28/08/2018 0 Σχόλια
Influence of the Q-switched and Er:YAG laser beam on the ablation of root canal dentine
Eirini Papagiakoumoua, Dimitrios N. Papadopoulos, Marouan G. Khabbazb, Mersini I. Makropoulou, Alexander A. Serafetinides
AbstractLaser based dental treatment is attractive to many researchers. Lasers in the 3 mm region, as the Er:YAG, are suitable especially for endodontic applications. In this study a pulsed free-running and Q-switched laser was used for the ablation experiments of root canal dentine. The laser beam was either directly focused on the dental tissue or delivered to it through an infrared fiber. For different spatial beam distributions, energies, number of pulses and both laser operations the quality characteristics (crater’s shape formation, ablation efficiency and surface characteristics modification) were evaluated using scanning electron microscopy (SEM).
The craters produced, generally, reflect the relevant beam profile. Inhomogeneous spatial beam profiles and short pulse duration result in cracks formation and lower tissue removal efficiency, while longer pulse durations cause hard dentine fusion. Any beam profile modification, due to laser characteristics variations and the specific delivering system properties, is directly reflected in the ablation crater shape and the tissue removal efficiency. Therefore, the laser parameters, as fluence, pulse repetition rate and number of pulses, have to be carefully adjusted in relation to the desirable result.
1. Introduction The use of lasers in dentistry was considered in the early days of medical lasers [1], as they can provide a number of clinical benefits that the conventional highspeed turbine lacks, such as pain-reduced cavity preparation, accurate and precision work and microsurgical operations.
During the last few years there is a great interest in laser treatment on the root canal dentin. Shaping of the surface of the root canal is essential for the successful outcome of the endodontic treatment. Various studies have shown that chemomechanical instrumentation, with different methods and techniques, was unable to remove totally the debris from the root canal walls [2–4].
Several laser sources have been used for dental applications, such as CO2, hydrogen-fluoride (HF) chemical laser, Er:YSSG, Er, Cr:YSSG, and Er:YAG [5,6]. To minimize the chances for thermal diffusion, during the photothermal ablation process, the laser absorption depth should be limited to the thinnest layer near the surface. The most important factors in limiting the conduction of heat from the laser exposure site to the adjoining material are: (i) the depth of energy deposition, (ii) the thermal diffusion and (iii) the change of state (i.e. vaporization). The depth of energy deposition varies with wavelength and defines the limit for minimizing thermal diffusion. If the basic absorber is water, the absorption depth is smaller for wavelengths in the mid-infrared region, where water presents maximum absorption [7]. As far as the dental hard tissue is concerned, its two other main constituents, namely, hydroxyapatite and collagen, have also absorption peaks in the 3.0 mm region [8,9]. The thermal diffusion, on the other hand, is associated with the thermal relaxation time of the material involved and the exposure duration to laser radiation. To avoid though thermal denaturation of the adjacent tissues to the laser exposed ones, high power pulses, with pulse width shorter than the thermal relaxation time, are needed [5]. But as the tissue removal per pulse is very small, high repetition rates are desirable in order to create a considerable tissue removal rate. Tissue ablation studies with picosecond and sub-picosecond laser pulses have been recently performed, providing interesting results [10–12]. These systems, emitting at wavelengths far from the main absorption peak of the water, ablate dental tissue with different mechanisms to the mid-infrared ones. Furthermore, in the majority of the cases, such systems remain rather costly for clinical use. Therefore, shorter pulses in the 3.0 mm region, at high repetition rates, are of primary importance [13]. Thus, TEA HF and Q-switched Er:YAG lasers are suitable as they can produce shorter pulses (of the order of 100 ns) than the tissue thermal relaxation time (of the order of 1 ms).
Q-switched Er:YAG laser producing 2.94 mm emission wavelength, short pulses, high power per pulse, high repetition rate, uniform reproducible beam seems to be the most advantageous, in terms of efficiency, pulse duration and spatial beam quality. Moreover, transmission of the 3.0 mm wavelength beams through flexible waveguides and fibers that is necessary for the application of lasers in dental treatments, is these days possible [14,15]. Especially as far as the transmission of the Er:YAG laser radiation is concerned, many studies have been made and several materials, appropriate for use in a clinical environment, have been tested in order to achieve low attenuation values, high power and short duration laser pulses delivery [16]. However, the transmission of laser radiation through waveguides and fibers induces changes in the temporal and mainly in the spatial beam profile. Therefore, it is very interesting to study how any alteration of the laser parameters (pulse duration, repetition rate, beam energy and spatial distribution of the beam profile) affects the cutting and ablating operations. It is expected that the result in the crater’s shape, depth and morphology is characteristic for the combination of these parameters. Especially, the laser beam quality delivered to the biological tissue may affect the crater, influencing both its shape and morphology [17,18]. The Er:YAG laser can produce a high quality beam in comparison to the HF and the CO2 laser, which are the other possible laser candidates for such applications.
The aim of this work was to investigate the influence of the Er:YAG laser beam spatial distribution in the ablation craters profile, by studying the photographic records obtained with scanning electron microscopy (SEM). The role of the mid-infrared laser pulse width was also studied and, therefore, Qswitched and free-running Er:YAG laser pulses were used, of different energies and repetition rates, delivered to the tissue either through the air or through a fiber.
2. Materials and methods 2.1. Laser sources A pulsed Er:YAG laser, emitting at 2.94 mm, in the Q-switched or free-running mode, was used, developed in our laboratory (N.T.U.A.). The laser rod was a Kigre Er:YAG rod, with dimensions 5 mm 150 mm. A laser cavity of 70 cm length was formed by an AR/75% for 2.94 mm flat/flat output coupler and a high reflector for 2.94 mm flat/flat. A Kigre K537 E. Papagiakoumou et al. / Applied Surface Science 233 (2004) 234–243 235 linear flash lamp filled with Xe (1500 Torr/5.5 in— 1850 V/56.5 J) was used to pump the crystal. In the free lasing mode, the maximum possible output energy was around 500 mJ with a pulse width of 80 ms at FWHM (full width at half maximum). The vertical and horizontal beam divergence was equal to 0.7 mrad and the maximum possible repetition rate with the available power supply was 10 Hz. By adding an frustrated total internal reflection (FTIR) prism modulator between the crystal and the output coupler inside the cavity, single Q-switched pulses were produced. The maximum possible output energy was around 80 mJ with a pulse width of 190 ns at FWHM. The beam divergence and the maximum possible repetition rate were the same as in the free lasing operation.
2.2. Sample preparation Several single rooted human teeth were used in this study. All teeth were radiographed to confirm the absence of complicated root canal anatomy and the presence of one root canal only. The crowns of the teeth were resected and root canals were measured, cleaned and shaped by the step-back preparation technique using the Flexofile (Maillefer, Switzerland) with filing motion. A 3% sodium hypochlorite solution was used as irrigant. Root canals were then dried with paper points.
Teeth were randomly divided in three groups (A, B and C). The roots were grooved longitudinally with diamond bur without penetration into the root canal and split into two halves. Teeth of group A were irradiated with Q-switched Er:YAG laser pulses of 190 ns pulse duration. Experiments were realized with different spatial distributions at several pulse energies, and thus several fluences, and the results were observed and evaluated in SEM images.
