U.S. patent application number 13/871665 was filed with the patent office on 2016-12-08 for apparatus and method for assessing percutaneous implant integrity.
This patent application is currently assigned to COVENANT HEALTH. The applicant listed for this patent is Gary FAULKNER, Donald Wayne RABOUD, Ryan Clair SWAIN, Johan Francis WOLFAARDT. Invention is credited to Gary FAULKNER, Donald Wayne RABOUD, Ryan Clair SWAIN, Johan Francis WOLFAARDT.
Application Number | 20160356745 13/871665 |
Document ID | / |
Family ID | 39081873 |
Filed Date | 2016-12-08 |
United States Patent
Application |
20160356745 |
Kind Code |
A9 |
FAULKNER; Gary ; et
al. |
December 8, 2016 |
APPARATUS AND METHOD FOR ASSESSING PERCUTANEOUS IMPLANT
INTEGRITY
Abstract
Provided is an apparatus for assessing interface integrity
between a medium and an implant. A first signal is translated from
a motion of an impact body during impact with an abutment connected
to the implant. In some embodiments, the first signal is filtered
using a zero phase shift filter and then used for assessing the
interface integrity. Since no phase shift is introduced, the
interface integrity is accurately assessed. In another embodiment,
the apparatus maintains a system model for impacting the impact
body against the abutment. The apparatus analytically determines an
interface property by applying a system property that has been
determined to the system model. An accurate system model allows for
an accurate assessment. According to another broad aspect, there is
provided a method of conducting the impact test. According to the
method, a person ensures that the impact body impacts against a
consistent portion of the abutment.
Inventors: |
FAULKNER; Gary; (Edmonton,
CA) ; RABOUD; Donald Wayne; (Edmonton, CA) ;
SWAIN; Ryan Clair; (Edmonton, CA) ; WOLFAARDT; Johan
Francis; (Edmonton, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
FAULKNER; Gary
RABOUD; Donald Wayne
SWAIN; Ryan Clair
WOLFAARDT; Johan Francis |
Edmonton
Edmonton
Edmonton
Edmonton |
|
CA
CA
CA
CA |
|
|
Assignee: |
COVENANT HEALTH
Edmonton
CA
|
Prior
Publication: |
|
Document Identifier |
Publication Date |
|
US 20140318247 A1 |
October 30, 2014 |
|
|
Family ID: |
39081873 |
Appl. No.: |
13/871665 |
Filed: |
April 26, 2013 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
12377768 |
Jan 13, 2011 |
8448516 |
|
|
PCT/CA07/01416 |
Aug 17, 2007 |
|
|
|
13871665 |
|
|
|
|
60822686 |
Aug 17, 2006 |
|
|
|
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
G01N 29/4472 20130101;
A61C 8/00 20130101; G01N 29/045 20130101; A61B 5/0051 20130101;
A61C 19/04 20130101; A61B 5/4504 20130101; A61B 5/682 20130101;
G01N 29/12 20130101; G01N 29/4418 20130101; A61B 5/11 20130101;
A61B 9/00 20130101; G01N 2291/0258 20130101 |
International
Class: |
G01N 29/12 20060101
G01N029/12 |
Claims
1. An apparatus for processing a signal for determining an
indication of an interface integrity between a medium and an
implant that is at least partially embedded therein, the apparatus
comprising: an input for receiving a first signal generated from a
motion of an impact body during impact with an abutment connected
to the implant; and a zero phase shift filter for filtering the
first signal thereby generating a filtered signal to be used for
determining the indication of the interface integrity.
2. The apparatus of claim 1, further comprising: an output for
providing the filtered signal to another entity that determines the
indication of the interface integrity based on the filtered
signal.
3. The apparatus of claim 1, further comprising: a property
determiner for determining the indication of the interface
integrity based on the filtered signal.
4. The apparatus of claim 3, wherein the indication of the
interface integrity determined by the property determiner is an
explicit indication of the integrity of the interface.
5. The apparatus of claim 3, wherein the property determiner is
operable to determine the indication of the interface integrity by:
determining a natural frequency of a system comprising the implant,
the abutment, and the impact body; and determining the indication
of the interface integrity based on the natural frequency.
6. The apparatus of claim 3, further comprising: a signal processor
implementing both the zero phase shift filter and the property
determiner.
7. The apparatus of claim 3, wherein the zero phase shift filter is
a moving average filter.
8. The apparatus of claim 3, further comprising: the impact body;
and a motion detector connected to the impact body for translating
the motion of the impact body during impact into the first
signal.
9. A method of processing a signal for determining an indication of
an interface integrity between a medium and an implant that is at
least partially embedded therein, the method comprising: receiving
a first signal generated from a motion of an impact body during
impact with an abutment connected to the implant; and filtering the
first signal using a zero phase shift filter thereby generating a
filtered signal to be used for determining the indication of the
interface integrity.
10. The method of claim 9, further comprising: providing the
filtered signal to another entity that determines the indication of
the interface integrity based on the filtered signal.
11. The method of claim 9, further comprising: determining the
indication of the interface integrity based on the filtered
signal.
12. The method of claim 11, wherein the indication of the interface
integrity is an explicit indication of integrity.
13. The method of claim 11, wherein determining the indication of
the interface integrity based on the filtered signal comprises:
determining a natural frequency of a system comprising the implant
and the abutment; and determining the indication of the interface
integrity based on the natural frequency.
14. The method of claim 11, wherein the zero phase shift filter is
a moving average filter.
15. The method of claim 11, further comprising the step of
translating the motion of the impact body during impact into the
first signal.
16. A non-transitory, computer readable medium having computer
executable instructions stored thereon for execution on a processor
to implement the method of claim 9.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a divisional of U.S. patent application
Ser. No. 12/377,768, which is the U.S. national phase application
of International Application PCT/CA07/01416, filed on Aug. 17,
2007, which claimed priority from U.S. Provisional Patent
Application 60/822,686 filed on Aug. 17, 2006, the contents of each
which is incorporated herein by reference in their entirety.
