**2.2. Principles of laser thermal therapy**

under stereotactic guidance, and the size of the lesion produced as a result of laser ablation is monitored using MR thermography in real time. The minimal invasiveness of the procedure makes it a good choice in patients that cannot tolerate a large operation due to either burden

Glioblastoma is a diffuse primary brain neoplasm with poor prognosis. The invasive nature and malignant potential of this tumor make its treatment a challenge. The current standard of care management paradigm consists of a multidisciplinary approach by combining maximal safe tumor resection with subsequent chemotherapy and radiation [1]. Despite maximal treatment, survival rates of glioblastoma remain poor with median survival of 14–16 months. Recent evidence indicates, however, that the extent of resection of high-grade gliomas correlates with patient survival [2–6]. In a similar sense, laser ablation can provide effective tumor cytoreduction to maximize the effectiveness of adjuvant treatments. Furthermore, there is promising evidence that hyperthermia may have additional synergistic effects with radiation, as well as disrupt blood-brain barrier (BBB) and thus facilitate chemotherapy delivery to target tissues [7, 8].

The following chapter will focus on describing the principles of laser ablation and the equipment used to deliver laser energy to brain tumors, as well as discussing current evidence for

The concept of using heat to destroy cancerous tissue has been attempted multiple times in the past. It was, however, difficult to develop a mechanism of heat delivery to the affected tissues that would allow controlled ablation of tissues in question. One of the earliest references to the efficacy of mild hyperthermia in cancer destruction is found in 1891 in the report by Dr. Coley, an orthopedic surgeon, who made an observation of complete resolution of an inoperable sarcoma in a patient after *Streptococcus pyogenes* infection [9]. He suggested that the high fevers that accompany the illness injured cancer cells sufficiently to destroy them. He followed up that work by describing a series of 10 patients that were successfully treated with "bacterial toxin therapy" [9]. Unfortunately, his results were not reproduced by others. In the years to come, radiation therapy and chemotherapy have established themselves as mainstream treatments for cancer. It was not until 1967 when Cavaliere et al. demonstrated selective sensitivity to heat of cancer cells, thus suggesting the use of hyperthermia as part of cancer therapy [10]. Follow-up work in animal models corroborated that notion. It was demonstrated that hyperthermia preferentially affects glioma cells compared to surrounding brain tissue. Local hypoxia and more acidic microenvironment within tumor contributes to this selective sensitivity to heat of glioblastoma [11]. Furthermore, hyperthermia potentiates

Other factors, however, influence the effectiveness of hyperthermia. In vitro experiments showed that only 50% of sarcomas were responsive to hyperthermia resulting in tumor

the use of laser thermal therapy in management of high-grade gliomas.

the effects of radiation and chemotherapy observed in vitro [12–14].

of disease or poor performance status.

188 Glioma - Contemporary Diagnostic and Therapeutic Approaches

**2. Background**

**2.1. The use of hyperthermia**

Laser interstitial therapy is an ablative technique that results in tissue destruction as a result of heating. Photons produced by the laser light are absorbed by the surrounding tissues causing excitation and release of excess energy as heat [18]. Once the critical thermal threshold is reached, protein denaturation and irreversible tissue coagulation ensues resulting in permanent tissue damage.

The first lasers that were attempted for tissue ablation were ruby-based [19]. The amount of energy and thus tissue damage produced by these lasers was difficult to control. In 1983, Bown et al. first described the use of the neodymium-doped yttrium aluminum garnet (Nd:YAG) laser for tissue ablation [20]. At present, two commercially available FDA-approved systems are available for use in neurosurgery in the United States: The NeuroBlate System (Monteris Medical, Plymouth, MN) and the Visualase Thermal Therapy System (Medtronic Inc., Minneapolis, MN). Laser ablation uses laser energy in the near-infrared range where the main tissue interactions are heating and coagulation (as opposed to cutting for the CO<sup>2</sup> laser). Both commercial laser ablation systems use diode lasers—one at the Nd:YAG wavelength (1064 nm) and the other at 980 nm. Although there have been claims that one is superior to the other, these are not founded in fact [21]. Lasers in this near-infrared range only penetrate a few millimeters into brain tissue. This heat is then propagated by conduction to allow for ablation radii that may extend up to 15–20 mm.

The wavelength of the laser light is what determines the efficiency of energy transfer to the tissues and, as a result, the volume of the lesion produced. Furthermore, the duration of tissue exposure to the laser light affects the amount of energy transferred and thus the amount of heating produced, with longer duration of exposure resulting in higher temperatures achieved in exposed tissues [22–24]. The design of the optical fiber and the laser catheter further affects the properties of the laser. Initially, lasers had to be used at very low power (1–5 W) to avoid excessive heating that results in tissue charring. Improvement in laser catheter design with the development of cooling mechanisms allowed use of higher power while still protecting nearby tissues. This also provides a non-stick catheter surface that allows the laser probe to glide easily through tissues. Current cooling systems employed in laser probe design use either cooled gas system with CO2 , or a continuous flow of saline through a sheath surrounding the optic fiber.

The laser probe tip is made of either sapphire or quartz to avoid altering the optical properties of the laser light. This design results in a spherical light distribution at the tip of the probe, and as a result, thermal energy is delivered in a symmetrical ellipsoid shape that is centered along the probe axis. The NeuroBlate System, in addition to the spherical probe design, also offers a side-firing probe which allows the surgeon to robotically control the direction of maximal heat distribution and may have an advantage in treating irregularly shaped lesions, or lesions near eloquent areas.

frequency, and this difference is used to interpolate local temperature using well defined relationship [18, 33]. MR thermography does not measure the actual temperature of tissue, rather the change in temperature over time, therefore an accurate core temperature is required at the start of each ablation. The Arrhenius model is then applied to estimate the degree of tissue damage that is produced based on the temperature and the amount of time that the tissues are exposed to a given temperature. Subsequently, computer software is used to visualize the temperature damage produced in real time with accurate temporal and special resolution.

