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Russian Physics Journal, Vol. 50, No. 9, 2007

MAGNETIC INDUCTION HYPERTHERMIA

V. N. Nikiforov UDC 621.318

A review of physical principles and experimental data on magnetic hyperthermia are presented. The main principles of magnetic hyperthermia are considered. Results of its application in the therapy of oncology diseases are presented.

1. PHYSICAL PRINCIPLES OF HYPERTHERMIA

Two components of ac electromagnetic fields E and H can cause heating of tissues. Based on the Maxwell equations and thermodynamic relations, not only the electrocaloric effect, namely, the heat absorption in substance caused by the electric field component E [6], but also the magnetocaloric effect caused by variations in H can be calculated. These effects are mainly determined by the dielectric ( and /) and magnetic () properties of substances, respectively. Because magnetism of biological objects is negligibly small, biologically compatible nontoxic magnetic nanoparticles (based on magnetite and so on) are used to strengthen the influence of an external magnetic field.

Thermodynamic relations for a magnet in a magnetic field are similar to those for a dielectric in an electric field [6]. However, an essential difference is that the magnetic field, unlike the electric field, does no work on charges moving in it, because the Lorentz force is perpendicular to the velocity vector of the moving charge. To calculate a change in the energy of the medium when the magnetic field is switched on, electric fields induced by magnetic field variations should be considered. In their turn, high-frequency electromagnetic fields cause heating due to the electric field component. At low frequencies, this effect is insignificant. At the same time, the role of the magnetic field component in magnet heating is significant only at low frequencies [6]. Therefore, inclusion of electromagnetic induction hyperthermia in a separate class of high-frequency phenomena [1], on the one hand, and of magnetic hyperthermia as a low-frequency influence, on the other hand, though relative, is justified.

Low-frequency (less than 100 kHz) electromagnetic fields can cause heating of magnetic nanoparticles at the expense of the magnetocaloric effect [7], magnetic reversal in the presence of a hysteresis loop, and magnetic crystal anisotropy of superparamagnetic particles. Physics of the magnetocaloric effect of this phenomenon is the following: elementary magnetic moments are directed chaotically without magnetic field, and hence the magnetic contribution to the entropy is significant. As the magnetic field increases, the magnetic moments are ordered along the field. As a result, the magnetic entropy component SM decreases. Because the magnetization process is close to adiabatic one, the total entropy S does not change, but the entropy component caused by thermal motion increases. Thus, the ferromagnet temperature increases with the magnetic field. Quantitatively, the temperature change is calculated from the formula

( )S

H

H

dMT dTT H

C = .

From the above formula it follows that the temperature can raise when the magnetic moment S changes with temperature, since the heat capacity C is always positive. The magnetocaloric effect of ferrites weakens significantly at

M. V. Lomonosov Moscow State University; e-mail: nvn@lt.phys.msu.ru. Translated from Izvestiya Vysshikh Uchebnykh Zavedenii, Fizika, No. 9, pp. 6072, September, 2007. Original article submitted December 28, 2006.

1064-8887/07/5009-0913 2007 Springer Science+Business Media, Inc.

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temperatures above the Curie point TC because the magnetic moment sharply decreases. It is well known that a ferromagnet placed in an ac magnetic field is heated. The thermal effect during magnetic reversal of the magnet is proportional to the area of the hysteresis loop in B, coordinates.

Superparamagnetic gels based on single-domain particles with sizes less than 810 nm have recently been used for hyperthermia. The thermal effect of magnetic reversal of particles is not connected with the hysteresis phenomena, because no domain system is present; it is caused by the energy change accompanying the magnetic moment rotation, that is, by the magnetic crystal anisotropy. Induction heating of tumors in vivo in an ac magnetic field is performed at frequencies, as a rule, below 200 kHz [8] with the help of preliminary implanted ferrite nanoparticles. The choice of the working frequency is determined by physiological properties of the electromagnetic field of this frequency, magnetic susceptibility, and ferrite particle sizes. Heating occurs due to losses in the hysteresis loop [9]. It should be noted that when sizes of ferromagnetic particles decrease, they transform to the superparamagnetic state; in this case, the heating mechanisms are caused by the magnetic crystal anisotropy.

