13.1 Introduction

Classical approaches to altering the shape of cartilage in the head, neck, and upper airway have focused on creating incisions in the cartilage to weaken it focally or using sutures to balance the forces which resist sustained deformation. Surgery as a whole is steeped in the use of these conventional approaches which require gaining access to individual cartilage specimens through incisions within the nose or neck.

Cartilage itself is a complex tissue which is triphasic in structure, having its mechanical state determined by the interplay of viscoelastic, hydrodynamic, and electrostatic forces. This specific behavior is well beyond the scope of this discussion here but has been examined in detail by Mow and Lai, among others [1–4]. Importantly, cartilage may be thought of as a charged polymer hydrogel, and if one examines cartilage from the vantage of a materials scientist, one can think of alternate ways of creating shape change, without the need potentially for either destructive techniques involving scalpels or techniques involving sutures. Early in the 1950s, Lewis Thomas, in fact, had examined the potential use of enzymes to locally disrupt the bonds between the glycosaminoglycans in the cartilage of rabbit ears and was able to demonstrate transient changes in tissue geometry [5–9]. While promising, the effect was noted to be completely reversible, which negated any further exploration of this approach toward clinical implementation.

In contemporary times, Emil Sobol of the Russian Academy of Sciences in Troitsk, while on sabbatical in Crete, had the opportunity to work with Emmanuel Helidonis, who is an otolaryngologist. Helidonis had an extensive medical laser facility, and Sobol spent his time identifying areas where advanced laser technology might optimize surgery. As an alternative to morselization, Sobol proposed that cartilage could be heated and then undergo a phase transformation that would lead to an alteration of shape. In this implementation, cartilage which is itself curved or misshapen is first mechanically deformed, and then laser energy is directed at areas where internal stress is concentrated. Focally within these regions of interest, temperature elevation leads to a local alteration in tissue mechanical properties and an acceleration of stress relaxation.

The net effect is shape change, which in this case early on was focused on changing the shape of the nasal septal cartilage. Regardless, the use of photothermal techniques to reshape cartilage is still investigational, though much basic research has focused on this application. The use of lasers to reshape cartilage has been studied in great detail, though the precise mechanism still remains elusive. The hope and goal of this technique is that traditional cut-and-suture approaches to altering cartilage shape could be replaced by methods which are minimally invasive, potentially transcutaneous, or delivered via fiber optics and small-bore needles.

Likewise, in parallel with research performed in laser reshaping, other techniques that alter the shape of facial cartilages have been developed. These include the application of radiofrequency energy as well as the creation of in situ redox reactions in the tissue. All of these approaches do share in common a fundamental difference from classic surgical technique in that these view cartilage as a plastic material.

The remainder of this chapter reviews the basic science behind facial cartilage reshaping using laser and related technologies and provides a comprehensive review of the literature.

13.2 Basic Science of Cartilage Reshaping

13.2.1 Laser Shaping of Cartilage

In living organisms, cartilage serves to support and fasten soft tissues and to absorb shock for skeletal bones. Cartilage is a dense connective tissue composed of 65–80% of water containing a small proportion of chondrocytes within an extracellular matrix (ECM). The ECM is a hydrated gel containing proteoglycans (5–15%) and collagen fibers (20–25%). The proteoglycan matrix possesses negatively charged ion groups (SO³⁻ and COO⁻ moieties). Therefore, cartilage can be thought of as a charged hydrogel where free space is filled with water in either partially bound or free states. Polarized water molecules bind weakly to the negatively charged groups attached to both proteoglycan and collagen molecules. Free water within the cartilage matrix contains dissolved minerals such as Na⁺ and Ca²⁺ ions, which are attracted to the negatively charged components of proteoglycan molecules (Fig. 13.1). These ions are therefore trapped in the matrix when free water is forced out of it during deformation or evaporation. A gradient in distribution of negatively charged groups in the tissue accounts for the intramolecular internal stresses of the tissue. Since the ECM is nonvascular, maintenance of its mechanical structure and nutrition for chondrocytes depends on the diffusion of fluids [11]. The mechanical properties of cartilage shape retention are accounted for by the development of areas of high internal stress upon mechanical deformation and the property of shape memorization whereby shape recovery is possible after deformation.