Free-running Er:YAG laser pulses of 80 ms pulse duration were applied to the root canal walls of dental samples of group B. The influence of several pulse energies and complicated spatial distributions of the beam profile, produced from different values of the pulse repetition rate was tested. The results are illustrated in SEM images.
Teeth of group C were irradiated with both Qswitched and free-running laser beams, delivered to the ablation area through sapphire optical fibers.
2.3. Er:YAG laser ablation Sequences of different fluences, in each Er:YAG laser operation, were applied to dental sample groups. In group A, the Q-switched Er:YAG laser was used, with pulse repetition rate of 1–4 Hz, energy 8–70 mJ/ pulse. Teeth were irradiated with 10–200 pulses of 190 ns duration. In group B, the free-running laser was used, with pulse repetition rate of 1–4 Hz, energy 6– 70 mJ/pulse. Teeth were irradiated with 20–200 pulses of 80 ms duration. For the irradiation of samples of group C, laser parameters in the same range with groups A and B were used. The laser beam was delivered to the tissue through 425 mm core diameter sapphire fibers of 25 cm length, with effective NA 0.1 and attenuation values of 2 dB/m. The evaluation of this kind of fibers and their performance details have been reported elsewhere [19]. Transmission rates of 88% were measured.
The laser beam was focused on the sample by means of a focusing lens (f ¼ 15:0 cm). Whenever the fiber was used, the beam was focused through a 300 mm diameter pinhole fixed in the focal plane, in order to achieve a nearly TEM00-like input beam. The fiber was aligned in contact to the pinhole and the tooth was placed very close or in contact to the fiber’s end. The laser energy was measured using calibrated laser pyroelectric detectors (model ED-200, Gen Tec). Images of the respective laser beam spatial profiles were acquired using a Pyrocam I pyroelectric array (from Spiricon) connected to a PC laser beam analyzer. During the laser irradiation, the root canal was kept constantly wet with saline solution irrigation (3 ml/min) to avoid dehydration of the tissue. After the laser irradiation all teeth were prepared for scanning electron microscopy (SEM) examination, for the evaluation of the ablation crater morphology.
3. Results 3.1. Q-switched Er:YAG laser Samples irradiated with TEM00 mode profile Qswitched laser beam showed the formation of round well-shaped craters, with dentinal tubules opened at the crater edge. Cracks were clearly observed during the SEM examination in the majority of the samples.
In all teeth used no carbonization or extended melting zones were observed. The ablation started at a fluence value of 2.3 J/cm² . At higher pulse fluences the ablation rates increased quickly quasi-logarithmically while plasma formation was macroscopically observed accompanied by abrupt saturation of the ablation rates for energy fluence more than 10 J/cm² .
In Figs. 1 and 2, SEM images of samples irradiated with beam of irregular spatial energy distribution (TEM01 and TEM03 mode profile beam, respectively), are presented. The spatial beam distribution changes, from the one of Fig. 1 to that of Fig. 2, due to the higher repetition rate, while the energy remains almost the same (50 and 40 mJ/pulse, respectively). In these images, as we can see, the crater shape represents more or less, depending on the energy and the number of pulses, the beam profile. More efficient tissue ablation has occurred by the main lobes of the beam. Dentinal tubules were clear without any obscurity from fused dentine. However, cracks were observed in most of the samples even for lower energies (3.5 40 mJ/pulse).
At this point we should notice that when a higher order multimode beam profile was used the degree of delineation was strongly depended on the applied energy. For a multimode profile the definition of an overall fluence based on the ratio of total energy to total area, although retains its practical meaning, becomes inaccurate. Different parts of multimode beams may show ‘‘local’’ fluences with great difference between them. Thus, above a certain energy level, related to the specific spatial distribution of the beam, it is possible that most parts of the beam exceed the saturation fluence threshold (due to plasma shielding) exhibiting, therefore, similar ablation results. In this way, even high multimode beams may produce smoothed crater’s shape. In case of lower energy, so that the fluences of the different parts of the beam are not in the saturation area, their partial ablation efficiency per pulse follows a quick quasilogarithmic rise. This results in the characteristic delineation of the beam profile on the tissue. The above effect is clearly presented in Fig. 2, where in the first case (Fig. 2a) smoothing of the tissue surface is observed from a multimode beam of 70 mJ/pulse energy, while Fig. 2b shows strong delineation of the same beam profile but for 40 mJ/pulse energy.
Measurements with samples of croup C have been made by transmitting the Q-switched Er:YAG laser radiation through a sapphire fiber of 425 mm core diameter and 25 cm length. The output beam exhibited a strongly spiking, highly multimode, profile. At these fiber experiments we did not sprinkle the root canal as the tooth was placed too close or in contact to the fiber. This procedure caused deep cracks on the tissue when a large number of pulses were applied.
Fig. 3a shows an example, where the ablation crater is the result of 200 pulses with 10 mJ energy each and the pulse repetition rate set to 1 Hz. In this case the tooth were placed at 2 mm distance from the fiber output end. No plasma formation was macroscopically observed during the irradiation. All the energy was absorbed from the tissue and considerable material removal happened, while deep cracks were produced. On the contrary, when the tooth was placed in contact to the fiber output end, with the same laser parameters, plasma was created in the fiber-tissue interface. A minimal tissue removal has occurred and no cracks were observed when a small number of pulses was used. An example of this case is shown in Fig. 3b, where the tissue was irradiated with 100 pulses but no sufficient tissue material was removed.
3.2. Free-running Er:YAG laser
For the comparison of the root canal ablation results with laser pulses of the same wavelength but different pulse duration, experiments were performed with freerunning Er:YAG laser. When a TEM00 laser beam profile was applied on the tissue well-shaped craters were formed. The ablation started at a fluence value of 3.5 J/cm² . At higher pulse fluences the behavior of the ablation rates was the same as in the case of the Qswitched laser. A quick increment was observed and abrupt saturation of the ablation rates with increasing fluence more than 6 J/cm² was noticed. In Figs. 4 and 5, samples irradiated with free-running Er:YAG laser are presented. Obscured dentinal tubules at all levels of the crater’s wall led to a periodical rippled structure on the tissue surface. This ripple formation was present in all the samples irradiated with free running laser pulses. Small cracks were also observed only in few samples, when energies from 50-70 mJ/pulse or large number of pulses (over 150) were used (Fig. 5).
Once more, as happened also when the Q-switched laser was employed, the degree of delineation was depended on the energy level. When low energy (20 mJ/pulse) was applied the crater’s shape resembled that of the beam profile (Fig. 5a). By increasing the energy (70 mJ/pulse) the abnormalities, caused by the odd profile, became smoother (Fig. 5b).
Finally, experiments with transmission of the free-running Er:YAG laser radiation through the 425 mm core diameter sapphire fiber, have been performed, resulting to the same spiking character of the crater’s shape, as in the case of the Q-switched Er:YAG laser.
4. Discussion The main objective of this study was to evaluate the influence of the Q-switched Er:YAG (190 ns duration) spatial beam energy distribution on the ablation efficiency per pulse of the laser. The fluence and the repetition rate involvement on the root canal surface ablation were also investigated. For comparison, irradiation of samples with 80 ms free-running Er:YAG laser beam pulses, with similar experimental parameters, has been performed. From our results it was shown that the ablation started at a fluence value of 2.3 J/cm² for the Q-switched Er:YAG laser and at 3.5 J/cm² for the free-running laser. Previous studies by Majaron et al. [20] concluded that a change in the pulse duration in the range of 50 ms to 1 ms had no effect on the ablation threshold of dentine. However, in the study of Apel et al. [21] is shown that the ablation threshold value 10 J/cm² was reduced to 6 J/ cm² when shorter pulses were applied during experiments in the 100–700 ms time range. In our study, a clear reduction of the ablation rate threshold of about 30% was also observed for the ns pulses of the Qswitched Er:YAG laser, a pulse duration well below the 100 ms of the previous study.
