr, Barski G, Thonier M (1972) Evolution of cell-mediated antitumor immunity in mice bearing a syngeneic chemically induced tumor. Influence of tumor growth, surgical removal and treatment with irradiated tumor cells. Int J Cancer 9: 461–469. Find this article online
  28. (1998) United Kingdom Co-ordinating Committee on Cancer Research (UKCCCR) Guidelines for the Welfare of Animals in Experimental Neoplasia (Second Edition). Br J Cancer 77: 1–10. Find this article online
  29. Becker SM, Kuznetsoz AV (2006) Numerical Modeling of In Vivo Plate Electroporation Thermal Dose Assessment. ASME J of Biomechanical Engineering 128: 76–84. Find this article online
  30. Damianou CA, Hynynen K, Fan X (1995) Evaluation of accuracy of a theoretical model for predicting the necrosed tissue volume during focused ultrasound surgery. IEEE Transactions on Ultrasonics, Ferroelectrics and Frequency Control 42: 182–187. Find this article online
  31. Sapareto S, Dewey W (1984) Thermal dose determination in cancer therapy. Int J radiation oncology Biol Phys 10: 787–800. Find this article online
  32. Davalos RV, Rubinsky B, Mir LM (2003) Theoretical analysis of the thermal effects during in vivo tissue electroporation. Bioelectrochemistry 61: 99–107. Find this article online
  33. Incropera FP, DeWitt DP (2002) Chapter 5: Transient Conduction. In: Incropera FP, DeWitt DP, editors. Introduction to Heat Transfer. New York: John Wiley and Sons.
  34. White FM (1988) Chapter 1: Introduction. Heat and Mass Transfer. Addision-Wesley Publishing Company, Inc. pp. 1–46.
  35. Deng ZS, Liu J (2001) Blood perfusion-based model for characterizing the temperature fluctuations in living tissue. Phys A STAT Mech Appl 300: 521–530. Find this article online
  36. Swarup A, Stuchly SS, Surowiec A (1991) Dielectric properties of mouse MCA1 fibrosarcoma at different stages of development. Bioelectromagnetics 12: 1–8. Find this article online
  37. White FM (1988) Appendix C: Properties of Metallic Solids Heat and Mass Transfer. Addision-Wesley Publishing Company, Inc. pp. 672–673.
  38. Al-Sakere B, Bernat C, André F, Connault E, Opolon P, Davalos RV, Mir LM (2007) A study of the immunological response to tumor ablation with irreversible electroporation. Technol Cancer Res Treat 6: 301–306. Find this article online
  39. Miklavcic D, Corovic S, Pucihar G, Pavselj N (2006) Importance of tumour coverage by sufficiently high local electric field for effective electrochemotherapy. Eur J Cancer Supplements 4: 45–51. Find this article online
  40. Miklavcic D, Beravs K, Semrov D, Cemazar M, Demsar F, et al. (1998) The importance of electric field distribution for effective in vivo electroporation of tissues. Biophys J 74: 2152–2158. Find this article online
  41. Miklavcic D, Semrov D, Mekid H, Mir LM (2000) A validated model of in vivo electric field distribution in tissues for electrochemotherapy and for DNA electrotransfer for gene therapy. Biochim Biophys Acta 1523: 73–83. Find this article online
  42. Miller L, Leor J, Rubinsky B (2005) Cancer cells ablation with irreversible electroporation. Technol Cancer Res Treat 4: 699–705. Find this article online
  43. Hahn EW, Alfieri AA, Kim JH (1978) Single dose X-irradiation and concomitant hyperthermia on a murine fibrosarcoma. Cancer 42: 2591–2595. Find this article online
  44. Mohamed F, Stuart OA, Glehen O, Urano M, Sugarbaker PH (2004) Docetaxel and hyperthermia: factors that modify thermal enhancement. J Surg Oncol 88: 14–20. Find this article online

Heating the patient: a promising approach?