In 1993, Emil Sobol et al. were the first to describe the use of lasers in altering cartilage shape [11]. Sobol hypothesized that local laser heating may lead to increased plasticity by relaxing the internal stresses, thereby leading to shape change with fluence rate profiles below the ablation threshold. Under ordinary conditions, mechanical resistance to sustained cartilage deformation is largely due to the intermolecular forces between water and proteoglycan models. Under moderate laser heating, Sobol hypothesized that there is a momentary relaxation in internal stress when water transitions from a state in which it is bound to proteoglycans to a liberated free state. If this bound-to-free phase transition of water can occur without damaging surrounding protein or carbohydrate molecules, then a stable modified cartilage configuration may be achieved.

Sobol et al. demonstrated laser shape change on samples of cartilage from 0.2 to 1.5 mm in thickness using a CO₂ laser with 1–10 W in average power output. Cartilage samples were mechanically fixed in a new shape and exposed to either repetitively pulsed (pulse duration of 0.2 s, pulse repetition rate of 1 Hz) or continuous wave treatment regimens. Cartilage samples treated in this manner retained their shape for several months and could be implanted into animal models or kept in storage for later use, though banking of autologous cartilage is by no means common practice in North America.

13.2.2 Ex Vivo Cartilage Reshaping

After Sobol’s initial studies on the use of lasers for reshaping cartilage, more detailed studies followed to characterize the effects of laser irradiation on tissue structure and viability and to attempt to identify optimal dosimetry. In subsequent studies, a holmium:YAG (Ho:YAG) laser (2.12-μm wavelength) was used to reshape a 1-mm-thick cartilage specimen without overheating or gross destruction of tissue near the surface [10]. Confocal microscopy and histology determined that thermal injury was deeper using Ho:YAG (up to more than 1800 μm) than with an Er:YAG laser (2.94 μm) (up to 70 μm) as expected, suggesting that Ho:YAG lasers should be used judiciously with pulse energies as low as possible to reduce collateral tissue damage [12].

Atomic force microscopy images of the fine structure of cartilage following CO₂ laser irradiation showed the formation of micro-channels of 100–400 nm in cross section, lending some credence to a proposed mechanism for laser-induced stress relaxation of cartilage being based on short-time polymerization and subsequent reformation of proteoglycan units. These micro-channels may facilitate transport of traces of proteoglycan units possessing a length of ~300–400 nm and a width of ~80 nm. These results were also consistent with changes in light-scattering behavior in cartilage after laser-induced stress relaxation, whereby it is thought that the number of scattering centers first increases due to the short-time liberation of proteoglycan units and then decreases after the new proteoglycan configuration has been formed. Furthermore, sodium carbonate crystals were observed, suggesting that prolonged laser heating of cartilage induces denaturation of proteoglycans and can lead to local mineralization of the cartilaginous matrix [13].

13.2.2.1 Dosimetry Studies

In 2000, Helidonis et al. performed histologic and morphological analysis on CO₂ laser-irradiated rabbit auricular cartilage to assess shape retention and viability [14]. Straight cartilage samples were removed from the ears of 21 rabbits, and the cartilage was reshaped using CO₂ laser at an output power of 3 W, a spot diameter of 2 mm, and exposure time of 0.5 s. Remodeled cartilage, along with control cartilage, was then implanted into the rabbits’ backs and retrieved 6–12 months later, after which histology and morphological analyses revealed shape retention and viability of chondrocytes.