FIELD OF THE INVENTION
[0002] This invention relates to techniques for assessing
percutaneous implant integrity, and in particular to careful
control of conditions under which impact-style techniques are
utilised, and the analysis of data received therefrom.
BACKGROUND
[0003] Osseointegrated implants are routinely utilized in a broad
range of oral and extraoral applications including removable and
fixed dental prostheses, in re-construction of the head and neck,
as a transmission path for bone anchored hearing aids (BAHA.TM.),
to provide anchorage in orthodontic treatment and in orthopedic
applications. FIG. 1 shows a cross-sectional side view of a typical
in-situ implant and abutment system. Such implants are typically
3-6 mm in diameter and range in length from 3-4 mm (BAHA and orbit
applications) to 7-20 mm (dental reconstructions). Such implants
are often formed of titanium.
[0004] The success of these implants is dependent on the quality of
the bone-implant bond at the interface of the implant. A direct
structural and functional connection between living bone and the
surface of a load-carrying implant is defined as osseointegration.
This process typically begins immediately after the implant has
been installed. If this does not occur, the development of
connective soft tissue in the bone-implant interface may begin and
can lead to failure of the implant. The status of the implant-bone
interface during this crucial time is extremely important in
evaluating when the implant can be put into service (loaded) or
whether further healing is necessary.
[0005] In addition, over time osseointegration can deteriorate
and/or the degree of bone in contact with the implant surface can
reduce. Although implant survival rates are high in many
applications, it is important to be able to determine if any change
in the health of this interface occurs. As a result of these
potential clinical conditions, there is an ongoing desire to
monitor the "health" or integrity of the bone-implant interface
from initial installation of the implant throughout the life of the
implant.
[0006] Conventional diagnostic techniques, such as radiography and
magnetic resonance imaging, are generally able to evaluate bone
quantity and in some cases may provide parameters that relate to
bone quality (eg. Hounsfield radiodensity scale). However, these
techniques are limited in their ability to monitor the actual
bone-implant interface, as the implant tends to shield this region
resulting in poor image resolution in this vital area. Therefore,
the condition of the bone-implant interface including the implant
threads and the adjacent tissue undergoing remodelling is much more
difficult to evaluate. When using radiography, the changes in bone
are often well advanced before becoming evident on radiographic
images. Furthermore, images obtained in this manner are costly and
high quality radiographs carry additional risks associated with
radiation exposure.
[0007] Other techniques such as measuring removal torque are too
invasive to be used in either the operating room or for clinical
visits. As a result, dynamic mechanical testing methods have been
proposed and are presently in use. These mechanical techniques are
all, in one form or another, based on determining the resonant
frequency of the implant-tissue system including the transducer. As
the resonant frequency is dependent on the manner in which the
implant is supported by the surrounding biological tissue, changes
in this resonant frequency (perhaps coupled with changes in the
internal damping) should be linked to changes in the status of this
interface. This, of course, assumes that there are no other changes
in the implant system (such as a loosening of the abutment/implant
joint) that may overshadow those in the interface.
[0008] Presently, the primary commercially available system
developed specifically for monitoring implants is Osstell.TM.,
which employs a transducer attached to the abutment or directly to
the implant. The transducer excites the system over a range of
frequencies while simultaneously monitoring the resulting
transducer motion to determine the resonant frequency of the
overall implant/transducer or implant/abutment/transducer system.
The results of several investigations using this system have
reported varying degrees of success in identifying changes in the
implant status. A disadvantage of the Osstell is that it is
designed to be used with retrievable systems only.
[0009] Alternative techniques to the Osstell are based on transient
measurements in which the abutment is excited using an external
impact. Subsequently, a measurement method was developed that
utilised an instrumented impact hammer to evaluate the mechanical
impedance variations caused by interface changes. One approach
involved an impacting rod to excite the abutment and the resulting
resonant frequency was determined from an acoustic signal obtained
from a microphone mounted in close proximity.
[0010] Another system that has been used is the Periotest.TM.,
which was originally developed to measure the mobility of natural
dentition. As shown in FIG. 2, there is a Periotest handpiece,
which contains a metal rod of approximately 9 grams. The metal rod
is accelerated towards the implant-abutment via an electromagnet.
The acceleration response of the rod, while in contact with the
implant-abutment, is measured using an accelerometer attached to
the rear of this rod. In particular, the acceleration signal is
used to determine the period of time during which the rod and tooth
remain in contact. This period of time is indicative of the
integrity of the tooth interface.
[0011] There are benefits to the Periotest system. The Periotest
handpiece provides a convenient means to dynamically excite the
implant abutment system in areas that may be too cramped to utilise
Osstell or impact hammer devices. Also, the Periotest handpiece can
be used on implant abutment systems with non-recoverable, cemented
restorations. As well, the output signal from the accelerometer may
contain information unavailable to the RFA systems, which can be
more completely utilised to determine the status of the interface
layer. For example, the handpiece has recently been adapted for use
in a system designed to measure the damping capacity of
materials.
[0012] Several researchers have attempted to adopt the Periotest in
monitoring the integrity of artificial implants instead of natural
teeth. The results of these investigations have shown varying
degrees of success. When used to monitor the mobility of natural
teeth, the contact time is not used directly but is used to
calculate a so-called Periotest value (PTV) which was originally
chosen to correspond to the established Miller Mobility Index for
natural teeth. For natural teeth, which are supported by
periodontal ligaments, the PTV's range is from approximately -8 to
50 with -8 representing a tooth with a very stiff supporting
structure and a PTV of 50 corresponds to a tooth which is
noticeably loose and moveable by finger pressure.
[0013] When used to measure artificial implants the contact times
involved correspond to PTV's that are significantly lower than for
natural teeth, as the bone to implant interface provides a much
stiffer supporting structure than periodontal ligaments. Since the
Periotest has a built in lower PTV limit of -8 and only produces
integer values, there is a limited range of PTV readings available
for a typical implant application. For example, it has been found
that well integrated implants have a range of PTV values between -7
and 0 in the mandible and -7 to +1 in the maxilla at the time of
abutment connection. This limited range does not provide enough
resolution to monitor subtle changes in the bone-implant interface
over time. The Periotest system cannot accurately determine a
contact time for very stiff implant interfaces, especially for
those that are extraoral.