Ferromagnetic materials are heated well at rather low frequencies

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this reason, a number of researchers prefer to use temperatures up to 55C [18]. Harmful effects on normal cells decrease when the nanoparticles are injected directly into a tumor. The power of particle heating is determined quantitatively as a rate of energy absorption or specific energy absorption; it is equal to the amount of energy transformed into heat at high temperatures by unit mass per unit time [19]. In addition to the particle sizes and shapes that influence their magnetic properties and hence the specific energy absorption, that is, their heating power, a dependence between the heating temperature and the magnetic field strength should be considered when comparing experiments with different parameters of tissues. Tumors with volumes of 300 mm3 can be heated to the required temperatures. No potential problems are expected for great volumes of tissues (for example, for volumes greater than 1000 mm3) given that the employed magnetic liquid is properly thermally regulated and nanoparticles are injected directly into the tumor [18]. The frequency of the ac electromagnetic field should be chosen based on a compromise, namely, it must be higher than the frequency capable of provoking neuromuscular response and lower than the frequencies causing overheating of healthy tissues. Some researchers believe that the frequency should be in the range 1001000 kHz [20]; in this case, correctly chosen frequencies and electromagnetic field strengths produce no side effects of the ac magnetic field on a human body.

2. MAGNETIC HYPERTHERMIA

The magnetic susceptibility of magnetic nanomaterials depends not only on the temperature but also on the magnetic field H. A dependence of the magnetic moment M on the magnetic field is characterized by an MH curve, where the magnetic moment M saturates at high H values. In addition, hysteresis caused by irreversibility during magnetization is observed for ferro- and ferrimagnetic materials. The hysteresis is caused by pinning of magnetic domain boundaries at impurities or grain boundaries in the material as well as by the effect of magnetic crystal anisotropy. The MH curves are called hysteresis loops. The shape of these loops is determined by the particle size: large particles (with micron sizes or larger) are split into domains (the polydomain state) thereby leading to a narrow hysteresis loop, because rather low magnetic field energy is required to displace domain walls. At the same time, for smaller particle sizes, another scheme can be realized with a number of individual domains, which yields a wide hysteresis loop. For even smaller sizes of magnetic nanoparticles (a few tens of nanometers or less), the superparamagnetism phenomenon can be observed when the magnetic moment of the particle is free and can respond to thermal energy, by analogy with a gas of noninteracting magnetic needles. At the same time, individual atomic magnetic moments inside a single particle are parallel to each other (the condition of magnetic saturation inside the particle). This leads to the absence of hysteresis, though the MH curve is saturated. This can be explained within the framework of the Curie theory considering that each particle has one domain. Physics of superparamagnetism is based on the activation law. The relaxation time of particle magnetization is [21, 22]

0B

exp Ek T

= . (1)

Here E is the energy barrier for the magnetic moment and kT is the thermal energy. For noninteracting particles, the pre-exponential factor 0 has an order of 10101012 s and depends weakly on the

temperature [22]. The origin of the energy barrier E is caused by several reasons, including the effects of magnetic crystal and shape anisotropies. In the most simple cases, it is given by the formula E = KV, where K is the energy density of the magnetic crystal anisotropy and V is the particle volume.

Direct proportionality between E and V is a reason why superparamagnetism the thermally activated process of variations of the magnetic moment orientation for a single-domain nanoparticle is important for small nanoparticles, because the barrier E for them is comparable with the thermal energy kT, for example, at room temperature. However, it is important to note that superparamagnetism observations depend not only on the temperature but also on the measurement time m. For

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particles demonstrate paramagnetic properties. For >> m, the rate of changing the magnetic moment direction is low, the time of the magnetic moment reversal is much greater than the observation period, and the so-called blocked state is observed for the system, that is, its quasi-static properties are manifested.

The blocking temperature Tb is determined as a temperature of an intermediate state between these two states with = m. In typical experiments, m can change from m = 102 s for magnetization in the quasi-stationary case to 101105 s for the ac susceptibility and to m = 107109 s for the Mssbauer 57Fe nucleus spectroscopy.