Subsequently, investigations were conducted to optimize parameters for cartilage reshaping and to define the therapeutic window within which cartilage is reshaped but not thermally damaged (Fig. 13.2). To this end, Wong et al. designed a computer-controlled instrument to evaluate the effect of laser dosimetry on shape change during laser-mediated cartilage reshaping using an Nd:YAG laser at a wavelength of 1.32 μm [15]. Real-time measurements of tissue optical properties and surface temperatures were obtained, and to determine optimal reshaping, the radius of curvature of the specimen was compared to that of the reshaping jig. Optimal reshaping was observed at 6 W with an irradiation time of 16 s or alternatively at 10 W with an irradiation time of 8 s.

Subsequent investigations focused on characterizing the safety of LCR by determining shape change and tissue viability as a function of laser dosimetry [16]. Similarly, Dobrikov et al. used changes in backscattered He-Ne light intensity to characterize the temperature range in which stress relaxation occurs [17]. Wong et al. utilized the same Nd:YAG laser at 1.32 μm as in previous studies but with a spot diameter of 5.4 mm instead of 5 mm; exposure times of 4, 6, 8, 10, 12, and 16 s; and powers of 4, 6, and 8 W. Cartilage surface temperature measured using infrared thermography and a live/dead viability assay combined with fluorescent confocal microscopy was used to determine the amount of thermal damage generated in irradiated specimens. Confocal microscopy identified dead cells spanning the entire cross-sectional thickness of cartilage specimen within the laser spot at laser power density and exposure times above 4 W and 6 s, with damage proportional to increases in time and irradiance. These results suggested that thermal tissue damage is concurrent with shape change and that significant cell death occurs at laser dosimetry parameters necessary to produce clinically relevant shape changes.

13.2.3 Cartilage Properties

Mow and Lai proposed a triphasic model of articular cartilage as an extension of their earlier biphasic theory [2, 3, 18]. The triphasic model consists of the following three phases: (1) an intrinsically incompressible porous permeable charged solid phase, (2) an intrinsically incompressible interstitial fluid phase, and (3) an ion phase with two monovalent ions (anion and cation). In this theory, the motive forces for water and ions are described by the gradient of chemical or electrochemical potentials. These driving forces are balanced by the frictional forces between the phases as one phase flows through the other. Stress and strain in cartilage is determined by a balance of tissue elastic properties, fluid flow or shift, and electrostatic charge. This model was later extended to incorporate multiple polyvalent ions by Gu et al. [19]

13.2.3.1 Optical Properties of Cartilage

In order to optimize laser cartilage reshaping (LACR), it was important to determine the optical properties of cartilage to facilitate computational modeling. It is well established in the literature that different types of cartilage have different structures and compositions. Sobol’s group in 1993 demonstrated that light scattering may identify a phase transformation in cartilage after laser irradiation [20]. Bagratashvili et al. demonstrated that human, pig, and bovine cartilages have similar transmission and reflection spectra, which opened the door for the development of animal models for in vivo studies [21]. Wong et al. through a series of studies that measured integrated backscattered light intensity of He:Ne laser light (λ = 632.8 nm) during laser irradiation by Nd:YAG laser (λ = 1.32 μm) observed an increase, plateau, and then decrease in diffuse reflectance during heating [22–24].

The above developments allowed for later studies by Youn et al. to further characterize optical and thermal properties of nasal septal cartilage using double integrating sphere experiments and thermocouple techniques [25]. Wong et al. demonstrated in 2001 that the tissue optical, mechanical, and biologic properties of septal cartilage varied spatially within each individual sample, as well as between animals within the same species [26]. These measurements would establish baseline values for tissue metabolism, cell density, and the basic biomechanical behavior of porcine and rabbit septal cartilage.

13.2.3.2 Thermal and Mechanical Properties of Cartilage

Stress Relaxation

From a materials science point of view, cartilage may be thought of as a charged polymer hydrogel, with a triphasic structure formed by the interplay of viscoelastic, hydrodynamic, and electrostatic forces. Early hypotheses by Sobol et al. on the mechanism of stress relaxation that occurs upon heating cartilage proposed that this phenomenon relied on a phase transformation dominated by the movement of the water through the matrix [27]. The underlying principle of this hypothesis is that water exists in two forms to form the structure of cartilage: bound, or non-exchangeable, and free, or exchangeable. Upon heating and deformation, a phase change between the two states of water within cartilage allows for stress relaxation to occur within the areas of highest stress, thereby allowing for shape change. The first of many studies to explore this hypothesis examined the thermodynamic characteristics of this “bound-to-free” phase transformation of water [28].