BRIEF DESCRIPTION OF THE DRAWINGS
[0014] FIG. 1 is a cross-sectional side view of a typical in-situ
implant and abutment system;
[0015] FIG. 2 is a cross-sectional side view of a typical Periotest
system;
[0016] FIG. 3 is a schematic drawing of a Periotest rod striking an
implant abutment;
[0017] FIG. 4 is a graph showing typical raw and conditioned
Periotest signals;
[0018] FIG. 5 is a schematic of an apparatus for determining an
indication of an interface integrity between a medium and an
implant that is at least partially embedded therein;
[0019] FIG. 6 is a flowchart of a method of processing a signal for
determining an indication of an interface integrity between a
medium and an implant that is at least partially embedded
therein;
[0020] FIG. 7 is a schematic of another apparatus for determining
an indication of an interface integrity between the medium and the
implant that is at least partially embedded therein;
[0021] FIG. 8 is a flowchart of another method of processing a
signal for determining an indication of an interface integrity
between a medium and an implant that is at least partially embedded
therein;
[0022] FIG. 9 is a photograph of a testing apparatus for an in
vitro model;
[0023] FIG. 10 is a graph showing a Periotest conditioned signal
and a moving average filtered signal;
[0024] FIG. 11 is a graph depicting a comparison of contact times
calculated based on the moving average filtered signal and the
conditioned Periotest signal;
[0025] FIG. 12 is a graph depicting a Periotest conditioned signal,
a moving average filtered signal and a strain gauge signal;
[0026] FIG. 13 is a chart depicting the repeatability and
reproducibility of the experimentation;
[0027] FIG. 14 is a schematic of an apparatus for determining a
property of an interface between a medium and an implant that is at
least partially embedded therein;
[0028] FIG. 15 is a flowchart of a method of determining a property
of an interface between a medium and an implant that is at least
partially embedded therein;
[0029] FIG. 16 is a schematic of the in vitro experimental model
for impact testing;
[0030] FIG. 17 is a schematic of a finite element analysis (FEA)
model for the impact test;
[0031] FIG. 18 is a graph depicting a typical transient analysis
signal for the FEA model of FIG. 17;
[0032] FIGS. 19A through 19D are graphs depicting changes in first
and second natural frequencies of the implant abutment as a
function of increasing loss of osseointegration and bone margin
height;
[0033] FIGS. 20A through 20D are graphs depicting changes in first
and second natural frequencies of the implant abutment as a
function of increasing interface layer stiffness;
[0034] FIG. 21 is a schematic of a four-degree of freedom model for
the impact system;
[0035] FIGS. 22A and 22B are graphs comparing measured acceleration
response and predicted acceleration response;
[0036] FIG. 23 is a graph comparing measured acceleration response
with damped model acceleration response;
[0037] FIGS. 24 and 25 are graphs comparing measured acceleration
response with damped model acceleration response with and without
flange;
[0038] FIG. 26A to 26D are graphs depicting modal acceleration
components;
[0039] FIGS. 27A and 27B are graphs comparing model results with
measurements for abutment strike at different heights;
[0040] FIGS. 28A to 28D are graphs comparing model results with
measurements for abutment strike at different points along the
abutment;
[0041] FIGS. 29A and 29B are graphs comparing model results with
measurements for implants of different length;
[0042] FIGS. 30A to 30D are graphs comparing model results with
measurements for abutments of different length;
[0043] FIGS. 31A and 31B are graphs depicting effects of varying
support stiffness on the first mode frequency for two abutment
lengths;
[0044] FIG. 32 is a graph depicting effects of changing the damping
coefficient on the model acceleration response;
[0045] FIGS. 33A and 33B are graphs depicting effects of bone loss
from the top of the implant towards the base on the first mode
resonant frequency;
[0046] FIGS. 34A and 34B are graphs depicting model results with
and without a flange at two different first mode frequencies;
[0047] FIG. 35 is a flowchart of an example method of conducting an
impact test;
[0048] FIG. 36 is a chart depicting natural frequency as a function
of the distance of the Periotest handpiece from the abutment;
[0049] FIG. 37 is a chart depicting natural frequency as a function
of abutment torque;
[0050] FIG. 38 is a chart depicting natural frequency as a function
of striking height;
[0051] FIG. 39 is a chart depicting natural frequency as a function
of handpiece angulation;
[0052] FIG. 40 is a photograph of a calibration block used during
in vivo measurements;
[0053] FIGS. 41A and 42B are graphs comparing model to measurement
values with and without a flange;
[0054] FIG. 42 is a graph depicting percent difference in predicted
interface stiffness for two different abutment geometry
measurements with and without a flange for 10 patients;
[0055] FIGS. 43A and 43B are graphs comparing impact measurements
with predicted model response at the 12 month measurement with two
different abutment lengths and no flange support;
[0056] FIG. 44 is a graph depicting average longitudinal interface
stiffness based on all patients compared to individual interface
stiffness results;
[0057] FIGS. 45A to 45F are graphs comparing acceleration
measurement to predicted model response at different patient visits
for a patient; and
[0058] FIGS. 46A to 46F are graphs comparing acceleration
measurement to predicted model response at different patient visits
for another patient.
SUMMARY OF THE INVENTION
[0059] According to a broad aspect, there is provided an apparatus
for processing a signal for determining an indication of an
interface integrity between a medium and an implant that is at
least partially embedded therein, the apparatus comprising: an
input for receiving a first signal generated from a motion of an
impact body during impact with an abutment connected to the
implant; and a zero phase shift filter for filtering the first
signal thereby generating a filtered signal to be used for
determining the indication of the interface integrity.
[0060] According to another broad aspect, there is provided a
method of processing a signal for determining an indication of an
interface integrity between a medium and an implant that is at
least partially embedded therein, the method comprising: receiving
a first signal generated from a motion of an impact body during
impact with an abutment connected to the implant; and filtering the
first signal using a zero phase shift filter thereby generating a
filtered signal to be used for determining the indication of the
interface integrity.