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Posted 03 Jul 2011 — by James Street
Category Hyperthermia

Abstract

 
Full Article PDF file

There is a clear rationale for using hyperthermia in cancer treatment. Treatment at temperatures between 40 and 44°C is cytotoxic for cells in an environment with a low pO2 and low pH, conditions that are found specifically within tumour tissue, due to insufficient blood perfusion. Under such conditions radiotherapy is less effective, and systemically applied cytotoxic agents will reach such areas in lower concentrations than in well perfused areas. Therefore, the addition of hyperthermia to radiotherapy or chemotherapy will result in at least an additive effect. Furthermore, the effects of both radiotherapy and many drugs are enhanced at an increased temperature. Hyperthermia can be applied by several methods: local hyperthermia by external or internal energy sources, regional hyperthermia by perfusion of organs or limbs, or by irrigation of body cavities, and whole body hyperthermia.

The use of hyperthermia alone has resulted in complete overall response rates of 13%. The clinical value of hyperthermia in addition to other treatment modalities has been shown in randomised trials. Significant improvement in clinical outcome has been demonstrated for tumours of the head and neck, breast, brain, bladder, cervix, rectum, lung, oesophagus, vulva and vagina, and also for melanoma. Additional hyperthermia resulted in remarkably higher (complete) response rates, accompanied by improved local tumour control rates, better palliative effects and/or better overall survival rates. Generally, when combined with radiotherapy, no increase in radiation toxicity could be demonstrated. Whether toxicity from chemotherapy is enhanced depends on sequence of the two modalities, and on which tissues are heated. Toxicity from hyperthermia cannot always be avoided, but is usually of limited clinical relevance.

Recent developments include improvements in heating techniques and thermometry, development of hyperthermia treatment planning models, studies on heat shock proteins and an effect on anti-cancer immune responses, drug targeting to tumours, bone marrow purging, combination with drugs targeting tumour vasculature, and the role of hyperthermia in gene therapy.

The clinical results achieved to date have confirmed the expectations raised by results from experimental studies. These findings justify using hyperthermia as part of standard treatment in tumour sites for which its efficacy has been proven and, furthermore, to initiate new studies with other tumours. Hyperthermia is certainly a promising approach and deserves more attention than it has received until now.

Hyperthermia in Cancer Treatment. National Cancer Institute

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Posted 02 Jul 2011 — by James Street
Category Hyperthermia

Key Points

  1. What is hyperthermia?Hyperthermia (also called thermal therapy or thermotherapy) is a type of cancer treatment in which body tissue is exposed to high temperatures (up to 113°F). Research has shown that high temperatures can damage and kill cancer cells, usually with minimal injury to normal tissues (1). By killing cancer cells and damaging proteins and structures within cells (2), hyperthermia may shrink tumors.

    Hyperthermia is under study in clinical trials (research studies with people) and is not widely available (see Question 5).

  2. How is hyperthermia used to treat cancer?Hyperthermia is almost always used with other forms of cancer therapy, such as radiation therapy and chemotherapy (1, 3). Hyperthermia may make some cancer cells more sensitive to radiation or harm other cancer cells that radiation cannot damage. When hyperthermia and radiation therapy are combined, they are often given within an hour of each other. Hyperthermia can also enhance the effects of certain anticancer drugs.

    Numerous clinical trials have studied hyperthermia in combination with radiation therapy and/or chemotherapy. These studies have focused on the treatment of many types of cancer, including sarcoma, melanoma, and cancers of the head and neck, brain, lung, esophagus, breast, bladder, rectum, liver, appendix, cervix, and peritoneal lining (mesothelioma) (1, 3, 4, 5, 6, 7). Many of these studies, but not all, have shown a significant reduction in tumor size when hyperthermia is combined with other treatments (1, 3, 6, 7). However, not all of these studies have shown increased survival in patients receiving the combined treatments (3, 5, 7).