An important observation was made by Sobol et al. in 1997, where it was noted that light scattering increases with stress relaxation and at a temperature exceeding 70 °C [29]. Changes in light scattering were thought to represent the bound-to-free water phase transition, beginning with the formation of nucleus centers or local regions of anomalous refractive index created when water bound to large proteoglycan molecules becomes liberated. When examining the mechanical properties of cartilage following this phase transition, it was observed that stress indeed decreases with some time delay after tissue temperature reaches 70 °C, which was initially hypothesized to represent the internal friction coefficient of cartilage. These observations were used to develop theoretical models which incorporated thermal and mass transfer in a tissue to study the effect of laser irradiation, water evaporation from the surface, and the temperature dependence of the diffusion coefficient [30]. From this model, it was shown that surface temperature reaches a plateau quicker than the maximal temperature, laser-induced mass transfer in cartilage is heterogeneous along the depth, and depth of the denatured area depends on laser fluence, wavelength, exposure time, and thickness of cartilage.

Wong et al. investigated the pattern of backscattered light intensity and internal stress and found that both tend to increase, plateau, and then decrease in similar ways during laser irradiation [31]. The plateau region occurred when the cartilage surface temperature approached 65 °C. These observations clarified the potential of using backscattered light intensity to control the process of laser-assisted cartilage reshaping, which would allow for greater precision of heating, and minimize nonspecific thermal injury due to uncontrolled heating. Bagratashvili et al. examined the phase change through multiple modalities including optical coherence tomography [32]. They describe a bleaching effect similar to the increase in backscattered light described by earlier studies. This bleaching effect was due to structural alterations in irradiated cartilage caused by the removal of water ; since water and cartilage matrix have different refractive indices, removal of water leads to increases in scattering signal.

Temperature Dependence of LCR

Although early reports show this transition to occur above 70 °C, later studies have described a range of critical temperatures T_c from 60 °C to 70 °C [31, 33]. To refine this range, Wong et al. studied temperature-dependent changes in thermal properties using modulated differential scanning calorimetry (MDSC), a very precise technique to measure temperature and heat flow associated with transitions in materials as a function of temperature and time [34]. It was observed that slow heating results in a lower critical transition temperature of around 55 °C, in contrast to the rapid heating associated with laser irradiation, with a critical transition temperature of around 65 °C. At around 70 °C, it was noted that heat flow into the specimen reaches a maximum and subsequently decreases, which is in agreement with previous results regarding temperature-dependent changes in optical and mechanical properties of cartilage. Further studies by Sobol et al. demonstrated that light scattering may be useful for measuring denaturation thresholds and kinetics for biological tissues, with denaturation thresholds showing an inverse correlation with the absorption spectrum of the tissue [35].

Mechanical Properties

The elastic modulus describes the intrinsic stress-strain relationships in a material independent of geometry and is the best characterization of the mechanical behavior of cartilage. Investigations into the elastic modulus of cartilage samples contributed to the understanding of shape change in cartilage during laser irradiation and the optimization of this process. Gaon et al. determined the elastic moduli of porcine cartilage before and after Nd:YAG laser irradiation (λ = 1.32 μm, 21.22 W/cm²) and found the elastic moduli to be much lower following irradiation [36]. This result confirms that cartilage becomes more flexible after undergoing stress relaxation due to photothermal heating. Gaon et al. also examined the changes in elastic moduli after total thermal denaturation of cartilage samples, which also resulted in a lower elastic modulus. However, the mechanical changes in elastic modulus after laser irradiation were reversible, whereas those following total thermal denaturation were not. Chao et al. conducted similar investigations on the elastic modulus of rabbit nasal septal cartilage and confirmed that this pattern of reversible decreases in elastic modulus is seen in the porcine model and is also seen in the rabbit model [37].