Thus, the hysteresis of the MH curve, characteristic of ferro- and ferrimagnets, indicates the need for application of external energy to overcome the energy barrier for domain wall motion caused by the magnetic crystal anisotropy, microstructural impurities, and grain boundaries in the material. The given energy is delivered by an applied magnetic field and can be characterized by the area of the hysteresis loop of the MH curve. When an ac magnetic field is applied to a ferro- or ferrimagnet, a constant energy flux caused by transformation of the magnetic energy into heat arises in the magnetic material. This provides the physical basis for the method of magnetic hyperthermia.

Similar reasoning can be used for superparamagnetic materials without magnetic hysteresis of MH curves. In superparamagnetic materials, the external magnetic energy is necessary to level coherently the magnetic moments of the particle to achieve magnetic saturation. These mechanisms for superparamagnetic nanoparticles are discussed below.

The magnetic characteristics for ferro- or ferrimagnetic nanoparticles injected into a blood vessel have analogous (to the above-described) character. The magnetic response of the blood vessel also involves the paramagnetic contribution, for example, from hemoglobin molecules containing iron, and the diamagnetic contribution, for example, from vessel proteins which comprise carbon, hydrogen, nitrogen, and oxygen atoms. A magnetic signal of injected particles, irrespective of their sizes, considerably exceeds the magnetic contribution of the blood vessel. This magnetic selectivity is one of the advantages of biomedical applications of the magnetic nanoparticles.

3. MECHANISMS OF NANOPARTICLE HEATING IN A MAGNETIC LIQUID

The ferromagnetic particles demonstrate hysteresis properties in an ac magnetic field; they are heated. The amount of heat produced in unit volume per unit time is determined by the frequency of the ac magnetic field multiplied by the hysteresis loop area in the MH coordinates:

1P f HdM= . (2) The given formula ignores other possible heating mechanisms, for example, induction heating by eddy currents and

a ferromagnetic resonance. Because particles used for magnetic hyperthermia have nanosizes (10100 nm) and frequencies of the ac magnetic field (f = 0.051.2 MHz), as a rule, are too low to generate noticeable eddy currents, the influence of induction heating can be neglected. The effect of ferromagnetic resonance also can take place, but only at frequencies much higher than those used for magnetic hyperthermia. For ferromagnetic particles with sizes greater than those of superparamagnetic nanoparticles, there is no explicit frequency dependence of the hysteresis loop area, so that the rate of heat emission can be determined from quasi-statical measurements of the hysteresis loop, for example, by means of a vibration or SQID magnetometer. Reorientation and growth of spontaneously magnetized domains in a ferromagnetic particle depends not only on the microstructural features like vacancies, impurities, or grain boundaries but also on the special features of the magnetic material like magnetic crystal anisotropy and particle shapes and sizes.

In principle, significant hysteresis heating of ferromagnetic particles can be obtained with the use of highly anisotropic magnetic materials like NdFeB or SmCo. However, restrictions on the ac magnetic field strength H which must be safe for a patient do not allow complete saturation of the magnetic moment M to be achieved and hence, loops of total saturation cannot be realized for these materials. Small (unsaturated) loops can be used; they also cause heating, but its level is much lower. In fact, as follows from Eq. (2), maximum thermal emission P is realized for a rectangular hysteresis

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loop. This case can be realized for an ensemble of uniaxial anisotropic magnetic particles with intrinsic magnetic moments controlled by the external magnetic field H. However, it is difficult to realize the given configuration in vivo.

In the last decade, magnetic hyperthermia has been given a new impetus due to superparamagnetic liquids ferroliquids with one-domain magnetic nanoparticles suspended in water or hydrocarbon solution and showing superparamagnetic properties that are used as working agents [3, 4, 23]. When a ferroliquid is removed from the magnetic field, its magnetization vanishes because of the thermal energy of the environment. This relaxation corresponds either to physical rotation of particles directly inside the liquid or to rotation of nuclear magnetic moments inside each particle. Rotation of particles is usually referred to as the Brownian rotation, and rotation of the magnetic moment inside each particle is called the Nel relaxation. Each of these processes is characterized by its own relaxation time: B for the Brownian rotation that depends on the hydrodynamic properties of the liquid and N for the Nel relaxation determined by the ratio of the energy of magnetic anisotropy of superparamagnetic particles to the thermal energy. Both Brownian and Nel processes can occur in a ferroliquid considering that the Nel relaxation time N is appropriate for spatially fixed superparamagnetic particles where none of physical rotations of the particle is possible. The relaxation times B and N depend differently on the particle size; the losses caused by the Brownian rotation are maximal at frequencies lower than the Nel relaxation frequencies for a preset nanoparticle size.