[0061] According to another broad aspect, there is provided a
computer readable medium having computer executable instructions
stored thereon for execution on a processor so as to implement the
method summarised above.
[0062] According to another broad aspect, there is provided an
apparatus for determining a property of an interface between a
medium and an implant that is at least partially embedded therein,
the apparatus comprising: an input for receiving a signal generated
from a motion of an impact body during impact with an abutment
connected to the implant; and a property determiner for: (a)
maintaining a mathematical model for impacting the impact body
against the abutment; (b) determining a system property from the
signal; and (c) analytically determining the property of the
interface by applying the system property to the mathematical
model.
[0063] According to another broad aspect, there is provided a
method of determining a property of an interface between a medium
and an implant that is at least partially embedded therein, the
method comprising: maintaining a mathematical model for impacting
an impact body against an abutment connected to the implant;
receiving a signal generated from a motion of the impact body
during impact with the abutment; determining a system property
based on the signal; and analytically determining the property of
the interface by applying the system property to the mathematical
model.
[0064] According to another broad aspect, there is provided a
computer readable medium having computer executable instructions
stored thereon for execution on a processor so as to implement the
method summarised above.
[0065] According to another broad aspect, there is provided a
method of conducting an impact test to assess integrity of a
plurality of implants using an impact-type testing system, each
implant being at least partially embedded in a medium and having an
abutment connected thereto, the method comprises: impacting an
impact body against each abutment; and ensuring that the impact
body impacts against each abutment at a consistent portion of the
abutment.
[0066] According to another broad aspect, there is provided a
method of conducting impact tests to assess integrity of an implant
over time using an impact-type testing system, the implant being at
least partially embedded in a medium and having an abutment
connected thereto, the method comprises: from time to time,
conducting an impact test by impacting an impact body against the
abutment; and ensuring that the impact body impacts against the
abutment at a consistent portion of the abutment for each impact
test.
[0067] According to another broad aspect, there is provided a
calibration block comprising: a medium; and a plurality of systems,
each system comprising a respective implant embedded in the medium
and a respective abutment connected to the implant; wherein each
system has a predetermined nominal value for a system property.
DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS
[0068] While the methods of the present invention are described in
the context of an impact test conducted on an abutment attached to
an artificial implant, it is to be understood that these methods
may also be employed in the context of natural dentition. Thus, in
this specification, the term "abutment" includes the crown of a
natural tooth, while "implant" includes the root of a tooth. It is
also to be understood that the present invention is applicable to
replacement teeth. In such applications, the "implant" is synthetic
and might for example be formed of titanium. The "abutment"
connected to the implant is also synthetic and is typically
designed to function as a tooth crown.
Section I: Zero Phase Shift Filter
Introduction
[0069] Referring now to FIG. 3, shown is a schematic drawing of a
Periotest rod striking an implant abutment. At point A, the
Periotest rod strikes the implant abutment. The Periotest rod and
the implant abutment remain in contact through points B and C. At
point C the accelerometer signal reaches zero. The interval between
A and C is termed the contact time and is indicative of the
integrity of the implant interface. The Periotest system measures
the contact time in order to assess the integrity of the implant
interface, but as noted above, the resolution of the Periotest
system is limited.
[0070] The limited resolution of the Periotest system is further
compounded by the fact that the Periotest unit does not base the
contact time on the accelerometer signal directly. Instead, the
accelerometer signal is first conditioned using a filter to smooth
the signal. The contact time is then based on this filtered signal.
However, the filter used can produce a noticeable and significant
phase shift in the accelerometer signal, which introduces a
distortion of the contact time.
[0071] In the Periotest system, the accelerometer signal is
filtered and processed to yield a quantitative measure of mobility
related to the Miller Mobility Index for natural dentition. An
example of the signal before and after filtering is shown in FIG. 4
for the corresponding motion of the implant and rod shown in FIG.
3. There are two major differences between the conditioned and
unconditioned signals in FIG. 4. A comparison of the two signals
shows the unconditioned having a distinct higher frequency
component that has been removed in the conditioning. In addition,
the time for the acceleration to return to zero is considerably
longer for the conditioned than the unconditioned signal.
[0072] The differences between the conditioned and unconditioned
signals suggest that perhaps the filtering discards information
that could be used for a more complete diagnosis. Also, the
differences suggest that the filtering alters the fundamental time
(to return to zero acceleration) used to calculate the response in
terms of a so-called Periotest value (PTV). While these differences
may not be significant for natural dentition, as the range of PTV
values is relatively large (-6 to 50), they have more significance
for implant-abutment systems where the majority of results have
PTV's over a much more limited range (-8 to 2).
[0073] The possibility of using the impact technique of the
Periotest system to more precisely monitor the status of the bone
to implant interface has been investigated. An issue to consider is
whether variables such as osseointegration levels and loss of bone
margin height have an appreciable effect on the overall response.
To investigate these issues, the raw accelerometer signal such as
the one shown in FIG. 4 has been investigated.
System and Method
[0074] Turning now to FIG. 5, shown is a schematic of an apparatus
11 for determining an indication of an interface integrity between
a medium 20 and an implant 22 that is at least partially embedded
therein. An abutment 24 is connected to the implant 22. The
apparatus 11 has a signal processor 10 connected to an impact body
26 via a coupling 28. The impact body 26 has a motion detector 27,
which might for example be an accelerometer. The signal processor
10 has an input 12, a zero-phase shift filter 14, and a property
determiner 16. The apparatus 11 may have other components, but they
are not shown for sake of simplicity.
[0075] In operation, a user impacts the impact body 26 against the
abutment 24. The impact body 26 might be accelerated towards the
abutment 24 for example via an electromagnet. The motion detector
27 translates the motion of the impact body 26 during impact into a
first signal, which is provided to the signal processor 10 over the
coupling 28. The coupling 28 is a wired connection, but in
alternative implementations might be a wireless connection. The
signal processor 10 receives the first signal over the input 12.
According to an embodiment of the invention, the first signal is
filtered with the zero phase shift filter 14 thereby generating a
filtered signal to be used for determining the indication of the
interface integrity. The property determiner 16 determines the
indication of the interface integrity based on the filtered signal.