Thermal Properties

It has been hypothesized that cartilage reshaping occurs due to a heat-induced transition that leads to the rearrangement of molecular bonds in the cartilage matrix macromolecules. Chae et al. studied the thermomechanical behavior of cartilage using dynamic mechanical analysis (DMA) and time-temperature superposition (TTS)—techniques used in the rheological sciences to characterize viscoelastic material properties, such as storage and loss modulus, and damping properties [38]. They identified a temperature transition range between 50 °C and 67 °C—consistent with previous results. By using TTS, Chae et al. were able to estimate the activation energy associated with the mechanical relaxation of cartilage as approximately 148 kJ/mole. This estimated activation energy for stress relaxation exceeds that of the evaporation of free water (41–44 kJ/mole), as well as the activation energy of water diffusion (30.6 kJ/mole). Additionally, it was found that a relatively larger activation energy was required for relatively lower concentration of water. This may be due to an increased amount of energy needed to facilitate water movement through the dense ECM and liberate bound water from the proteoglycan side groups.

In order to estimate the thermal influence on the physical shape of a cartilage sample, Wright et al. rapidly immersed porcine nasal septal cartilage in saline water baths and measured resulting bend angles [39]. This was performed to emulate uniform or bulk volumetric heating of thin cartilage specimens held in deformation. The largest bend angle was seen at 74 °C with an immersion time of 320 s. In a following study, Wright et al. examined the dependence of cartilage shape change on both temperature and laser dosimetry using laser irradiation in addition to saline bath immersion [40]. From this investigation, the critical transition temperature region was determined by the sharp increase in bend angle at consecutive times of immersion at the same temperature (59–68 °C and 62–68 °C for porcine and rabbit cartilage, respectively). As for laser irradiation, similar transition zones for dosimetry occurred below 20.4 W/cm² for both species.

In a later study by Chae et al., temperature modulate differential scanning calorimetry (TMDSC) is a technique that is used to differentiate between thermodynamic and kinetic components of heat flow [41]. From this analytic technique, two enthalpic events were identified in samples with low water loss, with the first event occurring between 50 °C and 52 °C. When water loss exceeded about 35–40%, only one endothermic event was observed. Thus, the water content of the sample has a profound effect on the temperature range of phase transformation. Laser heating of cartilage creates a localized region of dehydration within cartilage samples within the area of light distribution. This variation in water content could lead to changes in temperature thresholds for stress relaxation and thus phase transformation.

Modeling of Cartilage Reshaping

During laser irradiation of biological tissue, many important physical processes occur that determine temperature elevation and thermal damage rates. Of note are the propagation of light within a scattering media; transformation of laser light into photochemical, acoustic, or thermal energy; tissue-tissue and tissue-environment heat and mass transfer; and the occurrence of low-energy phase transformations. In order to optimize the reshaping process, it is essential to characterize the temperature-dependent stress relaxation and physical properties of cartilage, namely, elastic modulus, thermal diffusivity, and optical scattering. These processes have been used by Diaz et al. to create a finite element model (FEM) to predict the temperature distribution in a slab of porcine nasal cartilage during laser irradiation [42, 43]. These models can be used to make predictions of the onset, extent, and severity of thermal injury—information which may be used to develop dosimetry guidelines for medical applications of lasers.

In order to guide and optimize laser cartilage reshaping for clinical use in septal cartilage reshaping, Protsenko et al. used FEM to model the forces applied during cartilage straightening deformation before and after laser irradiation as a function of the number, pattern, and location of laser target sites [44]. From the FEM model, it was observed that straightening deformation produced a nonhomogeneous stress field with regions of tension and compression. With an increase in number of laser irradiation sites and delivered laser energy, it was noted that reaction force decreased. The model showed that in order to reduce reaction force by 95%, approximately 50% of thermal damage to septal cartilage would also occur.