The physical substantiation of heating of superparamagnetic nanoparticles by ac magnetic fields was given by Rosensweig [24]. He used the Debye model constructed for a description of dielectric dispersion in polar liquids [25] and understanding that restrictions on the rate of change of the magnetic moment M in a ferroliquid mean that the magnetic moment M is delayed from the ac external field H. For low field strengths under assumption of minimal interactions of superparamagnetic particles, magnetization of a ferroliquid in an external ac magnetic field can be described in terms of the complex magnetic susceptibility = + , where and are frequency-dependent. Exactly the component specifies the thermal emission power [24]:

22 0P f H= . (3)

Formula (3) can be interpreted as follows: the delay of the magnetic moment M from the ac magnetic field H is caused by the transformation of the magnetic energy into the internal energy of the superparamagnetic particle. This simple model is in good agreement with the available experimental data, namely, with a square-law dependence of P2 on H [26] and a dependence of on the frequency f of an ac magnetic field [2729].

The power of thermal emission from magnetic nanoparticles is conventionally described in terms of the specific energy absorption in units Wg1. The specific energy absorption multiplied by the particle density yields the thermal emission powers in unit volume P1 and P2 which allow the efficiency of heating of magnetic particles with different sizes to be compared [3, 3034]. A comparison demonstrates that the conventional ferromagnetic material requires application of ac magnetic fields with strength of 100 kAm1 or higher, which is less than the saturation field strength. Therefore, only small hysteresis loops can be used with allowance for a 15 kAm1 restriction on the field strength, which in its turn, results in low values of the specific energy absorption. On the contrary, superparamagnetic nanomaterials can effectively generate thermal power even in lower magnetic fields. For example, according to Hergt et al. [34], one of the ferroliquids has the specific energy absorption equal to 45 Wg1 in a field of 6.5 kAm1 with frequency f = 300 kHz. Extrapolation of this result to a field of 14 kAm1 yields a specific energy absorption of 209 Wg1. At the same time, for one of the best ferromagnetic liquids, the specific energy absorption is 75 Wg1 in a field of 14 kAm1. Obviously, the use of magnetic particles for magnetic hyperthermia is preferable in magnetic fields as low as possible.

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4. APPLICATION OF MAGNETIC HYPERTHERMIA IN ONCOLOGY

Prospects for destructing tumor cells by local hyperthermia has led to the development and creation of various devices intended for heating of oncology cells under conditions of survival of the surrounding healthy tissue [3537]. The first experimental studies of the feasibility of magnetic material application for hyperthermia were performed in 1957. Gilchrist et al. [16] placed various specimens of tumor tissue with 20100 nm -Fe2O3 nanoparticles injected into them in an ac magnetic field with a frequency of 1.2 MHz. Then numerous publications followed describing various schemes and using various types of magnetic materials, magnetic field strengths and frequencies, methods of preparation and covering with coatings, and nanoparticle delivery [23, 26, 30, 3853]. The procedure involves concentration of magnetic nanoparticles scattered in an organism in a tumor tissue and subsequent application of an ac magnetic field of sufficient strength and frequency to cause nanoparticle heating. The high temperature caused by nanoparticle heating results in destruction of tumor if the temperature can be maintained above a therapeutic threshold of 42C for 30 min or even longer. It should be noted that the use of most magnetic hyperthermia devices is limited because of accompanying heating of a healthy tissue. Nevertheless, magnetic hyperthermia with the help of nanoparticles still remains one of the perspective methods of tumor therapy due to its selective thermal influence on local tissue regions. In a number of clinics, it is used as one of the important perspective components of experimental cancer therapy [35, 12, 23, 54, 55].