At higher pulse fluences the ablation rates increased quickly quasi-logarithmically. This phenomenon is a result of absorption and scattering of the laser beam by the ejected debris, as it was also reported by others [17,22]. Abrupt saturation of the ablation rates with increasing fluence above 6 J/cm² for the free-running Er:YAG laser and 10 J/cm² for the Q-switched Er:YAG laser reveals plasma formation. Plasma absorbs most of the incident laser radiation, thereby decreasing the energy available for ablation, as it has been also suggested for CO² laser beam tissue ablation [23].
Generally, experiments showed that the degree of delineation of odd spatial energy distributions depends on the energy level. Thus, quick rise to a saturation level in the ablation rates, due to the different fluence levels in a multimode beam profile, is clearly reflected in the beam profile delineation on the tissue. Consequently, in a clinical application one should know the saturation levels of the ablation rates of the specific tissue that wants to irradiate, in order to choose the appropriate laser parameters. The fact that in a clinical application the laser pulses are not delivered to the same position, due to both patient and dentist movements, accentuate the necessity of careful laser use. The ablated tissue irregularities, like those observed in Figs. 2b and 5a, could randomly overlap due to pulse to pulse movement and thus influence the successful shaping of root canal walls.
It has been recommended that adjusting the pulse repetition rate minimizes thermal injury of the surrounding tissues by ensuring that the inter-pulse period is longer than the thermal relaxation time of the substrate and thus allows cooling between pulses [24,25]. Cooling of the tissue, in order to avoid thermal injuries, may also be aided by reducing the pulse duration [26]. The Q-switched Er:YAG laser, with 500 times shorter pulse width than the freerunning laser, is of great interest for applications in which extremely high surface quality of the irradiated biological tissue is the main concern. Free-running laser ablation is based on a quasi-continuous vaporization of tissue water during the laser pulse [27]. Therefore, some of the dehydrated collagen matrix still remains on the irradiated surface [28]. In contrast the Q-switched laser ablation is an explosive removal of the collagen matrix by rapidly evaporated tissue water due to the very high intensity and the short interaction time [29]. Only a small amount of heat penetrates the adjacent tissue, as the pulse mid-infrared lasers (λ ≈ 3:0 mm) exhibit strong absorption by hard dental tissues, restricting residual thermal damage to a minimum zone [30]. In our experiments, the only form of thermal alteration that we observed was fusion of dentine, which resulted in obscurity of dentinal tubules. And, according to the pre-mentioned theoretical predictions, this effect was observed only in samples irradiated in the free-lasing mode, independently of energy distribution and pulse repetition rates.
A very interesting pattern of the ablation surface, observed in several dental samples, was the circular ripple aspect of the crater. The spatial period of this rippled structure, which was closed to the irradiation wavelength (~3 mm) (shown very clearly in Fig. 4), probably indicates the occurrence of coherent structure formation processes [31]. Therefore, this ripple morphology could be considered as a result of interference between the incident and the reflected laser light with the scattered light near the interface [31]. The ripple aspect of the crater wall (grating-like patterns) was very common in polymer surfaces irradiated with laser light, particularly polarized light [32]. This formation was attributed to a periodic electromagnetic field which also results from interference of the incident beam with surface scattered radiation. To the best of our knowledge, this aspect was recorded for the first time in hard biological tissue surface.
As far as the crack formation is concerned, shorter pulses did create more and deeper cracks than those produced in few cases with the free-running Er:YAG laser. This is probably the result of stress transients produced during the short-pulsed ablation process. The laser-induced temperature rise in the primary zone of energy deposition, for pulse duration tp < d=us, where d is the primary zone of energy deposition and us is the velocity of sound in tissue, causes thermoelastic expansion of the tissue, which yields mechanical stress [33]. This procedure can be accompanied by plasma formation in the air–tissue surface. Actually, the macroscopic observations during the experimental procedure indicate that the higher Er:YAG laser power per pulse, in the Q-switched laser, created luminous and sounded plasma.
Besides the reduced thermal damage, another consequence of the short duration high power pulses is the reduction of the ablation rates to few micrometers per pulse at moderate fluences. In all the samples of group A the Q-switched Er:YAG laser demonstrated lower efficiency in tissue removal rates than the free-lasing one, at similar laser parameters. The ablation efficiency decreases when plasma is initiated on the sample surface. Plasma initiates due to the high stress transients induced by short laser pulses and ionization of the ablated material by the incident laser beam. This plasma shields the composite surface from the remaining laser energy in the tail of the incident laser pulse. The energy in the tail of the laser pulses is absorbed and reflected by the plasma plume in front of the sample, and it never reaches the surface to contribute to ablation [34]. Thus, a first consideration could be that the lower efficiency of Q-switched Er:YAG laser seems to be an obstacle for some clinical applications. However, another approximation could also be that shallower craters are created in this way and this permits a controlled ablation of the tissue.
The application of the Er:YAG laser in a clinical environment demands the efficient delivery of the laser beam through non-toxic fibers or waveguides. The propagation properties of the laser beam through large-core/multimode fibers and preexistent or induced by the use, input/output end and bulk fiber imperfections result in multimode output laser beam even when a TEM00 input laser beam is launched into the fiber. This consists the motivation for the investigation of the beam profile delineation on the tissue when fibers are used.
The use of optical fibers to transmit the laser beam to the tissue, in an effort to simulate real clinical conditions, gave us encouraging preliminary results. From the craters observed in both Q-switched and free-running Er:YAG laser seems that the result depends a lot on whether the laser radiation is delivered in the form of a free beam or whether comes through a fiber in contact with the tissue surface. According to Mu¨ller et al. [35], in the first case the microscale fast thermal explosion can expand into the open hemisphere and the ablation products will take a large amount of the energy away. In the fiber contact case the explosion happens in a closed chamber and the energy cannot escape. Therefore, the pressure increases much more and in addition to the direct photoablation effect the non-irradiated surrounding tissue becomes destroyed by a photohydraulic effect [36].
It should be noted that there are differences in holes created with the fiber in contact and not in contact with the tissue. A restriction factor of the ablation efficiency is the intense plasma formation in the area between the fiber’s output end and the tissue’s surface, which is strongly related to their distance. The above should be taken into account in consideration with the cone/cave crater’s shape observations by Mu¨ller [35].
5. Conclusions The interest in laser based endodontic treatment procedures (e.g. root canal enlargement) is continuously increasing, especially during the last years, as lasers offer convenient, fast and almost painless application and their experimental use is very encouraging. Promising hard dental tissue laser ablation systems are considered to be the mid-infrared lasers, like the Er:YAG laser. Ablation of root canal dentine using Er:YAG laser is an efficient procedure resulting in a very good crater quality, when the appropriate parameters are used. Moreover, short pulse durations provided by Q-switched Er:YAG laser systems seem to improve the ablation process in terms of reduced thermal damage.
