Pulpal Temperature Changes during Low-power Hard-tissue CO2
Laurence J. WALSH
University of Queensland Dental School, Brisbane, QLD, Australia
Braz Dent J (1996) 7(1): 5-11 ISSN 0103-6440
| Introduction | Material
and Methods | Results | Discussion
| Acknowledgments | References
Thermal insult to pulpal tissue is recognized as a major limitation
to the use of lasers for dental hard-tissue procedures. This study examined
thermal changes at the level of the dental pulp in human molar teeth irradiated
with a CO2 dental laser using a pulsed mode of operation. Sectioned
molar teeth were exposed, in vitro, to CO2 laser radiation.
The laser parameters were those used clinically for laser desensitization
and laser-enhanced fluoride treatment. Fissure regions and root surfaces
were irradiated. For settings which might reasonably be used clinically,
the temperature rise was not of a magnitude which would be expected to
cause pulpal inflammation or necrosis. With regard to thermal properties
of tooth structure, times taken to reach the maximum temperature reduced,
and times taken to cool to baseline increased with increasing laser exposures.
Key Words: lasers, dental pulp, thermal stress.
Carbon dioxide (CO2) lasers have attracted interest in recent
years for use in various dental hard-tissue procedures. CO2
laser radiation is within the infra-red region of the electromagnetic spectrum,
with a wavelength of 10.6 mm. This laser wavelength is absorbed strongly
in dental enamel, dentin and cementum (Featherstone and Nelson, 1987).
Consequently, useful working temperatures of hundreds to thousands of degrees
Celsius may be generated at the laser impact site, with, theoretically,
minimal transmission of thermal energy into the pulp (Fowler and Kuroda,
1986). Indeed, the temperature rise in the pulp chamber/root canal has
been shown, for the continuous wave mode of emission, to be minimal (Powell
et al., 1989). This likely reflects the inherent thermal conductivity of
dentin. Parameters for thermal injury to dental pulp have been defined
by Zach and Cohen (1965), who examined histological changes in pulps of
monkeys exposed to a non-laser heat source (soldering iron). Their work
defined a “critical” threshold for temperature rise of 10oF (5.5oC), above
which an unacceptably high incidence of pulpal necrosis occurred. Below
5.5oC, reversible and mild pulpitis occurred. Below 4oF (2.2°C), no
histological changes were discernible. Based on these threshold values,
subsequent studies using continuous wave CO2 lasers as a heat
source have concluded that exposures should not exceed 10 Joules (J), to
ensure that the “critical” temperature increase of 5.5°C does not occur
(Miserendino et al., 1989; Powell et al., 1989). These same studies determined
that, to avoid pulpal necrosis, continuous wave exposures should not exceed
30 J. However, there are no data available on the thermal effects of CO2
lasers operated in pulsed (“chopped”) modes. From a purely theoretical
standpoint, pulsed modes of operation are safer, in that cooling can occur
between laser pulses (Walsh, 1993). The present study was undertaken to
quantify temperature changes during lasing of crown and root regions using
pulsed lasing modes identical to those employed for laser desensitization
and enhanced fluoride uptake. The study had two aims. Firstly, to determine
which settings fell within the pulpal safety limits defined by Zach and
Cohen (1965), and secondly, to examine what differences existed between
temperature rises with different pulsed modes. The times taken for the
maximum temperature to be reached, as well as the times to return to baseline
temperature (“thermal relaxation”) were also examined.
Material and Methods
A Luxar LX-20I dental laser was used (Luxar Corporation, Bothell, Washington,
USA). The beam was delivered via a flexible hollow waveguide to which a
straight handpiece with a ceramic tip was attached. The laser spot size
was 0.8 mm. The laser power output was monitored, and the delivery system
was calibrated twice daily. The laser power was maintained at 2 Watts.
The energy dose (E) was calculated by the following formula: E = Power
(W) x duty cycle x time (s), where duty cycle (DC) = pulses per second
(Hz) x length of pulse (s) x 100%. Continuous wave mode corresponds to
a duty cycle of 100%. Surgically removed unerupted human third molar teeth
were sectioned using an Isomet diamond saw. Sections were cut either horizontally
(crown segments) or vertically (root segments) to achieve a consistent
thickness (crown segments 2.0 mm, root segments 2.5 mm), measured from
the tooth surface to the pulp chamber. Sectioning was performed so that
the thermocouple could be placed in the pulpal site at greatest risk of
thermal insult. The pulp contents were removed, and the tooth sections
mounted in plaster discs. Specimens were kept moist at all times in order
to simulate intra-oral conditions of humidity. The method used to assess
temperature change at the dental pulp has been described in detail previously
(Sandford and Walsh, 1994). In brief, a K-type bead thermocouple was located
securely against the pulpal tooth surface by means of non-heat-conductive
plaster weights. The thermocouple was connected in turn to a temperature-to-voltage
converter, which provided an output of 1 mV per degree Celsius. The voltage
output was measured with a digital multimeter which provided a digitally
encoded output to the serial interface of an IBM-compatible computer. Data
for temperature were recorded at intervals of five seconds. After establishing
the baseline temperature, specimens were lased within the fissure region
or coronal root surface, directly opposite the thermocouple bead. The laser
was held at a constant distance of 1.5 mm from the target site. Recordings
were made continuously during and after lasing, until the samples had cooled
to baseline temperature. Each trial was replicated a minimum of five times.