Many investigations have demonstrated the therapeutic efficiency of the given treatment for animals [56]. One of the problems in clinical application of this treatment method to oncology patients is the feasibility of delivering the required number of magnetic nanoparticles to obtain the temperature high enough to treat a tumor in clinically acceptable ac magnetic fields. Most laboratory experiments on animals model such magnetic field parameters that could not be used safely for oncology patients. In most cases, a decrease in the strength or frequency of an ac magnetic field down to a safe level often causes a significant decrease in thermal emission of the magnetic material and temperature, which makes the given method inefficient for medical applications.

It is important to understand what the main physical mechanisms induce high temperatures in small magnetic nanoparticles placed in ac magnetic fields. It is desirable that the amount of heat produced by magnetic nanoparticles was suffice to maintain the temperature of the tissue at least at a level of 42C for 30 min. In calculations of the required thermal emission norm, it is necessary to take into account that it is difficult to obtain necessary heating because of the presence of blood circulation in the tumor tissue. The above-indicated mechanisms dominate in cooling of the tissue and change very considerably during heating. A number of authors analyzed the problem of heat transfer when a certain volume of the tissue was heated from inside by uniformly distributed heat sources microscopic magnetic particles [5759]. The problem of temperature decrease due to blood circulation was not considered at all owing to its mathematical complexity and the lack of general results. In medical practice, the empirical rule is used that the thermal emission rate inside the tumor tissue should be 100 mWcm3. This approximate estimate is satisfactory in most cases.

The frequency and strength of the applied external ac magnetic field used for heating are limited by possible accompanying physiological responce of an organism to high-frequency magnetic fields [60, 61]. Among them are stimulation of peripheral and skeletal muscles and possible heart stimulation accompanying inductive heating of the tissue. Allowable ranges of magnetic field frequencies and strengths are f = 0.051.2 MHz and H = 015 kAm1. Experimental data on the use of higher-frequency fields were obtained by Oleson et al. [62] who developed a hyperthermia system based on inductive heating of tissue and Atkinson et al. [63] who developed a treatment system based on heating by eddy currents of implanted metal heat sources. Atkinson concluded empirically that local exposure to electromagnetic ac fields with strength and frequency f for which the product Hf does not exceed 4.85108 Am1s1 is safe for a patient and does not cause undesirable complications.

The amount of the magnetic material needed to produce the required amount of heat to maintain the required temperature depends largely on a concrete procedure of hyperthermia application. For example, direct intratumor injection of a magnetic liquid yields much higher concentrations of the magnetic agent in the tumor in comparison with the method of

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intravenous delivery upon exposure to a magnetic field or the method of target delivery of the magnetic agent by means of antibodies. On the other hand, the last two methods have some advantages.

It is generally believed that an intratumor concentration of magnetic material of 510 mgcm3 is sufficient for treatment of a patient using local magnetic hyperthermia. The choice of magnetic nanoparticle material, as a rule, is limited. Iron oxides Fe3O4 magnetite and -Fe2O3 are most intensively investigated nowadays owing to their good magnetic properties and biological compatibility. Alternative materials were also investigated in [6467]; however, their toxicity was not investigated, which made problematic their medical applications.

The preferable size of magnetic nanoparticles used as agents for local magnetic hyperthermia is less than 10 m, because small particles can be delivered to a tumor as a suspension in a transport liquid. The nanoparticles can be bonded to antibodies to facilitate their delivery to oncology cells in a definite way. The materials used for magnetic hyperthermia are typically ferro- or ferrimagnets. Depending on the nanoparticle sizes, they can be one- (superparamagnetic) or polydomain.

The rate of thermal emission and mechanisms of heating differ for different classes of materials, and both advantages and disadvantages are observed when the given types of materials are used as hyperthermia agents.

When hyperthermia is used in cancer therapy, the temperature of the patient body is increased up to 43 by different methods (using water, air, or electromagnetic fields) [1, 68]. Changes of such factors as and hyperglycemia accompanying hyperthermia result in a change of metabolism and weakening of tumor cells, that is, produce the antitumor effect.

