Heat Generated During Seating of Dental Implant Fixtures

Irrigation to address heat generated by the dental implant osteotomy drill is a highly controversial topic. Many clinicians do not use sterile water or saline irrigation to cool the implant drill with no detrimental outcomes.¹  However, there may be another source of heat generation that may have been overlooked by clinicians, that is, frictional heat generated during implant fixture seating into the osteotomy.

The implant osteotomy is generally 0.5 mm smaller in diameter than the definitive implant that is placed in the osteotomy. This undersizing is done to ensure that the implant is mechanically secure in the bone to prevent any micro movement that may disturb or prevent osseointegration during healing.²  During placement of the implant fixture, there may be heat generated due to the friction of the implant fixture against the fresh-cut bone. The magnitude and time duration for causation of any detrimental osteotomy heat are uncertain. Most heat-generation studies on this were done on osteotomies and in vitro or on lower animals in nondental implant circumstances.³ 

Bone has a thermal conductivity between low-conducting subcutaneous fat and high-conducting muscle.⁴  This relatively low-to-moderate thermal conductivity property may cause the bone to retain generated heat and induce osseous necrosis.⁴  This also causes bone to not readily absorb heat.

The purpose of this study was to determine what heat magnitude may be generated during dental implant fixture seating into the osteotomy.

The null hypothesis was that there is no heat produced during dental implant placement into osteotomies.

Clean bovine mandibular rami were secured from a local university pathobiology service (University of Connecticut, Storrs, Conn). The bone was sectioned to provide osseous segments for study convenience (Figure 1). The bone sections were stored in saline in a refrigerator at approximately 5°C prior to study. During the study testing, each segment was maintained in a saline bath at a range of 34°C–38°C. All sites were deemed to be dense type 1 (Misch) bone 8- to 10-mm thick. All implants were obtained from the author's private practice. All implants used were retrieved, clinically failed implants or rejected due to intraoperative sizing issues. The author places 5–600 implants annually, and the implants were collected over several years. All implants were acid-etch blasted-surfaced 3.7, 4.7, 5.7 × 10 mm (Legacy, Implant Direct Implants, Ventura, Calif) or 2.5 × 10 mm (IntraLock, Boca Raton, Fla).

The workflow consisted of the following:

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  Secure clean bovine ramus and prepare thickness

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  Select and drill the osteotomies for the selected implants 2.5, 3.7, 4.7, 5.7

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  Drill the thermoprobe hole for placement 0.5 mm from osteotomy and 1 mm deep

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  Place bone with thermoprobe and implant in water bath at 36°C and allow to equalize and stabilize

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  Initiate implant placement torqueing to 50 Ncm at 12 rpm noting starting temperature and peak temperature

• •

  Evaluate data

All sites were drilled in the usual sequence as recommended by the manufacturer to the desired test diameter that matched the test implant diameter. The sites were first drilled with the proprietary pilot drill and then the succeeding incremental drills for that particular final implant size. The sequence of the 2.5 implant was a 1.5-mm diameter drill and finally the 2.0-mm drill. For the 3.7 implant, the drill sequence was 2.0 and then 2.8 mm. The drilling sequence for the 4.7 implant was 2.0, 2.8, and finally 3.4 mm. For the 5.7 implant, the sequence was 2.0, 2.8, 3.4, and finally 5.1 mm (Table 1; Figure 1).

The final osteotomy drill was 2.0 mm for the 2.5-mm-diameter implant, 2.8 mm for the 3.7-mm, 3.4 mm for the 4.7-mm, and 5.1 for the 5.7-mm implant. A small hole was drilled with a No. 1/2 round latch burr approximately 0.5 mm away from each osteotomy rim and 1 mm deep to accept a thermal probe for the electronic thermometer thermocouple (Omega Engineering, Stamford, Conn; Figures 2 and 3). Each bone segment, implant, and implant drill was immersed in the water bath and allowed to accommodate to the water temperature to a temperature approximating 37°C. During testing, each bone temperature was recorded as a start temperature when the bone temperature was stable in the water bath. Each implant was then placed in each site and rotated in a handpiece torque driver until seated into each osteotomy at 12 rpm to 50 newton centimeters (Ncm). The insertion torque of 45–50 Ncm is commonly used for dental implants. In most sites, complete seating to the fixture platform did not occur because of the implant's binding against the dense bone. The start and peak temperature increase were viewed on the thermocouple readout and recorded (Tables 2–4).