  3. What are the different methods of hyperthermia?Several methods of hyperthermia are currently under study, including local, regional, and whole-body hyperthermia (1, 3, 4, 5, 6, 7, 8, 9).
    • In local hyperthermia, heat is applied to a small area, such as a tumor, using various techniques that deliver energy to heat the tumor. Different types of energy may be used to apply heat, including microwave, radiofrequency, and ultrasound. Depending on the tumor location, there are several approaches to local hyperthermia:
      • External approaches are used to treat tumors that are in or just below the skin. External applicators are positioned around or near the appropriate region, and energy is focused on the tumor to raise its temperature.
      • Intraluminal or endocavitary methods may be used to treat tumors within or near body cavities, such as the esophagus or rectum. Probes are placed inside the cavity and inserted into the tumor to deliver energy and heat the area directly.
      • Interstitial techniques are used to treat tumors deep within the body, such as brain tumors. This technique allows the tumor to be heated to higher temperatures than external techniques. Under anesthesia, probes or needles are inserted into the tumor. Imaging techniques, such as ultrasound, may be used to make sure the probe is properly positioned within the tumor. The heat source is then inserted into the probe. Radiofrequency ablation (RFA) is a type of interstitial hyperthermia that uses radio waves to heat and kill cancer cells.
    • In regional hyperthermia, various approaches may be used to heat large areas of tissue, such as a body cavity, organ, or limb.
      • Deep tissue approaches may be used to treat cancers within the body, such as cervical or bladder cancer. External applicators are positioned around the body cavity or organ to be treated, and microwave or radiofrequency energy is focused on the area to raise its temperature.
      • Regional perfusion techniques can be used to treat cancers in the arms and legs, such as melanoma, or cancer in some organs, such as the liver or lung. In this procedure, some of the patient’s blood is removed, heated, and then pumped (perfused) back into the limb or organ. Anticancer drugs are commonly given during this treatment.
      • Continuous hyperthermic peritoneal perfusion (CHPP) is a technique used to treat cancers within the peritoneal cavity (the space within the abdomen that contains the intestines, stomach, and liver), including primary peritoneal mesothelioma and stomach cancer. During surgery, heated anticancer drugs flow from a warming device through the peritoneal cavity. The peritoneal cavity temperature reaches 106–108°F.
    • Whole-body hyperthermia is used to treat metastatic cancer that has spread throughout the body. This can be accomplished by several techniques that raise the body temperature to 107–108°F, including the use of thermal chambers (similar to large incubators) or hot water blankets.

    The effectiveness of hyperthermia treatment is related to the temperature achieved during the treatment, as well as the length of treatment and cell and tissue characteristics (1, 2). To ensure that the desired temperature is reached, but not exceeded, the temperature of the tumor and surrounding tissue is monitored throughout hyperthermia treatment (3, 5, 7). Using local anesthesia, the doctor inserts small needles or tubes with tiny thermometers into the treatment area to monitor the temperature. Imaging techniques, such as CT (computed tomography), may be used to make sure the probes are properly positioned (5).

  4. Does hyperthermia have any complications or side effects?Most normal tissues are not damaged during hyperthermia if the temperature remains under 111°F. However, due to regional differences in tissue characteristics, higher temperatures may occur in various spots. This can result in burns, blisters, discomfort, or pain (1, 5, 7). Perfusion techniques can cause tissue swelling, blood clots, bleeding, and other damage to the normal tissues in the perfused area; however, most of these side effects are temporary. Whole-body hyperthermia can cause more serious side effects, including cardiac and vascular disorders, but these effects are uncommon (1, 3, 7). Diarrhea, nausea, and vomiting are commonly observed after whole-body hyperthermia (7).
  5. What does the future hold for hyperthermia?A number of challenges must be overcome before hyperthermia can be considered a standard treatment for cancer (1, 3, 6, 7). Many clinical trials are being conducted to evaluate the effectiveness of hyperthermia. Some trials continue to research hyperthermia in combination with other therapies for the treatment of different cancers. Other studies focus on improving hyperthermia techniques.

    To learn more about clinical trials, call the National Cancer Institute’s (NCI) Cancer Information Service at the telephone number listed below or visit NCI’s Clinical Trials Home Page.