Since no phase shift is introduced, the indication of the interface
integrity can be accurately determined from the filtered
signal.
[0076] There are many possibilities for the zero-phase shift filter
14. In some implementations, the zero-phase shift filter 14 is a
moving average filter. In other implementations, the zero-phase
shift filter 14 is a symmetrical filter such as a Gaussian filter
or a Hamming filter. Other zero-phase shift filters are possible.
Additionally, many digital filters that introduce a phase shift
(such as a Butterworth filter for example) can be made to be
zero-phase by applying the filter on the data a second time but in
reverse order. More generally, a "zero-phase shift filter" can
include any appropriate combination of components that provide
suitable filtering with a zero net phase shift. Other
implementations are possible. Note that a "zero-phase shift filter"
ideally introduces no phase shift at all, but in practical
implementations might introduce a very small amount of phase shift.
Therefore, a zero-phase shift filter is characterised in that it
introduces no meaningful phase shift. Any phase shift introduced by
such a filter is not detectable or is negligible for the purposes
described herein.
[0077] It is to be understood that the "abutment connected to the
implant" does not necessarily mean that the abutment and the
implant are formed of separate members. In some implementations,
the abutment and the implant are formed of a same continuous
member. In this manner, although the abutment and the implant are
referred to separately, they are still part of the same continuous
member. In other implementations, the abutment and the implant are
formed of separate members.
[0078] There are many possibilities for the indication of the
interface integrity. In some implementations, the indication is an
explicit indication of the interface integrity. In these
implementations, a measure of the interface integrity is determined
by the property determiner 16. In other implementations, the
indication is an implicit indication of the interface integrity. In
these implementations, a measure of the interface integrity may not
have been determined, but at least a variable or parameter has been
determined that is indicative of the interface integrity. Such
variable or parameter might for example be the contact time or the
natural frequency of the system. Note that the contact time and the
natural frequency of the system are not explicit measures of the
interface integrity, but are still indicative of the interface
integrity.
[0079] There are many ways for the property determiner 16 to
determine the indication of the interface integrity. In some
implementations, the property determiner 16 determines a natural
frequency of the system based on a contact time measured from the
filtered signal. Note that since no phase shift is introduced, the
indication of the contact time can be accurately determined from
the filtered signal. Upon determining the contact time, the natural
frequency of the system can be determined. Finally, the property of
the interface can be determined based on the natural frequency. In
some implementations, this is performed by applying the natural
frequency to predetermined correlations or look-up tables. In other
implementations, this is performed by applying the natural
frequency to a predetermined mathematical model for the system.
Note that the "system" includes many components such as the implant
22 and the abutment 24, and may include other components and/or
considerations depending on the complexity of the model. Further
details of system modelling are provided later.
[0080] In the illustrated example, signal processing is performed
by the signal processor 10. More generally, signal processing can
be implemented by hardware, firmware, software, or any appropriate
combination thereof. For software implementations, there is
provided a computer readable medium having computer executable
instructions stored thereon for execution on a processor for
implementing functionality described herein.
[0081] Referring now to FIG. 6, shown is a flowchart of a method of
processing a signal for determining an indication of an interface
integrity between a medium and an implant that is at least
partially embedded therein. This method can be implemented by a
signal processor, for example by the signal processor 10 shown in
FIG. 5. More generally, this method may be implemented in any
appropriate apparatus.
[0082] At step 6-1, the apparatus receives a first signal generated
from a motion of an impact body during impact with an abutment
connected to the implant. According to an embodiment of the
invention, at step 6-2 the apparatus filters the first signal using
a zero phase shift filter thereby generating a filtered signal to
be used for determining the indication of the interface integrity.
Examples of zero phase shift filters that can be used have been
described above. At step 6-3, the apparatus determines the
indication of the interface integrity based on the filtered signal.
Examples for the indication of the interface integrity have been
provided above. Since no phase shift is introduced, the indication
of the interface integrity can be accurately determined from the
filtered signal. Examples of how this might be accomplished have
been provided above.
[0083] In the examples described above with reference to FIGS. 5
and 6, it is assumed that the zero-phase shift filter and the
property determiner are implemented by the same component. In
alternative implementations, they are implemented separately. For
example, the zero-phase shift filter might be included as part of
the impact body. An example of this is described below with
reference to FIGS. 7 and 8. Other implementations are possible. For
example, all of the signal processing could be performed by
hardware that is part of the impact body. For such implementations,
there is no need for a separate processor coupled to the impact
rod.
Another System and Method
[0084] Turning now to FIG. 7, shown is a schematic of another
apparatus 11A for determining an indication of an interface
integrity between the medium 20 and the implant 22 that is at least
partially embedded therein. The apparatus 11A has a signal
processor 10A connected to an impact body 26A via a coupling 28A.
The impact body 26A has a motion detector 27A, which might for
example be an accelerometer, and a zero-phase shift filter 14A. The
zero-phase shift filter 14A has an input 12A and an output 12C. The
signal processor 10A is not shown with any components for sake of
simplicity. The apparatus 11A may have other components, but they
are not shown for sake of simplicity.
[0085] In operation, a user impacts the impact body 26A against the
abutment 24. The impact body 26A might be accelerated towards the
abutment 24 for example via an electromagnet. The motion detector
27A translates the motion of the impact body 26A during impact into
a first signal. According to an embodiment of the invention, the
first signal is filtered with the zero phase shift filter 14A
thereby generating a filtered signal to be used for determining the
indication of the interface integrity. Examples of zero phase shift
filters that can be used have been described above. The filtered
signal is provided over the coupling 28A to another entity that
determines the indication of the interface integrity based on the
filtered signal. The coupling 28A is a wired connection, but in
alternative implementations might be a wireless connection. In this
example, the "other entity" is the signal processor 10A. The
property signal processor 10A determines the indication of the
interface integrity based on the filtered signal. Examples for the
indication of the interface integrity have been provided above.
Since no phase shift is introduced, the indication of the interface
integrity can be accurately determined from the filtered signal.
Examples of how this might be accomplished have been provided
above.