The temperature-exposure regime for regional electromagnetic hyperthermia of tumors of external localization is typically 4247 inside the tumor for 12 h. In the case of generalization of tumor growth, whole body hyperthermia at 4143 is used for 2.55 h with general anesthesia and artificial ventilation of the lungs [1, 68]. It should be noted that hyperthermia itself (using hot water, air, or HF, UHF, microwave, ultrasonic, or laser radiation) without additional antitumor therapy is ineffective irrespective of the method of its application. Short-term tumor remission after hyperthermia is observed only in 1015% of cases. A combination of regional and whole-body hyperthermia with radiotheraphy (thermoradiotherapy) yields higher regression parameters (>74%) [69]. Thus, a combination of hyperthermia with beam and chemotherapy strengthens regional beam and medicinal influence on the tumor [7073]. In clinical practice, two methods of local heating of tumor tissues upon exposure to electromagnetic beams [1] are used:

1) Direct heating (capacitive or inductive) at frequencies >1 MHz when the patient body is a part of an electric circuit and different tissues of the body are heated differently depending on their electric properties;

2) Indirect (through an intermediate agent) heating of tumor tissues at frequencies from 0.05 to 1.00 MHz when ferromagnetic implants (special needles, catheters, capsules, grains, or microspheres) transforming the electromagnetic beam energy of the given frequencies into heat scattered inside the tumor serve as heating elements. The contribution of direct heating to the temperature raise of tissues in this case is usually insignificant. Methods of direct radiofrequency, ferromagnetic, and combined radiofrequency and ultrasonic heating of tissues were described in [1, 4].

Heating of tissues of experimental animals to temperatures exceeding +47 during hyperthermia causes thermal ablation (thermal destruction of cells) accompanied by acute necrosis, coagulation, and carbonation of tissues at long exposure times [1, 7375]. In clinical hyperthermia, such heating is excluded because of system complications (increase of blood pressure and heart attack) [1, 4, 10]. Therefore, for clinical whole-body hyperthermia, the temperature from +42 to +43C is conventionally used [1, 68], and for regional hyperthermia, the temperature from +42 to +46C is used [1, 4, 12].

The main efforts in the development of new methods of experimental hyperthermia are focused on optimization of temperature homogeneity in the tumor volume. More perfect systems of thermometry and magnetically controllable preparations with the Curie temperature TC = +44 are required. Above the Curie temperature, the magnetic properties of ferro- and ferrimagnetic materials are lost and hence the absorption of electromagnetic energy decreases almost to zero.

Heating of tissues at temperatures of 4344 causes in experimental animals moderate (depending on the dose) inactivation of cells in spite of the fact that the therapeutic dose-response curves obtained by registration of the results of experimental hyperthermia are similar to the curves obtained using chemotherapy. Not only DNA is destructed when performing hyperthermia, as in the case of application of alkylating preparations [76]. The remaining biomolecules of the

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cell change much more strongly for hyperthermia treatment than for chemotherapy [77, 78]. Especially sensitive to hyperthermia are regulatory proteins influencing the cell growth and differentiation and expression of receptor molecules involved in transfer of various signals. Therefore, many important changes in cells are caused by thermal inactivation of protein complexes participating in regulatory processes [68, 69, 75, 76]. Exactly hyperthermia induces primary convertible effects in cells and tissues [75]. It was proved that a few minutes after hyperthermia application, tumor cells express thermal-shock proteins that induce heat resistance of cells [76, 79]. Simultaneously, hyperthermia influences the activity of regulatory proteins, kinases, and cyclanes whose changes result in disorders of the cell cycle and can induce apoptosis [69, 7678, 80, 81].

Recent studies have been aimed at elucidation of a relationship between thermal and multiple medicinal stabilities. It is assumed that the combined effect of x-ray therapy and hyperthermia is caused by heating that leads to multifunctional reparative processes in a tumor cell after radiation-induced damages. The combined effect is less effective when heating is performed before or after damages caused by x-ray therapy, disregarding time intervals between hyperthermia, x-ray therapy, and chemotherapy [82]. The influence of hyperthermia on the tissue is accompanied by clearly pronounced changes in microvasculature and blood circulation as well in the energy state of oxygen [83].

Taking into account positive results of investigations in molecular biology and an increasing number of successful applications of experimental hyperthermia for treatment of different tumors of animals, further perfection of this method of tumor therapy is important [1, 8488]. Higher level of survival of experimental animals is achieved when hyperthermia is performed in combination with chemotherapy and sensitization of tumor tissues [8891].