The ablation rates were measured with the application of gaussian-like laser profiles to the tissue for both free-running and Q-switched laser operations. The ablation threshold values defined were 3.5 J/cm² for the Q-switched laser and 2.3 J/cm² for the free-running laser, while, in both cases, saturation, due to plasma formation, was observed above 10 J/cm² and 6 J/cm² fluences, respectively.
Well-defined, spatially uniform laser pulses allow accurate determination of material removal with good post-ablative tissue characteristics. Any beam profile modification, due to laser fatigue, resonator instabilities, misalignment, or elevated energy output characteristics, is directly reflected in the ablation crater shape and influences the tissue removal efficiency.
As a result of the different nature of the laser-tissue interaction between the two different laser operations, Q-switched and free-lasing, two main advantages of the short pulse ablation are deduced. First, the ablation rates with giant pulses are reduced to few micrometers per pulse at moderate fluences. This means that shallower craters are created and the ablation rate can be controlled more easily. Adjusting the pulse energy and/or the number of pulses can produce desirable crater depths. Secondly, the tissue ablation is performed with reduced lateral thermal damage and effects like fusion or carbonization are unlikely to happen.
However, giant pulses increase the possibility of stress transients creation due to stress confinement and this may lead to cracks formation in the crater’s edge or even at the bottom of the crater, when high powers or a large number of pulses are used. Still, a detailed histological analysis of the irradiated tissue has to be performed in order to detect deeper deteriorations that could be due to mechanical stress.
Use of fiber optics to enable laser beam propagation into the oral cavity clinical application of laser based dental treatment, is desirable. Results of the present study showed that someone has to be very careful, since a damaged end surface, defective points of the fiber core, variations of the core diameter, microbends, launching conditions and other numerous reasons are responsible for the extensive mode coupling in a fiber, which spoils the homogenous energy distribution applied on the tissue. Higher order modes can be excited and propagated even when a TEM00 input beam is used. This inevitable restriction on the beam profile quality in a clinical application necessitates the very careful control of the laser parameters (fluence, pulse duration, repetition rate, number of pulses) for the desirable result.
References [1] R.H. Stern, R.F. Sognaes, J. Dent. Res. 43 (1964) 873. [2] B.E. Mayer, O.A. Peters, F. Barbakow, Int. Endod. J. 35 (2002) 582. [3] M. Hulsmann, M. Schade, F. Schafers, Int. Endod. J. 34 (2001) 538.[4] J.F. Siqueira, M.C. Araujo, P.F. Garcia, R.C. Fraga, C.J. Saboia Dantas, J. Endodont. 23 (1997) 499. [5] M.L. Wolbarsht, IEEE J. Quantum Electron. 20 (1984) 1427.[6] P. Rechmann, D.S. Goldin, T. Hennig, Lasers in dentistry IV, Proc. SPIE 3248 (1998) 2. [7] J.G. Bayly, V.B. Karther, W.H. Stevens, Infrared Phys. 14 (1963) 211. [8] F.A. Miller, C.H. Wilkins, Anal. Chem. 24 (1952) 1253. [9] D.B. Doyle, E.G. Bendit, E.R. Blout, Biopolymers 14 (1975) 937. [10] B.M. Kim, M.D. Feit, A.M. Rubenchik, E.J. Joslin, J. Eichler, P.C. Stoller, L.B. Da Silva, Appl. Phys. Lett. 76 (2000) 4001. [11] J. Kru¨ger, W. Kautek, H. Appl. Phys. A 69 (1999) S403. [12] A.A. Serafetinides, M.I. Makropoulou, A.K. Kar, M. Khabbaz, Laser applications in medicine and dentistry, Proc. SPIE 2922 (1996) 200. [13] H. Lubatschowski, U. Hetzel, A. Ho¨gele, O. Kermani, C. Ziolek, W. Drommer, W. Ertmer, Medical applications of lasers in dermatology, opthalmology, dentistry and endoscopy, Proc. SPIE 3192 (1997) 175. [14] A.A. Serafetinides, E.T. Fabrikesi, G. Chourdakis, N. Anastasopoulou, Biomedical sensors, fibers and optical delivery systems, Proc. SPIE 3570 (1998) 28. [15] N. Anastasopoulou, C. Ziolek, A.A. Serafetinides, H. Lubatschowski, Optics Commun. 186 (2000) 167.[16] D.N. Papadopoulos, E. Papagiakoumou, A.A. Serafetinides, Optics in health care and biomedical optics: Diagnostics and treatment, Proc. SPIE 4916 (2002) 415.
[17] U. Keller, R. Hibst, Lasers Surg. Med. 9 (1989) 338.
[18] J. Meister, C. Apel, R. Franzen, N. Gutknecht, Lasers Med. Sci. 18 (2003) 112.
[19] E. Papagiakoumou, D.N. Papadopoulos, A.A. Serafetinides, SPIE Proceedings of the 3rd GR-I International Conference on New Laser Technologies and Applications, vol. 5131, 2002, p. 314.
[20] B. Majaron, M. Lukacˇ, D. Sˆusˆtercˇicˇ, N. Funduk, U. Skalericˇ, Laser applications in medicine and dentistry, Proc. SPIE 2922 (1996) 233. [21] C. Apel, R. Franzen, J. Meister, H. Sarrafzadegan, S. Thelen, N. Gutknecht, Lasers Med. Sci. 17 (2002) 253.
[22] B. Majaron, D. Sˆusˆtercˇicˇ, M. Lukacˇ, U. Skalericˇ, N. Funduk, Appl. Phys. B 66 (1998) 479.
[23] M. Forrer, M. Frenz, V. Romano, H.J. Altermatt, H.P. Weber, A. Silenok, M. Istomyn, V.I. Konov, Appl. Phys. B56 (1993) 104.
[24] J.T. Walsh, T.J. Flotte, R.R. Anderson, et al., Lasers Surg. Med. 8 (1987) 108.
[25] T.C. Vangsness, T. Watson, V. Saadatmanesh, et al., Lasers Surg. Med. 16 (1995) 61.
[26] W. Seka, D. Fried, J.D.B. Featherstone, et al., J. Den. Res. 74 (1995) 1086.
[27] A. Ho¨gele, C. Ziolek, H. Lubatschowski, S. Lohmann, H. Weilling, A. Olmes, W. Ertmer, Laser-Optik 29 (1997) 45.
[28] A. Olmes, H.G. Franke, M. Raible, H. Lubatschowski, W. Ertmer, E. Ba¨nsch, G. Dziuk, Laser-tissue interaction and tissue optics II, Proc. SPIE 2923 (1996) 144.
[29] J.L. Boulnois, Lasers Med. Sci. 1 (1986) 47.
[30] E.N. Sobol, A.A. Serafetinides, M.I. Makropoulou, E. Helidonis, G. Kavalos, Lasers Med. Sci. 10 (1995) 173.
[31] D. Ba¨uerle, Laser Processing and Chemistry, Part 5, SpringerVerlag, 2000 (Chapter 28).
[32] H. Sato, S. Nishio, Jou. Pho. Pho. C 2 (2001) 139.
[33] S.L. Jacques, Appl. Opt. 32 (1993) 2447.
[34] T. Dumore, D. Fried, Lasers Surg. Med. 27 (2000) 103.
[35] G. Mu¨ller, K. Do¨rschel, H. Kar, Lasers Med. Sci. 6 (1991) 241.