The following parameters were evaluated: (i) temperature rise (difference
between baseline temperature, and the maximum temperature recorded during
the trial); (ii) time to maximum temperature (calculated from the start
of lasing); and (iii) time to cool to baseline (calculated from the timepoint
at which the maximum temperature occurred). Differences were evaluated
using two-way paired analysis of variance.
Temperature rise - Crown
As shown in Figure 1, the temperature following lasing of the crown
increased with increasing laser exposure. However, temperature rises for
all but one of the settings (60 seconds/5% DC/6.0 J) were below the “pulpal
injury” threshold of 2.2°C. Differences between the 1% and 5% duty
cycles were significant for irradiation times of 5, 10, and 30 seconds
(P<0.05), but not for 60 seconds.
- Temperature rise following irradiation of the crown
region. Bars show mean values (N = 5 replicates).
Temperature rise - Root
As shown in Figure 2, temperature rises with the dental pulp for laser
irradiation of root surfaces increased in proportion to the laser exposure.
Again, as with the crown series, temperature rises for all but one of the
settings (60 seconds/5% DC/6.0 J) were below the threshold value of 2.2°C.
Differences between the 1% and 5% duty cycles were significant for lasing
times of 10, 30, and 60 seconds (P<0.01, P<0.01, and P<0.05 respectively).
Values for trials of 5 seconds laser exposure were not significantly different.
There was a linear relationship between laser exposure and pulp temperature
rise for each of the laser parameters investigated in the study. Figure
3 shows linear regression analysis, with values for gradients (m) and correlation
- Temperature rise following irradiation of the root surface. Bars show
mean values (N = 5 replicates).
Time to reach maximum temperature
The times taken to reach maximum temperature were related to the duration
of laser exposure, with the maximum temperature occurring within 10 seconds
of completion of lasing in all instances. Data for these times are presented
in Table 1. Each value is the mean of five
readings, rounded to the nearest second.
- Linear regression analysis of the relationship between laser exposure
(Joules) and temperature rise.
Time to cool to baseline
Cooling times were related to the duration of laser exposure, with longer
lasing periods associated with longer cooling times. Data for times are
presented in Table 2.
This study examined thermal changes in teeth following exposure to CO2
lasers operated in pulsed mode. The results indicate that the combination
of low power and low duty cycle induces minimal thermal changes at the
level of the dental pulp. Temperature rises for the 1% duty cycle were
less than those for the 5% duty cycle, and this reflects the increased
opportunity for cooling to occur between laser pulses. The 5% duty cycle
comprises ten 5 millisecond exposures per second, with 950 milliseconds
per second for cooling (95 milliseconds after each laser pulse). In contrast,
the 1% duty cycle setting comprises two 5 millisecond exposures per second,
allowing 990 milliseconds per second for cooling (495 seconds after each
laser pulse). It should be noted that the two instances in which differences
between these two modes were not significant reflect proportionally high
standard deviations in the sample groups. Significant differences may be
obtained with larger sample sizes. With pulsed modes of lasing, the maximum
temperature is always reached at or immediately after the end of lasing.
This may be attributed to the thermal conductivity of the dentin and enamel/cementum,
in that the thermal pulse dissipated within the tooth takes several seconds
to reach the thermocouple bead. Times to maximum temperature may be even
longer in the mouth than those recorded in this study, due to cooling from
pulpal blood flow and saliva. These same two factors would, similarly,
be expected to reduce cooling times. This study demonstrates that CO2
lasers, even at low duty cycles, invariably induce a temperature rise within
dental pulp during lasing in vitro of the crown or root surface at settings
used for desensitization (Forrest-Winchester and Walsh, 1992) or enhanced
fluoride treatment (Walsh, 1993). Only the highest of the energy settings
used in the study (6.0 J) produced temperature changes (both in crown and
root) which were greater than the 2.2°C pulpal damage threshold defined
by Zach and Cohen (1965). An irradiance of 6.0 J is clinically unrealistic
for these hard-tissue procedures. Thermal changes at low energy settings
are unlikely to be significant at the clinical level, as pulp symptomatology
and other adverse changes have not been observed (Walsh, 1994). Thus, pulsed
lasing modes are recommended strongly for dental hard-tissue procedures
because they provide the opportunity for cooling between exposure pulses.
This research was supported by the Australian Society of Endodontology.
The author thanks M. Sandford for technical assistance.
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Correspondence: Dr. L.J. Walsh, University of Queensland Dental
School, Turbot Street, Brisbane, QLD 4000, Australia.
Accepted October 31, 1995
Electronic publication: October, 1996
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