There was indeed a temperature rise during most of the implant seatings (Tables 2–4). The average temperature increases were 0.075°C for the 5.7-mm-diameter, 0.97°C for the 4.7-mm-diameter, 1.4°C for the 3.7-mm-diameter, and 8.6°C for the 2.5-mm-diameter implants. The temperature increase ranges were 0°C–0.2°C for the 5.7-mm diameter, 0.4°C–2.8°C for the 4.7-mm-diameter, 0.4°C–4.8°C for the 3.7-mm-diameter, and 1.8°c–18.0°C for the 2.5-mm-diameter implants. The temperature rise peak duration time in all tests was only momentary. All peak time durations were deemed to be less than 1 second and were not measured. Temperature rise and fall lasted approximately 2–5 seconds, but this was not measured. The temperature rises started a second or two after the implant seating had begun.

The data points of temperature differences revealed very different standard deviations, so a t test may be inappropriate and render incorrect analysis (Table 5). Log was used because the standard deviation was too different among the groups (one was large and one was small) to ensure a reliable analysis. The data points have been transformed to a log scale after adding 0.5 to the differences to avoid the logarithm of 0 (Figure 4). The spread of the data points then proved similar in spread. Analysis of variance (ANOVA) is the appropriate analysis to find differences among the implant diameter groups. The data were adjusted by the Tukey method for multiple testing, which increases the validity of the analysis. The result was that the 3.7- and 4.7-mm-diameter implant effects were very similar. The 2.5-mm and 5.7-mm results were very different from the 3.7-mm and 4.7-mm implants. The F value (between-group variability/within-group variability) as determined by ANOVA was 69.22 (P < .0001), showing significant differences between the groups considered as a whole. The P value of the log data points is significant in that it is less than the accepted .05 significance level.

A formula for frictional heat generation by one body moving in contact with another is as follows: q = nFv, where q = heat generation, n = a coefficient of friction based on the type of materials (eg, aluminum on steel is 0.61, metal on wet wood is 0.2), F = force on the body moving, and v = the speed of the motion.⁵  The metal–wet wood coefficient may be the closest approximation to titanium–vital bone.

According to this formula, as the force on the interface increases, the heat liberated increases. In addition, as the speed of contact increases, the heat liberation increases. The formula defines that if the implant moving against bone moves faster or if there is more force applied to the moving implant, there will be more heat generated. The friction is converted into thermal energy.

The heat generated is directly related to the frictional gradient of the two materials, bone and titanium. The titanium implants conduct and absorb thermal energy faster than vital or nonvital bone and thus will absorb any frictional heat generated at the bone-implant interface. This heat will then be transferred into the shank of the placement-driving instrument and into the handpiece. The heat transfer occurs where there is actual physical contact at the bone-interface contact and at areas of no contact via intervening tissue fluid. Thermal equilibrium will occur, but the conduction of the titanium implant is much greater than bone and will draw heat away from the bone-implant interface. The implant drill is indeed hot to touch when removed from the osteotomy, as clinicians will attest. However, this property of absorbing heat prevents any appreciable heat to be transferred and absorbed by the bone. Metals have free-moving electrons that enable rapid transfer of thermal energy. The covalent chemistry of organic polymers does not have this property. Cortical bone has a conductivity of approximately 0.16–0.34 watts of heat per meter per degrees Kelvin (W/m/K) and spongy bone 0.30 W/m/K,⁶  whereas titanium is 18.7–22.1 W/m/K⁵  or approximately 70 times the thermal conductivity of bone. Thus, titanium would be expected to absorb and conduct heat away from a bone-titanium interface.