All well-known methods of electromagnetic hyperthermia have definite limitations and are intended for external or interstitial application. For noninvasive whole body heating of mammals, electromagnetic systems of radio-frequency (10100 MHz) and microwave hyperthermia (>300 MHz) are used together with ultrasonic applicators and infrared radiators and baths [1, 4, 75, 92]. Depending on the geometry, arrangement, and number of applicators, energy is absorbed in surface or deep-lying tissues. For example, the BSD-2000 system supplied with the SIGMA 60 ring applicator (BSD Medical Corp., Salt Lake City, USA) is well known. For regional hyperthermia, temperature is raised in a limited region of the body. Radio-frequency hyperthermia systems are controlled on the basis of capacitive or inductive coupling of the electrodes. In modern regional hyperthermia systems, several pairs of radiators placed in the region being heated with two-capacitor control of energy distribution are used simultaneously. However, the above-considered systems generate ac electromagnetic fields propagating in a chosen direction, but neither their phase interactions, nor field strengths are controlled [83].

In regional hyperthermia systems, the temperature is monitored with the use of invasive thermometry with the help of thermistors, fiber-optical sensors, or thermocouples delivered through catheters 1.41.8 mm in diameter or implanted surgically or hypodermically. Because the three-dimensional anatomic pattern of the electromagnetic field distribution involves interfaces, for example, fat-muscle and bone-muscle, insufficiently fast control over the electromagnetic field along the body axis is a serious drawback, because it leads to the occurrence of thermal spots limiting the temperature raise in the required region, for example, because of intolerable pain, erythema, or other symptoms. To overcome the above-enumerated difficulties, systems and algorithms of calculations of electromagnetic fields were developed. Regional hyperthermia systems designed at present have much greater number of radiators, applicators, and other auxiliary adaptations compared to the already existing systems [1, 73, 92].

Radio-frequency hyperthermia systems, both microwave and ultrasonic, have limited applications. This is not only due to the interface effects but also due to high regional perfusion of tissues adjoining the tumor. Thermal convection, caused by strong blood perfusion, decreases the specific energy absorption and hence reduces the temperature of the tumor. Therefore, tumors located in regions with high perfusion, receive lower heat dose, and the hyperthermia efficiency decreases. The researchers failed to increase the temperature of tumors for cancer of the liver, lungs, and kidneys to 4445 and to maintain this temperature for a long time because of high perfusion of these organs; therefore, they are not treated by the methods of regional hyperthermia. At the same time, interface effects are the reason why cancer of the brain is treated now only with application of regional hyperthermia, including cranial trepanation. It is difficult to obtain a minimal temperature of +43 required for therapy of deep-lying tumors with the use of regional hyperthermia systems.

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In these cases, a combination of regional hyperthermia with x-ray therapy and chemotherapy strengthens the anti-tumor effect and frequently results in the desired remission. Therefore, regional hyperthermia is widely used in combination with x-ray therapy and chemotherapy.

According to experts opinion, treatment only with the help of external hyperthermia sources does not result in complete remission of tumors. Higher temperatures are achieved for interstitial hyperthermia, but the therapeutic effect in this case is limited by high temperature gradient at interfaces. To reduce the gradient, very traumatic hyperthermia methods are used [1, 73, 92].

The first ferromagnetic particles were fabricated from alloys with temperature TC higher than that required for hyperthermia [1]. However, it is desirable that TC of preparations did not exceed 45. Perfection of the preparations used for hyperthermia of tumors has led to the advent of polyfiber preparations with the specific energy absorption higher than that of usual magnetic particles of the same sizes [1, 73, 88, 92].

Advantages of ferromagnetic particles in comparison with others interstitial implants are that they can be produced in fragments having different TC to obtain asymmetric temperature structures and self-regulation to avoid overheating of normal tissues. Due to this method and three-dimensional control over heating, interstitial measurements of temperature can be avoided. To obtain a maximal calculated specific energy absorption, magnetic particles are implanted surgically being oriented exactly parallel to the magnetic field. The external ac magnetic field heating magnetic thermal grains is created mainly by rings of an inductance of various designs [1].