[0086] In the illustrated example, the zero-phase filter 14A is
implemented as hardware. More generally, the zero-phase filter 14A
can be implemented by hardware, firmware, software, or any
appropriate combination thereof. For software implementations,
there is provided a computer readable medium having computer
executable instructions stored thereon for execution on a processor
for implementing functionality described herein.
[0087] Referring now to FIG. 8, shown is a flowchart of another
method of processing a signal for determining an indication of an
interface integrity between a medium and an implant that is at
least partially embedded therein. This method can be implemented by
an impact body, for example by the zero-phase shift filter 14A of
the impact body 26A shown in FIG. 7. More generally, this method
may be implemented in any appropriate apparatus.
[0088] At step 8-1, the apparatus receives a first signal generated
from a motion of an impact body during impact with an abutment
connected to the implant. According to an embodiment of the
invention, at step 8-2 the apparatus filters the first signal using
a zero phase shift filter thereby generating a filtered signal to
be used for determining the indication of the interface integrity.
Examples of zero phase shift filters that can be used have been
described above. At step 8-3, the apparatus provides the filtered
signal to another entity that determines the indication of the
interface integrity based on the filtered signal. Examples for the
indication of the interface integrity have been provided above.
Since no phase shift is introduced, the indication of the interface
integrity can be accurately determined from the filtered signal.
Examples of how this might be accomplished have been provided
above.
Testing Apparatus
[0089] Referring now to FIG. 9, shown is a photograph of a testing
apparatus for an in vitro model. A disk is clamped in a circular
trough which is in turn mounted in a clamping device that also
supports the clamped Periotest handpiece. The clamped handpiece is
mounted on a microscope stage to allow adjustment of the position
of the rod relative to the abutment.
[0090] To simulate bone anchored implants, two implant/abutment
systems were chosen to simulate a range of implant applications.
The implants used were a 4 mm flanged extra-oral implant (4
mm.times.O3.75 mm, SEC 002-0, Entific Medical Systems, Toronto,
Ontario, Canada) and a 10 mm intra-oral implant (10 mm.times.O3.75
mm, Nobel Biocare, Toronto, Ontario, Canada). The implants were
mounted in 41 mm diameter discs of Photoelastic FRB-10 plastic
(Measurements Group Inc., Raleigh N.C., USA). Implants were
installed into the discs by drilling an appropriate diameter hole
and then cutting threads using a tap matched to the implant. The 4
mm implant was inserted into a disc of 5 mm thickness while the 10
mm implant was in a 10 mm thick disc. Both implants were secured to
the discs with epoxy cement (5 Minute Epoxy, Devcon, Danvers,
Mass., USA) to ensure as uniform an interface as possible. FRB-10
was chosen as its elastic modulus of 9.3 GPa is of the same order
as that reported for cortical bone and for dense cancellous bone
(1.3-25.8 GPa).
[0091] Two different abutments were used in the experiments, a
standard 5.5 mm (SDCA 005-0, Nobel Biocare, Toronto, Ontario,
Canada), and a standard 10 mm abutment (SDCA 043-0, Nobel Biocare,
Toronto, Ontario, Canada). The abutments were attached to the
implants using a torque wrench (DIB 038, Nobel Biocare, Toronto,
Ontario, Canada) and torqued to 20 Ncm unless otherwise specified.
The FRB discs were then mounted in a circular steel base that was
in turn mounted to a stand which also held the Periotest
handpiece.
[0092] The Periotest handpiece was mounted on a custom built
adjustable stand that allowed for vertical, horizontal and angular
rotations of the handpiece. The holder had two micrometer
attachments (Vickers Instrument Ltd., England) to control the
horizontal and vertical displacements. Handpiece angulation was
measured using a standard bevel gauge (not shown). The implant and
abutment were formed of a single aluminum post.
[0093] Mechanical properties and sizes of the components are given
in Table 1 and Table 2.
TABLE-US-00001 TABLE 1 Model Dimensions Oral Model Dimensions Post
Radius (P.sub.r) 2 mm Post Height (P.sub.h) 20 mm Abutment Height
(A.sub.h) 10 mm Engagement 9 mm Length (E.sub.L) Interface
Thickness (I.sub.t) 0.38 mm Interface Height (I.sub.h) 9 mm Disk
Radius (D.sub.r) 20 mm Disk Height (D.sub.h) 9 mm Periotest .RTM. 1
mm Periotest .RTM. 20 mm Rod Radius (R.sub.r) Rod Length (R.sub.L)
BAHA Model Dimensions Post Radius (P.sub.r) 2 mm Post Height
(P.sub.h) 20 mm Abutment Height (A.sub.h) 5 mm Engagement 4 mm
Length (E.sub.L) Interface Thickness (I.sub.t) 0.38 mm Interface
Height (I.sub.h) 9 mm Disk Radius (D.sub.r) 20 mm Disk Height
(D.sub.h) 9 mm Periotest .RTM. 1 mm Periotest .RTM. 20 mm Rod
Radius (R.sub.r) Rod Length (R.sub.L)
TABLE-US-00002 TABLE 2 Model Properties Young's Poisson's Density
Component Modulus (GPa) Ratio (kg/m.sup.3) FRB Disk 8.4 0.31 1800
Aluminium Post 73 0.32 2800 Acrylic Interface Layer 0.5 0.30 1800
Periotest .RTM. Rod 200 0.30 9.4 grams
[0094] To measure the Periotest signal and the un-modified
acceleration signal simultaneously, a DAP 5400a sampling card
(Microstar Laboratories, Bellevue, Wash., USA) with a sampling rate
of 2 MHz was used. The un-modified acceleration signal collected
from the Periotest handpiece was filtered by a moving average
filter, so as not to introduce phase shift and distortion of the
contact time. After filtering, the contact time was measured. The
fundamental mode dominates the response and with the removal of
higher frequency components in the signal the contact times
calculated serve as an approximation of the half period of
vibration of the system's first mode during impact. The resonant
frequency of the system was then calculated using
Freq = 1 2 ( ContactTime ) ##EQU00001##
[0095] To measure the extent of the differences between the moving
average filtered signal and the Periotest signal the contact times
for three different systems have been evaluated: [0096] 4 mm
implant with a 10 mm abutment to simulate a less stiff system
(longer contact time), [0097] 10 mm implant with a 3 mm abutment to
simulate a stiff system (shorter contact time), and -10 mm implant
with a 10 mm abutment to evaluate the intermediate case.