Fourier's heat conduction equation describes the transfer of heat: q = −kA(dT/dx), where q is the rate of heat transfer, k is the constant of the conductivity of the material expressed as a negative, A is the contact area, and dT/dx is the temperature gradient, where dT is the temperature difference, that is, the temperature applied versus the temperature of the entity, and dx is the distance into the entity.⁷  Heat conduction will progress from a body of higher temperature to a body of lower temperature. The larger the difference in temperature, the faster the heat will transfer. The higher the conductivity, the faster the heat transfer. The negative describes the direction of the heat transfer to or away from the material of higher thermal conductivity. Thus, heat generated at the titanium-bone interface would be transferred away from the bone and into the titanium metal and further into the placement attachment. In addition, the area of contact will enable faster transfer of heat.⁷ 

As the surface area of contact increases, the heat transfer increases. The heat of the force is then dissipated over the larger surface area, and the body with the greater capacity for heat absorption then indeed does absorb the heat. This is the heat sink effect. The force and seating torque on all the implants was the same, 50 Ncm at 12 rpm. The 2.5-mm-diameter implants generated the most in bone insertion heat, have the smallest surface area, and would not transfer as much heat as a larger implant.

For the sake of simplicity in these calculations, the implants are assumed to be cylinders. The formula for the surface area of a cylinder excluding the coronal and apical surfaces is as follows: surface area = h × 2 × pi × r, where h is the length of the implant, pi is the constant 3.14, and r is the radius of the implant (Table 5). The implant ends do not contact bone, so these surfaces are not considered. The surface area of the 3.7-mm-diameter, 10-mm-length implant has an approximate surface area that interfaces with the bone of 10 × 2 × 3.14 × 3.7/2 = 10.618 mm². The approximate surface area of the 5.7-mm-diameter implant that interfaces with the bone is 10 × 2 × 3.14 × 5.7/3 = 17.898 mm². Thus, the surface area of the larger 5.7-mm implant is 170% larger than the smaller 3.7-mm-diameter implant. The surface of the 2.5 implant is 2.28 times smaller than the 5.7 implant, and the 5.7 implant produced less than 1% of the in-bone heat increase. The differences between the average temperature rises of the implants confirm the predicted outcome of Fourier's formula, where the area of the surface interface is a major influence on frictional heat transfer.

The differences between the average temperature rises of the 5.7, 4.7, and 3.7 implants was less than 1.5°C, but there was a much larger 7.2°C difference between the 3.7-mm-diameter implants and the 2.5-mm-diameter implants This large difference may be explained by the area and volume differences among the implants, which impart heat absorption properties to the physical sizes of the implants. The volume of a cylinder is pi × radius squared × height. The large 5.7-mm-diameter implants had a very small to no temperature rise. This may be due to the heat sink effect of the larger surface area. In addition, the volume of titanium would draw any heat generated into the much more thermoconductive metal and then into the placement attachment on the drill handpiece. The contact of the mineral portion of the cortical bone to the titanium oxide surface may be the major cause of the friction and thus the generated heat.⁸  The organic portion is unlikely to generate frictional heat. Heat is generated by friction of the implant against the bone and compression of the compressible bone during seating.⁹  The implant is generally slightly larger than the osteotomy according to contemporary placement techniques and compresses the bone away from the advancing implant.

The heat generated was apparently not enough to make a significant difference as measured by the thermoprobe, since the titanium has 70 times greater potential for heat transmittal than bone.

The thermoprobe was located approximately 0.5 mm from the implant-bone interface. This placement distance has been used in past work.¹⁰  The temperature at the bone-implant juncture was likely slightly higher than the temperature measured by the thermoprobe. Since bone has a relatively low conductivity, the heat at the implant-bone interface may be higher than the heat detected at 0.5 mm from that interface.

It is this author's experience that most clinical implant seating placements are in bone that is less dense than type 1 (Misch) bone and may generate even less heat than this study found. Nevertheless, one recent study found temperature increases in nonvital porcine rib bone when placing 4.0-mm implants.¹¹  A total of 288 self-tapping implants were placed in osteotomies to 30, 35, and 40 Ncm. Thermocouples monitored the bone at 5-, 1.5-, and 10-mm depths. The temperature increases were found to be significant at the cortical level but not so at deeper trabecular levels of bone (5- to 10-mm deep).¹¹  Lower insertion torques used did not cause a significant temperature increase. The thermoprobes were placed 5-mm away from the osteotomies, which may not reveal the temperature increases that may occur in closer proximity. Because the greatest friction may occur at the cortical bone at the osteotomy opening, the deep placed thermoprobes did not demonstrate any significant temperature increases.

The friction that occurs may be mostly at the implant cervical–cortical bone interface in all densities of bone (Figure 5). Thus, the temperature rise may be approximately the same in dense or less-dense bone sites if the cortical bone thickness is equal.¹¹  Thus, most of the heat would be generated at the ridge crest cortical bone, but the transfer of this heat to bone is dependent on the ability of the implant to absorb that heat. The larger the displacement and thermal conductivity, the less heat will be transferred to the bone.