A magnetic liquid based on magnetite dextran was first used for treatment of rats with tumors of the mammary gland: 100 mg of magnetite with particle diameter of 6 nm were injected into the tail vein of rats for 10 min, and in 48 h, animals were exposed to an ac magnetic field (0.45 MHz) for 12 min. The temperature was measured with a thermometer during resection of the tumor performed right after termination of heating. The temperature raised by +8 for 12 min, and tumors in 11 of 12 animals necrosed. In a week, animals were killed and it was found that the most part of magnetite was concentrated in the liver, spleen, and kidneys. A certain amount of magnetite was present in tumor cells [93]. In the case of trans-arterial injection of suspension of silicoferrite particles shaped as needles 0.11 m long to rabbits with VX2 tumors with subsequent hyperthermia in an ac magnetic field, tumor necrosis was observed. After hyperthermia, rabbits lived 3 more years [94].

Glazed ceramics particles 1.5 m in diameter were injected intravenously to balb/c mice with Meth-A sarcoma; the mice were subsequently exposed to a 10-kHz ac magnetic field of 40 kA/m. Tumors were absent in 50% of mice 5 days after treatment. The number of mice that survived 497 days after treatment decreased to 12% [95]. In 10 years, the results obtained in [93] were repeatedly tested [74] for animal of the same type with the same tumors under the same conditions. Different results were obtained: no selectivity to field exposure was observed with a significant excess of the magnetite absorption by the liver, spleen, and lungs in comparison with the previous results. Studies of the tumor growth rate demonstrated a significant increase of the period of tumor volume doubling, that is, the response of the tumor to treatment was a delay of its growth. Some tumors examined 60 days after treatment were no longer dependent on the number of exposures to an ac magnetic field. No acute toxicity of ferrite was established [74].

The method of magnetic liquid hyperthermia has been developing for more than 10 years; it has passed successful tests on animals [1, 96]. In experiments on rats with an induced tumor of the brain, only two therapeutic sessions allowed the length of life of experimental animals to be increased more than four times from 8 to 35 days. In Germany, clinical tests of the method of magnetic liquid hyperthermia of tumors have already been started. Nowadays, the method is used for glioblastoma therapy. The magnetite nanoparticles are delivered to the tumor with the help of a supersensitive electronic navigation system. This allows glioblasnoma of tumors lying deeply in the brain tissue or localized near the brain fragments responsible for speech and locomotor functions to be treated [4, 45].

At present, magnetic liquid hyperthermia has a lot of problems to be solved. Among them is a problem of maintenance of a strictly preset temperature (in the interval 4547C) in the tumor volume when performing magnetic liquid hyperthermia as well as a problem of temperature monitoring. One more problem of magnetic liquid hyperthermia, in particular, is the lack of automatic control over the temperature of tumor tissues during hyperthermia.

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This problem can be solved by creation of magnetic particles with the Curie temperature equal to the temperature used for hyperthermia of tumors, that is, 4447C [88]. Under these conditions, the temperature in the tumor will be automatically maintained within the preset interval unlimited time. At present, the temperature in the tumor during application of magnetic liquid hyperthermia is regulated by changing the parameters of the ac magnetic field (in the on-off mode), and its control is performed with complex and expensive fluorooptical sensors. Application of nanoparticles with TC = 4547 can completely solve this problem. At present, the magnetite particles, in particular, with dextran coating (Fe3O4 with TC = 58.5C) are widely used as magnetic agents [1, 23, 54, 55, 86, 87, 96, 97]. A search for magnetic nanoparticles is actively conducted [6467].

CONCLUSIONS

The main efforts in the development of new methods of experimental hyperthermia are focused on optimization of temperature homogeneity in the tumor volume. More perfect thermometry systems are also required. Magnetically controllable preparations with TC = +44 are promising.

We would like to acknowledge N. A. Brusentsov, Doctor of Pharmacological Sciences, for her help and support. This work was supported in part by the Russian Foundation for Basic Research (grants Nos. 05-03-08215-ofi_a, 06-

02-01839-_b, 07-02-01513-b, and 07-08-00639-b).

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