Selected References

  1. van der Zee J. Heating the patient: A promising approach? Annals of Oncology 2002; 13:1173–1184.
  2. Hildebrandt B, Wust P, Ahlers O, et al. The cellular and molecular basis of hyperthermia. Critical Reviews in Oncology/Hematology 2002; 43:33–56.
  3. Wust P, Hildebrandt B, Sreenivasa G, et al. Hyperthermia in combined treatment of cancer. The Lancet Oncology 2002; 3:487–497.
  4. Alexander HR. Isolation perfusion. In: DeVita VT Jr., Hellman S, Rosenberg SA, editors. Cancer: Principles and Practice of Oncology. Vol. 1 and 2. 6th ed. Philadelphia: Lippincott Williams and Wilkins, 2001.
  5. Falk MH, Issels RD. Hyperthermia in oncology. International Journal of Hyperthermia 2001; 17(1):1–18.
  6. Dewhirst MW, Gibbs FA Jr, Roemer RB, Samulski TV. Hyperthermia. In: Gunderson LL, Tepper JE, editors. Clinical Radiation Oncology. 1st ed. New York, NY: Churchill Livingstone, 2000.
  7. Kapp DS, Hahn GM, Carlson RW. Principles of Hyperthermia. In: Bast RC Jr., Kufe DW, Pollock RE, et al., editors. Cancer Medicine e.5. 5th ed. Hamilton, Ontario: B.C. Decker Inc., 2000.
  8. Feldman AL, Libutti SK, Pingpank JF, et al. Analysis of factors associated with outcome in patients with malignant peritoneal mesothelioma undergoing surgical debulking and intraperitoneal chemotherapy. Journal of Clinical Oncology 2003; 21(24):4560–4567.
  9. Chang E, Alexander HR, Libutti SK, et al. Laparoscopic continuous hyperthermic peritoneal perfusion. Journal of the American College of Surgeons 2001; 193(2):225–229.

The effects inhibiting the proliferation of cancer cells by far-infrared radiation (FIR) are controlled by the basal expression level of heat shock protein (HSP) 70A.

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Posted 02 Jul 2011 — by James Street
Category HSP70 gene, Hyperthermia

Ishibashi J, Yamashita K, Ishikawa T, Hosokawa H, Sumida K, Nagayama M, Kitamura S.

Department of Oral and Maxillofacial Anatomy, Medical Science for Oral and Maxillofacial Regeneration, Graduate School of Health Biosciences, University of Tokushima, 3-18-15 Kuramoto, Tokushima 770-8504, Japan.

Abstract

We developed a tissue culture incubator that can continuously irradiate cells with far-infrared radiation (FIR) of wavelengths between 4 and 20 microm with a peak of 7-12 microm, and found that FIR caused different inhibiting effects to five human cancer cell lines, namely A431 (vulva), HSC3 (tongue), Sa3 (gingiva), A549 (lung), and MCF7 (breast). Then, in order to make clear the control system for the effect of FIR, the gene expression concerned to the inhibition effect by FIR were analyzed. In consequence, basal expression level of HSP70A mRNA was higher in A431 and MCF7 cells than in the FIR-sensitive HSC3, Sa3, and A549 cells. Also, the over expression of HSP70 inhibited FIR-induced growth arrest in HSC3 cells, and an HSP70 siRNA inhibited the proliferation of A431 cells by irradiation with FIR. These results indicate that the effect of a body temperature range of FIR suppressing the proliferation of some cancer cells is controlled by the basal expression level of heat shock protein (HSP) 70A. This finding suggested that FIR should be very effective medical treatment for some cancer cells which have a low level of HSP70. Still more, if the level of HSP70 in any cancer of a patient was measured, the effect of medical treatment by FIR can be foreseen for the cancer.

Magnetic Nanoparticles Fry Tumors

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Posted 01 Jul 2011 — by James Street
Category Hyperthermia, NanoTechnology, Physics and Engineering
by Tim Wogan on 1 July 2011, 3:21 PMAny parent fretting over a child’s fever knows that temperatures just a few degrees above normal can kill. But cancer researchers have now found a way to make high temperatures heal. In a new study, a team found that injecting mice with tiny magnets and cranking up the heat eliminated tumors from the animals’ bodies with no apparent side effects.