[36] R. Shrinivasan, K. Casey, IEEE J. Quantum Electron. 26 (1990) 2279.Q_switched_and_free_running_ErYAG_laser.pdf
AbstractLaser based dental treatment is attractive to many researchers. Lasers in the 3 mm region, as the Er:YAG, are suitable especially for endodontic applications. In this study a pulsed free-running and Q-switched laser was used for the ablation experiments of root canal dentine. The laser beam was either directly focused on the dental tissue or delivered to it through an infrared fiber. For different spatial beam distributions, energies, number of pulses and both laser operations the quality characteristics (crater’s shape formation, ablation efficiency and surface characteristics modification) were evaluated using scanning electron microscopy (SEM).
The craters produced, generally, reflect the relevant beam profile. Inhomogeneous spatial beam profiles and short pulse duration result in cracks formation and lower tissue removal efficiency, while longer pulse durations cause hard dentine fusion. Any beam profile modification, due to laser characteristics variations and the specific delivering system properties, is directly reflected in the ablation crater shape and the tissue removal efficiency. Therefore, the laser parameters, as fluence, pulse repetition rate and number of pulses, have to be carefully adjusted in relation to the desirable result.
1. Introduction The use of lasers in dentistry was considered in the early days of medical lasers [1], as they can provide a number of clinical benefits that the conventional highspeed turbine lacks, such as pain-reduced cavity preparation, accurate and precision work and microsurgical operations.
During the last few years there is a great interest in laser treatment on the root canal dentin. Shaping of the surface of the root canal is essential for the successful outcome of the endodontic treatment. Various studies have shown that chemomechanical instrumentation, with different methods and techniques, was unable to remove totally the debris from the root canal walls [2–4].
Several laser sources have been used for dental applications, such as CO2, hydrogen-fluoride (HF) chemical laser, Er:YSSG, Er, Cr:YSSG, and Er:YAG [5,6]. To minimize the chances for thermal diffusion, during the photothermal ablation process, the laser absorption depth should be limited to the thinnest layer near the surface. The most important factors in limiting the conduction of heat from the laser exposure site to the adjoining material are: (i) the depth of energy deposition, (ii) the thermal diffusion and (iii) the change of state (i.e. vaporization). The depth of energy deposition varies with wavelength and defines the limit for minimizing thermal diffusion. If the basic absorber is water, the absorption depth is smaller for wavelengths in the mid-infrared region, where water presents maximum absorption [7]. As far as the dental hard tissue is concerned, its two other main constituents, namely, hydroxyapatite and collagen, have also absorption peaks in the 3.0 mm region [8,9]. The thermal diffusion, on the other hand, is associated with the thermal relaxation time of the material involved and the exposure duration to laser radiation. To avoid though thermal denaturation of the adjacent tissues to the laser exposed ones, high power pulses, with pulse width shorter than the thermal relaxation time, are needed [5]. But as the tissue removal per pulse is very small, high repetition rates are desirable in order to create a considerable tissue removal rate. Tissue ablation studies with picosecond and sub-picosecond laser pulses have been recently performed, providing interesting results [10–12]. These systems, emitting at wavelengths far from the main absorption peak of the water, ablate dental tissue with different mechanisms to the mid-infrared ones. Furthermore, in the majority of the cases, such systems remain rather costly for clinical use. Therefore, shorter pulses in the 3.0 mm region, at high repetition rates, are of primary importance [13]. Thus, TEA HF and Q-switched Er:YAG lasers are suitable as they can produce shorter pulses (of the order of 100 ns) than the tissue thermal relaxation time (of the order of 1 ms).
Q-switched Er:YAG laser producing 2.94 mm emission wavelength, short pulses, high power per pulse, high repetition rate, uniform reproducible beam seems to be the most advantageous, in terms of efficiency, pulse duration and spatial beam quality. Moreover, transmission of the 3.0 mm wavelength beams through flexible waveguides and fibers that is necessary for the application of lasers in dental treatments, is these days possible [14,15]. Especially as far as the transmission of the Er:YAG laser radiation is concerned, many studies have been made and several materials, appropriate for use in a clinical environment, have been tested in order to achieve low attenuation values, high power and short duration laser pulses delivery [16]. However, the transmission of laser radiation through waveguides and fibers induces changes in the temporal and mainly in the spatial beam profile. Therefore, it is very interesting to study how any alteration of the laser parameters (pulse duration, repetition rate, beam energy and spatial distribution of the beam profile) affects the cutting and ablating operations. It is expected that the result in the crater’s shape, depth and morphology is characteristic for the combination of these parameters. Especially, the laser beam quality delivered to the biological tissue may affect the crater, influencing both its shape and morphology [17,18]. The Er:YAG laser can produce a high quality beam in comparison to the HF and the CO2 laser, which are the other possible laser candidates for such applications.
The aim of this work was to investigate the influence of the Er:YAG laser beam spatial distribution in the ablation craters profile, by studying the photographic records obtained with scanning electron microscopy (SEM). The role of the mid-infrared laser pulse width was also studied and, therefore, Qswitched and free-running Er:YAG laser pulses were used, of different energies and repetition rates, delivered to the tissue either through the air or through a fiber.
2. Materials and methods 2.1. Laser sources A pulsed Er:YAG laser, emitting at 2.94 mm, in the Q-switched or free-running mode, was used, developed in our laboratory (N.T.U.A.). The laser rod was a Kigre Er:YAG rod, with dimensions 5 mm 150 mm. A laser cavity of 70 cm length was formed by an AR/75% for 2.94 mm flat/flat output coupler and a high reflector for 2.94 mm flat/flat. A Kigre K537 E. Papagiakoumou et al. / Applied Surface Science 233 (2004) 234–243 235 linear flash lamp filled with Xe (1500 Torr/5.5 in— 1850 V/56.5 J) was used to pump the crystal. In the free lasing mode, the maximum possible output energy was around 500 mJ with a pulse width of 80 ms at FWHM (full width at half maximum). The vertical and horizontal beam divergence was equal to 0.7 mrad and the maximum possible repetition rate with the available power supply was 10 Hz. By adding an frustrated total internal reflection (FTIR) prism modulator between the crystal and the output coupler inside the cavity, single Q-switched pulses were produced. The maximum possible output energy was around 80 mJ with a pulse width of 190 ns at FWHM. The beam divergence and the maximum possible repetition rate were the same as in the free lasing operation.
2.2. Sample preparation Several single rooted human teeth were used in this study. All teeth were radiographed to confirm the absence of complicated root canal anatomy and the presence of one root canal only. The crowns of the teeth were resected and root canals were measured, cleaned and shaped by the step-back preparation technique using the Flexofile (Maillefer, Switzerland) with filing motion. A 3% sodium hypochlorite solution was used as irrigant. Root canals were then dried with paper points.
Teeth were randomly divided in three groups (A, B and C). The roots were grooved longitudinally with diamond bur without penetration into the root canal and split into two halves. Teeth of group A were irradiated with Q-switched Er:YAG laser pulses of 190 ns pulse duration. Experiments were realized with different spatial distributions at several pulse energies, and thus several fluences, and the results were observed and evaluated in SEM images.
Free-running Er:YAG laser pulses of 80 ms pulse duration were applied to the root canal walls of dental samples of group B. The influence of several pulse energies and complicated spatial distributions of the beam profile, produced from different values of the pulse repetition rate was tested. The results are illustrated in SEM images.