Higher insertion torques can produce higher temperature rises in bone.¹¹  Nevertheless, a larger displacement of the implant being seated may transmit heat away from the contacting bone. Small 2.5-mm-diameter implants may thus generate the same heat per square millimeter of surface area during insertion, but the smaller surface area and displacement of the implant may not remove as much generated heat that would then be transmitted to the bone. This may explain informal clinical reports of implant failures in the anterior of the mandible where implants were placed in very dense bone.

Many manufacturers make osteotomy drills that are approximately 0.5 mm smaller in diameter than the diameter of the implant placed in that site. This difference provides initial stability to enable osseointegration.⁸  There is compression of the bone during seating of the tapered implant and friction of the rough implant surface rotating against the freshly cut osseous surface. Blood and tissue exudate may act as lubricants to lessen any heat generation. In addition, the implant-bone surfaces may engage significantly only at the peaks of the implant threads, minimizing the actual contact area. The thread pitch may contribute to the speed of seating and thermal increases but probably not significantly. Osteotomies are generally cylindrical or tapered to match the shape of the selected implant. Thus, the implant as it seats has most contact with the bone at the osteotomy opening, that is, the cortical bone (Figure 4). This minimal contact may maximize the heat sink effect of the metal implant/drill handpiece. A cylindrical implant/osteotomy may generate less heat on seating than a tapered implant because of probable less surface contact and contact force.

The nonvital bovine ramus bone is much denser than vital human bone and was used for its ready availability. The unyielding nature of the bovine ramus prevented complete seating of most of the implants, and no attempt to ensure complete seating was done to ensure consistency of the trials. This is an in vitro study and makes no pretense as to its importance or clinical relevance. Obviously, seating temperatures will be different in vital bone. Further study is needed to clarify this issue. The use of bone taps is clinically done, but taps were not used to keep the trials consistent. Taps may indeed alter the generation of heat. A subsequent study may be done to ascertain this. This is beyond the scope of this in vitro study.

There has been some discussion in the dental implant literature about osseous thermal damage during dental implant surgery.^12–14  There is some debate as to at what temperature is vital bone irreversibly damaged.¹  The temperature at which there is thermal damage to vital bone may yet be an undetermined parameter for any individual patient for a given osseous site. A study by Eriksson et al¹⁴  found that bone necrosis can occur at 47°C that is held for 1 minute. Nevertheless, there may be different temperatures at which osseous cell damage does or does not occur depending on the bone density, blood supply, and the instrumentation that is used that causes the temperature rise. Except for the 2.5-mm-diameter implants, the in vitro results herein had a temperature rise below the reported destructive temperature found by Eriksson et al.¹⁴  It must be stressed that the present tests were done in dense nonvital bovine bone with no vascularity.

Vital osseous heating was studied by Eriksson et al.¹⁴  In their work, vital rabbit bone was heated to 53°C for 1 minute. Bone necrosis occurred but was reparable, and at some time during the 3–5 postoperative week, regeneration did take place. In addition, a temperature rise of 4.3°C may reduce the quality of subsequently repaired bone.^13,14  Blood perfusion may be the important parameter here. The authors reported that perfusion increased with the temperature increase but then stopped when 53°C was reached. Thus, this debate may be an issue of the importance of blood supply, bone physiology, and density. The bone formation enzyme, alkaline phosphatase, denatures at 53°C and thus would be inactivated at this temperature.¹²  Bone healing may be impaired. After removal of the thermal source, repair and regeneration may occur. In yet another Eriksson study, heating vital bone to 44°C to 47°C for 2 minutes produced bone necrosis.^13,15  This necrosis was irreparable. The temperature increases found in the present in vitro study herein were below previously reported detrimental temperatures and may not be detrimental to vital bone. The increase with 2.5-mm implants approached or exceeded the necrosis temperature but only momentarily, and again, the test bone is nonvital and may generate higher temperature increases than may be seen in vital bone.