The idea of killing cancer with heat isn’t new. Researchers know that, like normal cells, cancer cells start to die when the mercury rises above 43˚C. The trick is figuring out how to kill the cancer without harming the body’s own cells. One promising idea, known as magnetic hyperthermia, involves injecting minuscule “nanoparticles,” basically microscopic lumps of iron oxide or other compounds, into tumors to make them magnetic. The patient is put into a magnetic field that reverses direction thousands of times every second. The magnetic nanoparticles are excited by the applied field and begin to get hot, heating and potentially destroying the surrounding cancer tissue. Because healthy tissue is not altered by the magnetic field, it does not heat up and is not damaged.

But the therapy has yet to make its way to the clinic, with only a single reported trial in humans (with modest success). This is largely because conventional nanoparticles interact only weakly with the applied field, so quite a large dose is needed to generate enough heat to damage the tumor. Although nanoparticles aren’t particularly toxic, in large quantities they can trigger the body’s immune system to attack them, causing allergic reactions.

Nanoscientist Jinwoo Cheon of Yonsei University in Seoul and colleagues set out to create a nanoparticle that would get hotter than traditional nanoparticles so that not as many would need to be injected into the body. They made two-layer nanoparticles, each containing a core of one magnetic mineral inside a shell of another. Because of an esoteric interaction between the two minerals, called exchange coupling, these “core-shell” nanoparticles interacted far more strongly with the magnetic field than do traditional nanoparticles and released up to 10 times as much heat. That means one would need to give only 10% of the original dose to patients to achieve the same degree of hyperthermia as with traditional nanoparticles.

The team tested its technique on three mice whose abdomens had been grafted with cells from human brain cancer. The researchers injected the tumors with core-shell nanoparticles and placed the mice inside a coil of wire (see illustration). They turned on an alternating current in the coil, creating an alternating magnetic field. Although the researchers weren’t able to measure the precise temperatures inside the tumors, their estimates are between 43˚ and 48˚C. After 10 minutes, the team removed the mice from the coil and monitored the tumors for the next 4 weeks.

All traces of cancer disappeared from the mice with no apparent side effects, the team reported online 26 June in Nature Nanotechnology. For comparison, another group of mice were treated instead with a single dose of doxorubicin, a traditional anticancer drug. Although it initially shrunk some of the tumors, they grew back to four times their original size by the end of the trial. Heat treatment after an injection of traditional iron oxide nanoparticles had no significant effect on the tumors.

Nanoengineer Naomi Halas of Rice University in Houston, Texas, is impressed. “This group has solved the key impasse that has arrested the development of magnetic nanotherapies, that is, the weak response of the nanoparticle to the applied magnetic field,” she says. “I am so happy that more of these types of nanoparticle-based hyperthermal therapies are being developed to increase the arsenal of weapons against cancer.”

Feature: Duke Physicians Turn Up Heat on Tumors To Hasten Their Demise

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Posted 22 Jun 2011 — by James Street
Category Hyperthermia

By Duke Medicine News and Communications

DURHAM, N.C. — Writings from the ancient Egyptians claim they used an instrument called a “fire drill” to cauterize cancers, but 3,000 years later doctors have not yet mastered the art of directing heat to the desired spot to kill cancers. Targeting a tumor deep within the body or a limb is like trying to bake a single cookie in an oven that remains cool to the touch, say researchers at Duke University Medical Center’s Hyperthermia Program. Thus, practical barriers have stymied the widespread use of heat to shrink tumors: the tumor is hard to access, the target is hard to hit and physicians cannot easily measure its temperature.

Recently, the Duke program received a $19 million continuation grant from the National Cancer Institute (NCI) to study and apply the benefits of using heat, or hyperthermia, to treat patients with cancer. The grant is currently entering its 19th year of continuous funding and, according to program members, Duke has the only federally funded research program dedicated to making heat treatment a viable option for patients with a wide variety of cancers.

The new funds are helping the program’s researchers refine modern-day tools to implement the ancient idea of “targeted fire” to kill cancers. They are using microwave antennae to beam heat at a precise spot in the body; leg cuffs that encircle the affected area and deliver targeted heat; and a miniature water Jacuzzi that transmits microwave heat selectively to cancerous breasts.