Teeth of group C were irradiated with both Qswitched and free-running laser beams, delivered to the ablation area through sapphire optical fibers.
2.3. Er:YAG laser ablation Sequences of different fluences, in each Er:YAG laser operation, were applied to dental sample groups. In group A, the Q-switched Er:YAG laser was used, with pulse repetition rate of 1–4 Hz, energy 8–70 mJ/ pulse. Teeth were irradiated with 10–200 pulses of 190 ns duration. In group B, the free-running laser was used, with pulse repetition rate of 1–4 Hz, energy 6– 70 mJ/pulse. Teeth were irradiated with 20–200 pulses of 80 ms duration. For the irradiation of samples of group C, laser parameters in the same range with groups A and B were used. The laser beam was delivered to the tissue through 425 mm core diameter sapphire fibers of 25 cm length, with effective NA 0.1 and attenuation values of 2 dB/m. The evaluation of this kind of fibers and their performance details have been reported elsewhere [19]. Transmission rates of 88% were measured.
The laser beam was focused on the sample by means of a focusing lens (f ¼ 15:0 cm). Whenever the fiber was used, the beam was focused through a 300 mm diameter pinhole fixed in the focal plane, in order to achieve a nearly TEM00-like input beam. The fiber was aligned in contact to the pinhole and the tooth was placed very close or in contact to the fiber’s end. The laser energy was measured using calibrated laser pyroelectric detectors (model ED-200, Gen Tec). Images of the respective laser beam spatial profiles were acquired using a Pyrocam I pyroelectric array (from Spiricon) connected to a PC laser beam analyzer. During the laser irradiation, the root canal was kept constantly wet with saline solution irrigation (3 ml/min) to avoid dehydration of the tissue. After the laser irradiation all teeth were prepared for scanning electron microscopy (SEM) examination, for the evaluation of the ablation crater morphology.
3. Results 3.1. Q-switched Er:YAG laser Samples irradiated with TEM00 mode profile Qswitched laser beam showed the formation of round well-shaped craters, with dentinal tubules opened at the crater edge. Cracks were clearly observed during the SEM examination in the majority of the samples.
In all teeth used no carbonization or extended melting zones were observed. The ablation started at a fluence value of 2.3 J/cm² . At higher pulse fluences the ablation rates increased quickly quasi-logarithmically while plasma formation was macroscopically observed accompanied by abrupt saturation of the ablation rates for energy fluence more than 10 J/cm² .
In Figs. 1 and 2, SEM images of samples irradiated with beam of irregular spatial energy distribution (TEM01 and TEM03 mode profile beam, respectively), are presented. The spatial beam distribution changes, from the one of Fig. 1 to that of Fig. 2, due to the higher repetition rate, while the energy remains almost the same (50 and 40 mJ/pulse, respectively). In these images, as we can see, the crater shape represents more or less, depending on the energy and the number of pulses, the beam profile. More efficient tissue ablation has occurred by the main lobes of the beam. Dentinal tubules were clear without any obscurity from fused dentine. However, cracks were observed in most of the samples even for lower energies (3.5 40 mJ/pulse).
At this point we should notice that when a higher order multimode beam profile was used the degree of delineation was strongly depended on the applied energy. For a multimode profile the definition of an overall fluence based on the ratio of total energy to total area, although retains its practical meaning, becomes inaccurate. Different parts of multimode beams may show ‘‘local’’ fluences with great difference between them. Thus, above a certain energy level, related to the specific spatial distribution of the beam, it is possible that most parts of the beam exceed the saturation fluence threshold (due to plasma shielding) exhibiting, therefore, similar ablation results. In this way, even high multimode beams may produce smoothed crater’s shape. In case of lower energy, so that the fluences of the different parts of the beam are not in the saturation area, their partial ablation efficiency per pulse follows a quick quasilogarithmic rise. This results in the characteristic delineation of the beam profile on the tissue. The above effect is clearly presented in Fig. 2, where in the first case (Fig. 2a) smoothing of the tissue surface is observed from a multimode beam of 70 mJ/pulse energy, while Fig. 2b shows strong delineation of the same beam profile but for 40 mJ/pulse energy.
Measurements with samples of croup C have been made by transmitting the Q-switched Er:YAG laser radiation through a sapphire fiber of 425 mm core diameter and 25 cm length. The output beam exhibited a strongly spiking, highly multimode, profile. At these fiber experiments we did not sprinkle the root canal as the tooth was placed too close or in contact to the fiber. This procedure caused deep cracks on the tissue when a large number of pulses were applied.
Fig. 3a shows an example, where the ablation crater is the result of 200 pulses with 10 mJ energy each and the pulse repetition rate set to 1 Hz. In this case the tooth were placed at 2 mm distance from the fiber output end. No plasma formation was macroscopically observed during the irradiation. All the energy was absorbed from the tissue and considerable material removal happened, while deep cracks were produced. On the contrary, when the tooth was placed in contact to the fiber output end, with the same laser parameters, plasma was created in the fiber-tissue interface. A minimal tissue removal has occurred and no cracks were observed when a small number of pulses was used. An example of this case is shown in Fig. 3b, where the tissue was irradiated with 100 pulses but no sufficient tissue material was removed.
3.2. Free-running Er:YAG laser
For the comparison of the root canal ablation results with laser pulses of the same wavelength but different pulse duration, experiments were performed with freerunning Er:YAG laser. When a TEM00 laser beam profile was applied on the tissue well-shaped craters were formed. The ablation started at a fluence value of 3.5 J/cm² . At higher pulse fluences the behavior of the ablation rates was the same as in the case of the Qswitched laser. A quick increment was observed and abrupt saturation of the ablation rates with increasing fluence more than 6 J/cm² was noticed. In Figs. 4 and 5, samples irradiated with free-running Er:YAG laser are presented. Obscured dentinal tubules at all levels of the crater’s wall led to a periodical rippled structure on the tissue surface. This ripple formation was present in all the samples irradiated with free running laser pulses. Small cracks were also observed only in few samples, when energies from 50-70 mJ/pulse or large number of pulses (over 150) were used (Fig. 5).
Once more, as happened also when the Q-switched laser was employed, the degree of delineation was depended on the energy level. When low energy (20 mJ/pulse) was applied the crater’s shape resembled that of the beam profile (Fig. 5a). By increasing the energy (70 mJ/pulse) the abnormalities, caused by the odd profile, became smoother (Fig. 5b).
Finally, experiments with transmission of the free-running Er:YAG laser radiation through the 425 mm core diameter sapphire fiber, have been performed, resulting to the same spiking character of the crater’s shape, as in the case of the Q-switched Er:YAG laser.
4. Discussion The main objective of this study was to evaluate the influence of the Q-switched Er:YAG (190 ns duration) spatial beam energy distribution on the ablation efficiency per pulse of the laser. The fluence and the repetition rate involvement on the root canal surface ablation were also investigated. For comparison, irradiation of samples with 80 ms free-running Er:YAG laser beam pulses, with similar experimental parameters, has been performed. From our results it was shown that the ablation started at a fluence value of 2.3 J/cm² for the Q-switched Er:YAG laser and at 3.5 J/cm² for the free-running laser. Previous studies by Majaron et al. [20] concluded that a change in the pulse duration in the range of 50 ms to 1 ms had no effect on the ablation threshold of dentine. However, in the study of Apel et al. [21] is shown that the ablation threshold value 10 J/cm² was reduced to 6 J/ cm² when shorter pulses were applied during experiments in the 100–700 ms time range. In our study, a clear reduction of the ablation rate threshold of about 30% was also observed for the ns pulses of the Qswitched Er:YAG laser, a pulse duration well below the 100 ms of the previous study.