Results of Testing
[0098] Referring now to FIG. 10, shown is a graph depicting a
Periotest conditioned signal and a moving average filtered signal.
It can be seen that there is a significant difference between the
contact time based on the Periotest filtered signal (A to C) and
the contact time based on a signal that has been filtered with a
moving average filter (A' to C') which does not introduce any phase
shift in the signal. While this difference between the signals may
not be important for natural teeth with relatively long contact
times, it becomes very significant for the smaller contact times
associated with artificial implant measurements.
[0099] Due to the filtering distortion of the signal and the
limited resolution of the PTV scale it is preferable to measure the
resonant frequency based on an accelerometer signal which has not
been distorted. For an implant/abutment system with a PTV range
between -8 and 0 the resonant frequency will have values ranging
from 2700 Hz to 1300 Hz (higher frequency corresponds to a more
stable system and lower frequency a less stable system).
[0100] Referring now to FIG. 11, shown is a graph depicting a
comparison of contact times calculated based on the moving average
filtered signal and the conditioned Periotest signal. The
differences between the conditioned signal used by the Periotest to
calculate the PTV and the alternative signal conditioned using a
moving average filter technique, can be significantly
different--especially for more rigidly mounted implants. The
largest difference in contact time was 88 .mu.s (the 10 mm implant
with a 3 mm abutment), which is over 40% of the moving average
filtered value. The results show that as the stiffness of the
implant/abutment increases the difference between the Periotest
conditioned signal and the moving average filter increases. The
difference was 8% for a 4 mm implant and a 10 mm abutment, while
for the 10 mm implant with a 3 mm abutment this difference
increased to 40%.
[0101] To independently monitor the motion of the implant/abutment
system, a strain gauge was mounted on a separate abutment to
measure the bending strain during the impact by the Periotest rod.
A linear strain gauge, type EA-06-015EH-120 (Micro-Measurements,
Measurements Group Inc., Raleigh, N.C., USA), was mounted
vertically on the exterior surface of a 5.5 mm abutment on the side
impacted by the rod. The strain gauge was attached using M-Bond 200
(Micro-Measurements) adhesive and then coated with M-Coat D acrylic
(Micro-Measurements). The lead wires from the strain gauge were
0.005-inch diameter type 7X00157 (California Fine Wire, California,
USA). The strain gauge measurements utilized a DAP 5400a sampling
card (Microstar Laboratories, Bellevue, Wash., USA) with a sampling
rate of 2 MHz which could simultaneously monitor the strain gauge
signal, the moving average filtered accelerometer signal and the
Periotest acceleration signal. The strain gauge abutment was then
attached to the 4 mm implant and measurements were taken by
striking the top of the 5.5 mm abutment.
[0102] Referring now to FIG. 12, shown is a graph depecting a
Periotest conditioned signal, a moving average filtered signal and
a strain gauge signal. This shows one of the 16 strikes taken on
the 5.5 mm strain gauged abutment with a 4 mm implant. The contact
time based on the strain signal matches the moving average
accelerometer signal almost identically, while the Periotest
filtered signal shows a significantly longer contact time. The
analysis of the accelerometer signal from the handpiece coupled
with that from the strain gauge mounted on the abutment showed that
the moving average filtered signal is a better measure of the
actual motion of the implant/abutment system and provides a more
representative measure of the resonant frequency (and thus the
stiffness) of the system.
[0103] To evaluate the repeatability and reproducibility of the
measurement system, seven sets of five consecutive measurements
were taken on the 4 mm implant with a 5.5 mm abutment. The
handpiece was set at an angle of 5.quadrature. from an axis
perpendicular to the implant. The distance between the end of the
handpiece and the abutment was set to 1.5 mm. The micrometer was
set so that the Periotest rod would strike the rim of the 5.5 mm
abutment. Between each set of five readings the stand was moved and
then re-aligned to strike the rim of the abutment in an attempt to
replicate the previous set of readings.
[0104] Referring now to FIG. 13, shown is a chart depicting the
repeatability and reproducibility of the experimentation. The mean
resonant frequency of the 4 mm implant/5.5 mm abutment system for
35 readings (7 sets of 5 readings) was found to be 2083.+-.12 Hz
(n=35). Within a single group of five consecutive readings the
largest standard deviation was 12 Hz (n=5). Of the seven sets of
readings the lowest average value was 2070.+-.12 Hz and the highest
average reading was 2095.+-.3 Hz. The error bars on the plot (and
subsequent plots for following sections) are one standard deviation
of the measurements.
[0105] The repeatability and reproducibility measurements show that
for 95% confidence (.+-.2 standard deviations) the resonant
frequency can be determined to within .+-.24 Hz when using the
moving average filter. With a range of resonant frequencies between
1300 and 2700 Hz for implant/abutment systems this technique
provides 58 resolution steps while the PTV scale offers eight (PTV
readings between -8 and 0).
Conclusion
[0106] While the Periotest system as a whole has some shortcomings
when used to monitor implant integrity, the concept of an impact
test remains a viable one. The Periotest handpiece itself provides
a very convenient method to dynamically excite the implant/abutment
system. In fact, the Periotest handpiece has been incorporated as
part of a system to measure the damping capacity of materials. The
Periotest handpiece was used to develop an improved impact test to
monitor implant integrity. Alternate signal processing which avoids
the phase shift in the accelerometer signal reducing or eliminating
distortions in the contact time was analyzed. Additionally, the
effect of critical clinical parameters on the results of the
proposed technique was examined such that appropriate clinical
protocols could be developed.