Another study compared the heat generated by self-drilling screws versus self-tapping screws in rabbit calvaria.¹⁶  The screws were 0.2 or 0.4 mm larger in diameter than the prepared osteotomies. The screws were 1.5 or 2.0 mm in diameter and 5-mm long. The authors found that bone damage did occur during insertion of nonirrigated self-tapping screws, but this may not be clinically significant. The osteotomy itself will induce bone cell damage.¹⁷  The dental implants used in the present study were self-tapping screw-type implants of much larger diameter and length and were installed in a water bath kept at or near body temperature, which may simulate clinical conditions more closely. The water bath and the much larger implant sizes used in the present study herein may dissipate any heat and prevent a larger rise in temperature. Blood and tissue exudate may act as lubricants to minimize friction in vital bone.

Time duration may be an issue in that irreparable bone necrosis may occur after 2 minutes of temperature elevation.^14,15  The seating heat increases seen in these tests were momentary and would probably not incur serious osseous necrosis in vital bone. Nevertheless, if the generated heat is not conducted away from the implant bone interface by a large displacement or surface area, the generated heat may induce necrosis.

The point velocity of the larger-diameter implant is faster than the point velocity of the smaller-diameter implant since all insertions were done at 12 rpm. The 5.7-mm-diameter implant rotated at a velocity of 68.4 mm per minute, while the 3.7-mm-diameter implant rotated at 14.4 mm per minute. According to Fourier's formula, the higher point velocity of the larger implant would generate more heat than the smaller-diameter implant. Even though more heat may be generated, this apparently does not outweigh the ability of the larger implant to remove heat from the interface. To minimize the heat of surface movement velocity, a very slow placement, less than 12 rpm, of the small implant may be indicated, with the implant cooled before insertion and/or the osteotomy site cooled. Leaving the placement instrument connected to the small implant may dissipate any generated heat.

Vital osseous sites may not absorb or accept generated heat as readily as nonvital bone, and this may make any temperature increase clinically insignificant.¹  That is, vital bone with a blood supply and tissue fluid and blood immersing the osteotomy may not incur a significant frictional seating temperature increase. In previously published work by this author, no temperature increase was seen when performing maxillary osteotomies in two living patients without irrigation.¹ 

The implant seating heat generated in this in vitro testing in dense type 1 (Misch) bovine bone was generally found to be below the purported detrimental temperature and duration magnitudes. A large implant surface area and displacement produced less of a temperature rise as the implants were seated. Titanium may be 70 times more heat conductive than bone. Larger-diameter titanium implants may act as a heat sink that draws away any generated heat from the bone-implant interface. Thus, a larger implant, with larger surface area and displacement, may have a capacity to absorb more heat, more readily, than smaller-diameter implants. Except for the 2.5-mm implants, the temperature increases found in this study are well below the purported detrimental temperature level and time duration in larger-diameter implants. The nonvital bovine ramus bone used in this testing does not mimic vital conditions. Vital bone may not allow dramatic temperature increases because of blood supply, blood itself, and tissue exudate that may act as lubricants. Frictional heat generated during implant fixture seating may be considered clinically insignificant in most vital bone sites because of the ability of the metal implant to draw heat away from the bone. Implant surface area and thermal conductivity are important factors in generated heat transfer at the bone-implant interface. Smaller-diameter implants may not remove generated heat as compared with larger implants so that generated heat is transmitted to the surrounding bone. This heat may be detrimental in some sites and may cause or contribute to a short-term failure to integrate, called an early failure. Nevertheless, further investigation is needed to discover any truly significant issues related to implant seating in vital bone osteotomy sites.

Placing and seating implants in prepared osteotomies does generate heat at the bone-implant interface. The implant surface area and the higher thermal conductivity of the titanium implant draw the generated heat away from the bone. This may prevent a detrimental increase in temperature. The null hypothesis is disproved. This in vitro animal bone study may not be applied to the clinical situation, and clinical trials need to be conducted to verify these results in vital human bone. Nevertheless, caution should be shown when placing small-diameter implants. A very low insertion rotation rate may be observed to minimize any heat generation, especially in dense bone with low vascularity, such as the in anterior mandible. In addition, leaving the handpiece in contact with the implant for an extended period after seating may conduct heat away from the implant-bone interface.

The author acknowledges the technical support of Mss. Erika Paegle, CDA, Alicia Gosselin, CDA, Britani Ames, CDA, Danielle Green, DMD, and Martie Flanagan, BSRN.