At higher pulse fluences the ablation rates increased quickly quasi-logarithmically. This phenomenon is a result of absorption and scattering of the laser beam by the ejected debris, as it was also reported by others [17,22]. Abrupt saturation of the ablation rates with increasing fluence above 6 J/cm² for the free-running Er:YAG laser and 10 J/cm² for the Q-switched Er:YAG laser reveals plasma formation. Plasma absorbs most of the incident laser radiation, thereby decreasing the energy available for ablation, as it has been also suggested for CO² laser beam tissue ablation [23].
Generally, experiments showed that the degree of delineation of odd spatial energy distributions depends on the energy level. Thus, quick rise to a saturation level in the ablation rates, due to the different fluence levels in a multimode beam profile, is clearly reflected in the beam profile delineation on the tissue. Consequently, in a clinical application one should know the saturation levels of the ablation rates of the specific tissue that wants to irradiate, in order to choose the appropriate laser parameters. The fact that in a clinical application the laser pulses are not delivered to the same position, due to both patient and dentist movements, accentuate the necessity of careful laser use. The ablated tissue irregularities, like those observed in Figs. 2b and 5a, could randomly overlap due to pulse to pulse movement and thus influence the successful shaping of root canal walls.
It has been recommended that adjusting the pulse repetition rate minimizes thermal injury of the surrounding tissues by ensuring that the inter-pulse period is longer than the thermal relaxation time of the substrate and thus allows cooling between pulses [24,25]. Cooling of the tissue, in order to avoid thermal injuries, may also be aided by reducing the pulse duration [26]. The Q-switched Er:YAG laser, with 500 times shorter pulse width than the freerunning laser, is of great interest for applications in which extremely high surface quality of the irradiated biological tissue is the main concern. Free-running laser ablation is based on a quasi-continuous vaporization of tissue water during the laser pulse [27]. Therefore, some of the dehydrated collagen matrix still remains on the irradiated surface [28]. In contrast the Q-switched laser ablation is an explosive removal of the collagen matrix by rapidly evaporated tissue water due to the very high intensity and the short interaction time [29]. Only a small amount of heat penetrates the adjacent tissue, as the pulse mid-infrared lasers (λ ≈ 3:0 mm) exhibit strong absorption by hard dental tissues, restricting residual thermal damage to a minimum zone [30]. In our experiments, the only form of thermal alteration that we observed was fusion of dentine, which resulted in obscurity of dentinal tubules. And, according to the pre-mentioned theoretical predictions, this effect was observed only in samples irradiated in the free-lasing mode, independently of energy distribution and pulse repetition rates.
A very interesting pattern of the ablation surface, observed in several dental samples, was the circular ripple aspect of the crater. The spatial period of this rippled structure, which was closed to the irradiation wavelength (~3 mm) (shown very clearly in Fig. 4), probably indicates the occurrence of coherent structure formation processes [31]. Therefore, this ripple morphology could be considered as a result of interference between the incident and the reflected laser light with the scattered light near the interface [31]. The ripple aspect of the crater wall (grating-like patterns) was very common in polymer surfaces irradiated with laser light, particularly polarized light [32]. This formation was attributed to a periodic electromagnetic field which also results from interference of the incident beam with surface scattered radiation. To the best of our knowledge, this aspect was recorded for the first time in hard biological tissue surface.
As far as the crack formation is concerned, shorter pulses did create more and deeper cracks than those produced in few cases with the free-running Er:YAG laser. This is probably the result of stress transients produced during the short-pulsed ablation process. The laser-induced temperature rise in the primary zone of energy deposition, for pulse duration tp < d=us, where d is the primary zone of energy deposition and us is the velocity of sound in tissue, causes thermoelastic expansion of the tissue, which yields mechanical stress [33]. This procedure can be accompanied by plasma formation in the air–tissue surface. Actually, the macroscopic observations during the experimental procedure indicate that the higher Er:YAG laser power per pulse, in the Q-switched laser, created luminous and sounded plasma.
Besides the reduced thermal damage, another consequence of the short duration high power pulses is the reduction of the ablation rates to few micrometers per pulse at moderate fluences. In all the samples of group A the Q-switched Er:YAG laser demonstrated lower efficiency in tissue removal rates than the free-lasing one, at similar laser parameters. The ablation efficiency decreases when plasma is initiated on the sample surface. Plasma initiates due to the high stress transients induced by short laser pulses and ionization of the ablated material by the incident laser beam. This plasma shields the composite surface from the remaining laser energy in the tail of the incident laser pulse. The energy in the tail of the laser pulses is absorbed and reflected by the plasma plume in front of the sample, and it never reaches the surface to contribute to ablation [34]. Thus, a first consideration could be that the lower efficiency of Q-switched Er:YAG laser seems to be an obstacle for some clinical applications. However, another approximation could also be that shallower craters are created in this way and this permits a controlled ablation of the tissue.
The application of the Er:YAG laser in a clinical environment demands the efficient delivery of the laser beam through non-toxic fibers or waveguides. The propagation properties of the laser beam through large-core/multimode fibers and preexistent or induced by the use, input/output end and bulk fiber imperfections result in multimode output laser beam even when a TEM00 input laser beam is launched into the fiber. This consists the motivation for the investigation of the beam profile delineation on the tissue when fibers are used.
The use of optical fibers to transmit the laser beam to the tissue, in an effort to simulate real clinical conditions, gave us encouraging preliminary results. From the craters observed in both Q-switched and free-running Er:YAG laser seems that the result depends a lot on whether the laser radiation is delivered in the form of a free beam or whether comes through a fiber in contact with the tissue surface. According to Mu¨ller et al. [35], in the first case the microscale fast thermal explosion can expand into the open hemisphere and the ablation products will take a large amount of the energy away. In the fiber contact case the explosion happens in a closed chamber and the energy cannot escape. Therefore, the pressure increases much more and in addition to the direct photoablation effect the non-irradiated surrounding tissue becomes destroyed by a photohydraulic effect [36].
It should be noted that there are differences in holes created with the fiber in contact and not in contact with the tissue. A restriction factor of the ablation efficiency is the intense plasma formation in the area between the fiber’s output end and the tissue’s surface, which is strongly related to their distance. The above should be taken into account in consideration with the cone/cave crater’s shape observations by Mu¨ller [35].
5. Conclusions The interest in laser based endodontic treatment procedures (e.g. root canal enlargement) is continuously increasing, especially during the last years, as lasers offer convenient, fast and almost painless application and their experimental use is very encouraging. Promising hard dental tissue laser ablation systems are considered to be the mid-infrared lasers, like the Er:YAG laser. Ablation of root canal dentine using Er:YAG laser is an efficient procedure resulting in a very good crater quality, when the appropriate parameters are used. Moreover, short pulse durations provided by Q-switched Er:YAG laser systems seem to improve the ablation process in terms of reduced thermal damage.
The ablation rates were measured with the application of gaussian-like laser profiles to the tissue for both free-running and Q-switched laser operations. The ablation threshold values defined were 3.5 J/cm² for the Q-switched laser and 2.3 J/cm² for the free-running laser, while, in both cases, saturation, due to plasma formation, was observed above 10 J/cm² and 6 J/cm² fluences, respectively.