Laser Interstitial Thermal Therapy in Glioblastoma http://dx.doi.org/10.5772/intechopen.77078 191

Heating tissue results in different types of tissue damage. Several different zones of tissue damage have been described. Heating tissues to up to 40°C typically does not disrupt cellular homeostasis. Once the temperature increases in the range of 42–45°C, the cells display marked susceptibility to cellular damage [34]. This range is typically explored in hyperthermia experiments. Further increase in temperature from 46 to 60°C results in significant cytotoxicity and consequent rapid cell death [35]. At temperatures exceeding 60°C, the damage sustained by mitochondrial enzymes, as well as cellular nucleic acids and proteins is so severe that coagulative necrosis takes place [36]. Finally, heating tissues to near boiling temperatures results in charring, tissue evaporation and carbonization, that may result in life-threatening intracranial pressure increases if not immediately relieved. In addition to temperature thresholds, the length of time that the tissue exposed to a particular temperature determines the extent of tissue damage with longer exposures resulting in equivalent damage that is observed at higher temperatures [18]. For instance, heating tissues to 43°C for 2 min will result in reversible tissue damage. Whereas heating tissues to this temperature for 10 min will result in permanent

As tissue heating occurs, concentric zones of damage can be identified [18, 37–39]. In the area around the fiber, the temperatures can reach high numbers in excess of 60°C resulting in central core area of coagulative necrosis. If the temperature in the area adjacent to the fiber inadvertently reaches 100°C, tissue vaporization occurs and a pseudocavity is formed. Immediately outside the core area lies the intermediate zone of permanently damaged tissue with increased interstitial fluid content. The outermost zone of damage that represents marginal zone consists of edematous but viable brain tissue. Histologically, the marginal zone is defined by lack of evidence of apoptosis and vessel thrombosis, and containing axonal swelling, shrinking neurons, and hypertrophied endothelial cells—markers of reversible tissue injury. Following a laser ablation procedure, tissues typically exhibit an increase in size due to the presence of necrotic tissue and perilesional edema. Over time, however, the necrotic core of the lesion is replaced by granulation tissue resulting in lesion shrinkage and scar formation.

The typical appearance of high grade glioma is an irregular and heterogeneously enhancing lesion on T1-weighted images. After treatment, there are typical changes that are observed on subsequent imaging studies [40–42]. Immediately after procedure one can appreciate an area of hyperintensity within the lesion on T1-weighted MRI images. This finding corresponds to

**2.4. Biological effect of LITT**

**2.5. Radiographic appearance**

injury, and for 60 min will result in coagulative necrosis.

The Visualase system uses a 15 W 980 nm diode laser that is cooled with circulating sterile saline solution [25]. The diameter of the catheter is 1.65 mm. The laser probe tip comes with a light diffusing tip that results in spherical light distribution producing an ellipsoid area of tissue damage. This non-pulsed system produces faster lesions but the application of heat is limited to several minutes. The system is connected to a workstation that displays real-time thermography data as "thermal" and "damage" images [26, 27]. A number of safe points can be set on the pre-treatment MRI, and when the set temperature is reached at that point, the laser is deactivated.

The NeuroBlate System uses a 12 W solid-state Dornier diode laser that operates at Nd:YAG wavelength of 1064 nm [28]. The laser catheter is cooled with CO2 gas [29]. The probes come in two diameters: 3.2 and 2.2 mm. The light diffusing tip comes in two configurations: spherical, used to produce elliptical lesions along the probe axis, and side-firing probes, that enable treatment of complex and irregularly shaped lesions. The computer interface displays thermal damage as thermal-damage-threshold (TDT) lines. The yellow line represents tissue volume that is exposed to the equivalent of 43°C for 2 min, the blue line is equivalent to exposure to 43°C for 10 min, and the white line surround the volume that received the equivalent of thermal energy of 43°C for 60 min. Based on the Arrhenius equation, the higher the temperature, the less time it takes to generate each TDT-line.

### **2.3. MR thermography**

The use of laser ablation for treatment of tumors was first described by Bown in 1983 [20]. The first report of intracranial use for brain lesion laser ablation came out in 1990 [22]. Despite that, laser interstitial thermal therapy did not gain wide-spread use due to lack of the ability to monitor the extent of ablation and tissue damage. A variety of methods were attempted to measure thermal energy delivered to tissues and included skin thermometers, subcutaneous and interstitial probes, infrared detectors, and thermographic cameras, none of which were accurate enough to predict the size of the resulting thermal lesion [30–32]. Introduction of MR thermography revolutionized the application of laser thermal therapy since for the first time it allowed monitoring of the extent of tissue damage in real time [27]. The principle of MR thermography relies on detecting differential temperature-specific proton resonance frequency in the water molecules. At a given temperature, a proportion of water molecules are interconnected in space via hydrogen bonds between molecules. As the temperature of tissues increase during laser ablation, more water molecules are freed up from the hydrogen bonds between H2 O molecules. During application of the magnetic field, proton nuclei within free water molecules are mobilized more effectively resulting in a different proton resonance frequency, and this difference is used to interpolate local temperature using well defined relationship [18, 33]. MR thermography does not measure the actual temperature of tissue, rather the change in temperature over time, therefore an accurate core temperature is required at the start of each ablation. The Arrhenius model is then applied to estimate the degree of tissue damage that is produced based on the temperature and the amount of time that the tissues are exposed to a given temperature. Subsequently, computer software is used to visualize the temperature damage produced in real time with accurate temporal and special resolution.