13.2.3.3 Biophysical Properties and Cartilage Behavior

Cartilage is a charged, hydrated, protein-based polymer that is at risk for denaturation upon heating. It is well known that laser heating may result in thermal decomposition of biopolymers as intermolecular bonds break as temperature rises. The safe practice of LCR necessitates investigation into thresholds of denaturation to minimize the risk of uncontrolled thermal injury. Tissue changes due to laser irradiation may be monitored by examining the number and size of light-scattering centers in the tissue, which has been discussed in previous sections for the application of monitoring the phase transformation of water in cartilage during laser-induced stress relaxation.

Sobol et al. used data on the time dependence of light scattering in tissue to estimate the approximate values of kinetic parameters for denaturation as a function of laser wavelength and radiant exposure [45]. An inverse correlation between denaturation thresholds and the absorption spectrum of the tissue was observed for many wavelengths. This was observed except at wavelengths near 3 and 6 μm, where denaturation threshold is instead governed by heating kinetics of tissue, as the initial absorption coefficient is very high. In a following study, Sobol et al. examined the alterations in the absorption of tissue water by laser heating with a short, single laser pulse with negligible movement and evaporation of water [46]. For temperatures less than 50 °C, it was observed that the absorption coefficient for cartilage remained approximately constant. However, for temperatures above this critical threshold temperature, the absorption of coefficient decreases at a nearly constant rate. Therefore, the critical threshold temperature is the characteristic temperature for a change in the molecular structure of the tissue, and the changes observed are due to a decrease of intermolecular interaction energy, such as by the disaggregation of water molecules.

In another series of studies, Ignat’eva et al. investigated the thermal stability of collagen in cartilage and factors that may alter the degree of denaturation upon heating [47, 48]. Cartilage is composed primarily of type II collagen fibers embedded in a mesh-like network of proteoglycan fibers. In their 2004 results, Ignat’eva et al. generated the curve of endothermic melting of collagen and observed three peaks with maximums at 60, 65, and 70 °C. These peaks correspond to melting of three fractions of collagen: tropocollagen, fibril surface collagen, and fibrillary collagen, respectively. Analysis of thermal and thermomechanical behavior of the samples revealed that the initial melting point of the first fraction corresponded to the phase of softening of the preparations (40–50 °C), whereas the initial melting point of the third fraction (65 °C) corresponded to abrupt changes in sample shape. In their subsequent study, Ignat’eva et al. determined that hyaline cartilage, such as that of the nasal septum, is thermally stable and remains incompletely denatured up to 100 °C. However, partial destruction of glycosaminoglycans in hyaline cartilage leads to an increase in degree of denaturation of collagen II upon heating. Proteoglycan aggregates therefore may play a key role in creating topological hindrances for moving polypeptide chains, reducing the configurational entropy of collagen macromolecules during denaturation. Later, Hajiioannou et al. determined the distinct role of the collagen network in cartilage shape and tensile strength preservation by enzymatically incubating cartilage strips and subsequently using laser irradiation for reshaping [49]. Collagen degradation was observed to be a substantial factor leading to the release of cartilage tensile stresses.

Polarization-sensitive optical coherence tomography (PS-OCT) has been used to characterize the polarization state of backscattered light as a function of optical path length in birefringent biological tissues. Birefringence in cartilage is due to asymmetrical collagen fibril structure, and changes in birefringence may signal disruption of cartilaginous structure due to laser irradiation. Youn et al. investigated the use of PS-OCT to measure thermodynamically induced changes of phase retardation in cartilage during LACR. It was observed that the retardation of light in the cartilage sample was changed due to laser irradiation, with two possible causes: dehydration and thermal denaturation (Fig. 13.3). The two conditions were then tested via either dehydration in glycerol or thermal denaturation in heated physiological saline. The results suggested that the observed retardation changes in cartilage were primarily due to dehydration.

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