Such modern inventions – many developed and built at Duke — are enabling physicians to more effectively employ heat to target and destroy tumors. As a result, studies at Duke and elsewhere have demonstrated that hyperthermia boosts the killing power of radiation and chemotherapy by up to ten times greater than without heat. When the tumor reaches the desired temperature, physicians blast it with chemotherapy or radiation.

“The question isn’t whether hyperthermia works, but how do we apply the therapy so it achieves the desired goal and how do we inform physicians on its proper use,” said Mark Dewhirst, DVM, Ph.D., professor of radiation oncology at Duke and director of the hyperthermia program. “Our goal is to enable doctors to write a dose prescription for heat treatment that is user-friendly and can be ‘filled’ as we would any other prescription.”

During the past year alone, Duke researchers demonstrated precisely how heat can dramatically increase the ability of radiation to shrink recurrent tumors in the breast, chest wall, head and neck and skin. In addition, a Duke pilot study showed that a combination of hyperthermia, radiation and chemotherapy caused a complete clinical response in 10 of 12 patients with locally advanced cervix cancer, meaning the tumors shrank to insignificant size.

Heating tumors elicits a series of important changes that hasten the tumor’s demise, said Dewhirst. Heat makes blood vessels leakier and thus enables chemotherapy to penetrate the tumor more effectively. Heat also increases oxygen levels within the tumor, and oxygen is critical to the proper functioning of radiation and chemotherapy inside a cell. Finally, heat amplifies the level of DNA damage that chemotherapy and radiation inflict upon the cancer cell by inhibiting enzymes that normally repair such DNA damage.

Still, persistent roadblocks have made hyperthermia difficult to administer and even harder to control and measure, said Dewhirst. Physicians lack clear-cut guidelines that dictate optimum temperatures and heating times for specific cancers. Technological and physical barriers have prevented physicians from determining whether the heat is reaching its intended target and, if so, how to accurately measure the tumor’s temperature. As a result, doctors have not been able to deliver a uniform heat dose for all patients, he said.

Dewhirst’s team has turned to the exquisite imaging capabilities of magnetic resonance imaging (MRI) to better visualize the tumor’s location and gauge its temperature in real time, as it is being heated. A twist on the traditional MRI, the Duke team of biomedical engineers, physicists and radiation oncologists have created the MRI “imaging thermometer” that measures a tumor’s temperature by measuring, in part, how fast water moves around the tissue. Because water moves faster when heated, the MRI detects this shift and depicts the hot spot as a red glow on the computer screen. When optimum temperatures are reached, the tumor is then treated with radiation or chemotherapy.

“We essentially create a three dimensional temperature map using the color red to depict the hottest region and blue to depict the coldest, so we can see exactly what we are heating and how hot the tumor becomes,” said Dewhirst.

The new imaging thermometer is part of a dedicated MRI device used exclusively for hyperthermia treatment, he said. The NCI grant will allow the Duke team to fine-tune the device so it can be developed as an attachment to existing MRIs, allowing a broader range of treatment facilities to offer hyperthermia.

Indeed, Duke engineers have built a variety of devices that enable physicians to better target and heat tumors. Among the latest designs is a hyperthermia “cuff” that surrounds limbs that contain tumors, said Dewhirst. The cuff is a cylindrical tube with multiple antennae inside. The patient places the affected arm or leg inside the tube and doctors adjust each antenna to aim the microwave energy toward the tumor. Precisely aiming the heat at the tumor ensures the heat reaches its intended target.

“We know where we want the hot spot to be,” said Dewhirst. “The issue is how to tune the heating device to deliver the heat to the right spot. Now we have developed a way to tune the microwaves so that we can target the right spot with the right dose.”

For breast cancer patients, Duke engineers have designed a hyperthermia treatment table with a small opening through which the cancerous breast protrudes. The patient lies face down while the breast, resting in a small cup of water, is heated via microwave energy. The heat triggers the chemotherapy that has just been infused to settle inside the tumor. Once there, it trickles out of its protective coating — a tiny fat bubble called a liposome – and attacks the tumor’s genetic machinery. The body’s normal tissues remain unheated, so the drug is not preferentially delivered there

“Encapsulating the chemotherapy inside of liposomes enables us to deliver 30 times more chemotherapy than we normally could to the tumor site, without poisoning the rest of the body,” said Duke oncologist Kimberly Blackwell, M.D., who leads the studies using this new technology. “Heat also boosts the drugs’ potency by interfering with mechanisms that control a cancer cell’s ability to replicate.”