Well-defined, spatially uniform laser pulses allow accurate determination of material removal with good post-ablative tissue characteristics. Any beam profile modification, due to laser fatigue, resonator instabilities, misalignment, or elevated energy output characteristics, is directly reflected in the ablation crater shape and influences the tissue removal efficiency.
As a result of the different nature of the laser-tissue interaction between the two different laser operations, Q-switched and free-lasing, two main advantages of the short pulse ablation are deduced. First, the ablation rates with giant pulses are reduced to few micrometers per pulse at moderate fluences. This means that shallower craters are created and the ablation rate can be controlled more easily. Adjusting the pulse energy and/or the number of pulses can produce desirable crater depths. Secondly, the tissue ablation is performed with reduced lateral thermal damage and effects like fusion or carbonization are unlikely to happen.
However, giant pulses increase the possibility of stress transients creation due to stress confinement and this may lead to cracks formation in the crater’s edge or even at the bottom of the crater, when high powers or a large number of pulses are used. Still, a detailed histological analysis of the irradiated tissue has to be performed in order to detect deeper deteriorations that could be due to mechanical stress.
Use of fiber optics to enable laser beam propagation into the oral cavity clinical application of laser based dental treatment, is desirable. Results of the present study showed that someone has to be very careful, since a damaged end surface, defective points of the fiber core, variations of the core diameter, microbends, launching conditions and other numerous reasons are responsible for the extensive mode coupling in a fiber, which spoils the homogenous energy distribution applied on the tissue. Higher order modes can be excited and propagated even when a TEM00 input beam is used. This inevitable restriction on the beam profile quality in a clinical application necessitates the very careful control of the laser parameters (fluence, pulse duration, repetition rate, number of pulses) for the desirable result.
References [1] R.H. Stern, R.F. Sognaes, J. Dent. Res. 43 (1964) 873. [2] B.E. Mayer, O.A. Peters, F. Barbakow, Int. Endod. J. 35 (2002) 582. [3] M. Hulsmann, M. Schade, F. Schafers, Int. Endod. J. 34 (2001) 538.[4] J.F. Siqueira, M.C. Araujo, P.F. Garcia, R.C. Fraga, C.J. Saboia Dantas, J. Endodont. 23 (1997) 499. [5] M.L. Wolbarsht, IEEE J. Quantum Electron. 20 (1984) 1427.[6] P. Rechmann, D.S. Goldin, T. Hennig, Lasers in dentistry IV, Proc. SPIE 3248 (1998) 2. [7] J.G. Bayly, V.B. Karther, W.H. Stevens, Infrared Phys. 14 (1963) 211. [8] F.A. Miller, C.H. Wilkins, Anal. Chem. 24 (1952) 1253. [9] D.B. Doyle, E.G. Bendit, E.R. Blout, Biopolymers 14 (1975) 937. [10] B.M. Kim, M.D. Feit, A.M. Rubenchik, E.J. Joslin, J. Eichler, P.C. Stoller, L.B. Da Silva, Appl. Phys. Lett. 76 (2000) 4001. [11] J. Kru¨ger, W. Kautek, H. Appl. Phys. A 69 (1999) S403. [12] A.A. Serafetinides, M.I. Makropoulou, A.K. Kar, M. Khabbaz, Laser applications in medicine and dentistry, Proc. SPIE 2922 (1996) 200. [13] H. Lubatschowski, U. Hetzel, A. Ho¨gele, O. Kermani, C. Ziolek, W. Drommer, W. Ertmer, Medical applications of lasers in dermatology, opthalmology, dentistry and endoscopy, Proc. SPIE 3192 (1997) 175. [14] A.A. Serafetinides, E.T. Fabrikesi, G. Chourdakis, N. Anastasopoulou, Biomedical sensors, fibers and optical delivery systems, Proc. SPIE 3570 (1998) 28. [15] N. Anastasopoulou, C. Ziolek, A.A. Serafetinides, H. Lubatschowski, Optics Commun. 186 (2000) 167.[16] D.N. Papadopoulos, E. Papagiakoumou, A.A. Serafetinides, Optics in health care and biomedical optics: Diagnostics and treatment, Proc. SPIE 4916 (2002) 415.
[17] U. Keller, R. Hibst, Lasers Surg. Med. 9 (1989) 338.
[18] J. Meister, C. Apel, R. Franzen, N. Gutknecht, Lasers Med. Sci. 18 (2003) 112.
[19] E. Papagiakoumou, D.N. Papadopoulos, A.A. Serafetinides, SPIE Proceedings of the 3rd GR-I International Conference on New Laser Technologies and Applications, vol. 5131, 2002, p. 314.
[20] B. Majaron, M. Lukacˇ, D. Sˆusˆtercˇicˇ, N. Funduk, U. Skalericˇ, Laser applications in medicine and dentistry, Proc. SPIE 2922 (1996) 233. [21] C. Apel, R. Franzen, J. Meister, H. Sarrafzadegan, S. Thelen, N. Gutknecht, Lasers Med. Sci. 17 (2002) 253.
[22] B. Majaron, D. Sˆusˆtercˇicˇ, M. Lukacˇ, U. Skalericˇ, N. Funduk, Appl. Phys. B 66 (1998) 479.
[23] M. Forrer, M. Frenz, V. Romano, H.J. Altermatt, H.P. Weber, A. Silenok, M. Istomyn, V.I. Konov, Appl. Phys. B56 (1993) 104.
[24] J.T. Walsh, T.J. Flotte, R.R. Anderson, et al., Lasers Surg. Med. 8 (1987) 108.
[25] T.C. Vangsness, T. Watson, V. Saadatmanesh, et al., Lasers Surg. Med. 16 (1995) 61.
[26] W. Seka, D. Fried, J.D.B. Featherstone, et al., J. Den. Res. 74 (1995) 1086.
[27] A. Ho¨gele, C. Ziolek, H. Lubatschowski, S. Lohmann, H. Weilling, A. Olmes, W. Ertmer, Laser-Optik 29 (1997) 45.
[28] A. Olmes, H.G. Franke, M. Raible, H. Lubatschowski, W. Ertmer, E. Ba¨nsch, G. Dziuk, Laser-tissue interaction and tissue optics II, Proc. SPIE 2923 (1996) 144.
[29] J.L. Boulnois, Lasers Med. Sci. 1 (1986) 47.
[30] E.N. Sobol, A.A. Serafetinides, M.I. Makropoulou, E. Helidonis, G. Kavalos, Lasers Med. Sci. 10 (1995) 173.
[31] D. Ba¨uerle, Laser Processing and Chemistry, Part 5, SpringerVerlag, 2000 (Chapter 28).
[32] H. Sato, S. Nishio, Jou. Pho. Pho. C 2 (2001) 139.
[33] S.L. Jacques, Appl. Opt. 32 (1993) 2447.
[34] T. Dumore, D. Fried, Lasers Surg. Med. 27 (2000) 103.
[35] G. Mu¨ller, K. Do¨rschel, H. Kar, Lasers Med. Sci. 6 (1991) 241.
[36] R. Shrinivasan, K. Casey, IEEE J. Quantum Electron. 26 (1990) 2279.Q_switched_and_free_running_ErYAG_laser.pdf
Σχόλια