Heat, while powerful, is only part of the equation, said Dewhirst. The tumor itself provides a highly sophisticated network of signals – a so-called microenvironment — that drives how the tumor behaves in response to the treatment. Heating a tumor produces different results than heating normal tissue; thus, Dewhirst’s team is studying how a tumor’s oxygen levels, pH, glucose levels and other characteristics affect its response to heat.

INTO THE CLINIC

Armed with the latest technology, the Duke team has initiated a range of new studies to assess hyperthermia’s ability to treat a wider variety of cancers. Among the latest Duke trials are studies examining:
• Hyperthermia and cervical cancer — Although easily treated in the U.S.in its early stages, locally advanced cervical cancer is the number-one cancer killer of women worldwide. The cure rate in the U.S. is considerably reduced among women whose cancer is not diagnosed in its early stages. Duke is conducting an international phase III study in the U.S. and Europe, adding hyperthermia to the current standard of care for advanced cervical cancer to determine whether the addition of heat reduces mortality.

• Hyperthermia and breast cancer — Women with chest wall recurrences of breast cancer will receive the Duke-developed heat-sensitive fat liposome containing the drug doxorubicin in a phase I trial. Subsequently, the drug will enter a phase I/II trial in women with locally advanced breast cancer. Currently enrolling are phase I/II trials for women with locally advanced breast cancer to receive the commercially-available liposome, Doxil.

• Genetic studies of locally advanced breast cancer — Researchers will create genetic profiles of breast and cervix cancer tumor tissue to try to determine if certain genetic traits can predict recurrence after chemotherapy treatment. The researchers will study markers of low oxygen levels in tumors, which often are associated with poor outcome, and elevation of genetic markers of inflammation, another predictor of poor outcome.

• Sarcoma — Two clinical trials for patients with tumors of connective tissue in the arms or legs (sarcomas) will test the Duke-designed heat delivery cuff together with radiation or chemotherapy. Ellen Jones, M.D., will study hyperthermia’s effects on boosting tumor oxygen levels and in turn how increased oxygen enhances radiation’s cancer-killing effects.

• Melanoma — Douglas Tyler, M.D., will test whether heat treatment increases the effectiveness of chemotherapy in a rare, aggressive form of melanoma of the arm or leg. The limb’s circulation will be cut off from the rest of the body during treatment to shield the vital organs from high dose chemotherapy.

 

Quercetin and tamoxifen sensitize human melanoma cells to hyperthermia

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Posted 09 Jun 2011 — by James Street
Category Hyperthermia, Melanoma, quercetin

Piantelli, M.; Tatone, D.; Castrilli, G.; Savini, F.; Maggiano, N.; Larocca, L. M.; Ranelletti, F. O.; Natali, P. G.*

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Abstract

Hyperthermia produces regression of human cancer. Because hyperthermia has produced only limited results, attention has focused on searching for substances able to sensitize tumour cells to the effects of hyperthermia. The flavonoid quercetin has been reported to be a hyperthermic sensitizer in ovarian and uterine cervical tumours and in leukaemia. Quercetin and tamoxifen inhibit melanoma cell growth. We therefore investigated whether quercetin and tamoxifen can sensitize M10, M14 and MNT1 human melanoma cells to hyperthermia. We observed that both quercetin and tamoxifen synergize with hyperthermia (42.5°C) in reducing the clonogenic activity of M14 and MNT1 and in inducing apoptotic cell death in all three cell lines. As revealed by flow cytometric and Northern blot analyses, quercetin and tamoxifen reduced heat shock protein-70 expression at both protein and mRNA levels. Our results suggest that quercetin and tamoxifen can be usefully combined with hyperthermia in the therapy of recurrent and/or metastatic melanoma.