Section II: Mathematical Model
Introduction
[0107] The drive for a clinically effective, non-invasive technique
for monitoring implant stability has led to a number of testing
methods based on the concept of resonant frequency. Resonant
frequency measurements are an indirect measure of the bone-implant
interface integrity and do not provide any specific measures of the
physical properties of the interface itself. While initial testing
by some researchers suggested that Periotest was an objective and
easily applied measurement technique for stability assessment of
implants, recent literature reviews of the Periotest discuss some
of the failings of the instrument including the effect clinical
variables have on the measurements as well as the reduced
resolution and low sensitivity when measuring implant-abutment
systems. Some of the inconsistencies in the reported Periotest
results may be due to a lack of understanding of how the system
being measured responds when excited.
[0108] For instance, there has been some debate as to what the
higher frequency component found in the raw accelerometer signal
(see FIG. 4) represents. Some suggest that the higher frequency is
a result of partial separation between the impact tool and the
implant, resulting in a "bouncing" affect. It has also been
hypothesised that this frequency is merely electrical noise on the
accelerometer signal or the second mode of vibration of the
implant-tissue system. Simulations and modelling are performed to
understand the source of this component of the signal and if it can
be used to better understand the status of the interface. The
higher frequency component can potentially be used to glean more
information about the integrity of the implant interface.
[0109] In order to gain a greater understanding of the Periotest
measurement system, mechanical models of the system have been
developed. One approach was to model a Periotest impact. The system
shown in FIG. 1 was modeled analytically as a single degree of
freedom system in which the implant-abutment was assumed to be a
rigid body pinned at the implant base. The model was used to
estimate the force of the impact and relate the PTV to an overall
equivalent interface stiffness. However, analytical results were
purely theoretical and were not verified by directly comparing the
theoretical results to in vitro or in vivo experiments.
[0110] A subsequent approach involved a two degree of freedom
analytical model in which the bone-implant properties were modeled
as a series of springs acting along the length of the implant.
Model results were then correlated to in vitro measurements for
extraoral implant-abutment systems. The in vitro testing and model
results determined that implant diameter, length of engagement
between bone and implant, angulation of Periotest handpiece and
striking height along the abutment all influenced the output of the
Periotest. An in vivo patient study was also attempted, however,
results were inconclusive due to what the authors believed was a
poor understanding of the effects due to measurement parameters and
lack of a rigorous clinical testing protocol. Unfortunately, this
study utilized the filtered accelerometer signal from the Periotest
which, as previously discussed, is not an accurate reflection of
the impact response.
[0111] More recently, FEA was used for the system shown in FIG. 1
to produce a complete transient simulation of the impact by the
rod. This study utilized the un-filtered (raw) accelerometer
response and compared finite element solutions with in vitro data
for oral and extraoral implants. It was shown that the stiffness of
the components as well as the junctions between them significantly
affect the overall response and that the implant and abutment do
not act as a single rigid body during the contact. One difficulty
in using this technique was the very long processing time involved
and the necessity of doing a somewhat imprecise frequency analysis
on the transient response.
[0112] It is desired to develop a better understanding of the
dynamics that occur during the impact and how this affects the
accelerometer response during measurements. To achieve this, an
analytical model of the implant/abutment/Periotest system is
developed to aid in interpreting the acceleration signal and in
particular how the supporting bone properties affect this signal
over a range of implant applications. Analytical model results are
compared directly to in vitro measurements. Studies done with the
Periotest often erroneously refer to the device as measuring the
damping characteristics of the interface. The analytical model can
be used to help clear some of the confusion about what bone
properties are currently being measured. Additionally, the
developed analytical model can be used to simulate changes in the
bone stiffness supporting the implant and to determine the effect
of bone loss around the neck of the implant. Finally, since some
implants currently used incorporate a flange, the model can be used
to understand the influence the flange has on the impact
accelerometer response.
System and Method
[0113] Turning now to FIG. 14, shown is a schematic of an apparatus
31 for determining a property of an interface between a medium 40
and an implant 42 that is at least partially embedded therein. The
property of the interface might for example be a measure of the
integrity of the interface. The measure of the integrity of the
interface might for example be a stiffness of the interface. An
abutment 44 is connected to the implant 42. The apparatus 31 has a
signal processor 30 connected to an impact body 46 via a coupling
48. The impact body 46 has a motion detector 47, which might for
example be an accelerometer. The signal processor 30 has an input
32, a filter 34, and a property determiner 36. The property
determiner 36 has a mathematical model 37 for impacting the impact
body 46 against the abutment 44. The apparatus 31 may have other
components, but they are not shown for sake of simplicity.
[0114] In operation, a user impacts the impact body 46 against the
abutment 44. The impact body 46 might be accelerated towards the
abutment 44 for example via an electromagnet. The motion detector
47 translates the motion of the impact body 46 during impact into a
first signal, which is provided to the signal processor 30 over the
coupling 48. The coupling 48 is a wired connection, but in
alternative implementations might be a wireless connection. The
signal processor 30 receives the first signal over the input 32.
The first signal is filtered by the filter 34 thereby generating a
filtered signal to be used for determining the property of the
interface. The property determiner 36 determines a system property
based on the signal. The system property might for example be a
natural frequency of the system. According to an embodiment of the
application, the property determiner 36 analytically determines the
property of the interface by applying the system property to the
mathematical model 37. Therefore, the property determiner 36 solves
for the property of the interface based on the mathematical model
37 and the system property that has been determined.
[0115] It is to be understood that the "abutment connected to the
implant" does not necessarily mean that the abutment and the
implant are formed of separate members. In some implementations,
the abutment and the implant are formed of a same continuous
member. In this manner, although the abutment and the implant are
referred to separately, they are still part of the same continuous
member. In other implementations, the abutment and the implant are
formed of separate members.
[0116] In the illustrated example, the first signal is filtered by
the signal processor 30. In other implementations, the first signal
is filtered before reaching the signal processor 30. In some
implementations, a zero-phase shift filter is implemented. Example
zero-phase shift filters that can be used have been described
above. In other implementations, the first signal is not filtered
at all.
[0117] There are many possibilities for the mathematical model 37.
The mathematical model 37 can have varying complexity depending on
how many components and/or considerations the model is to include.
In some implementations, the mathematical model 37 has
three-degrees of movement. In other implementations, the
mathematical model 37 has four-degrees of movement. Other
implementations are possible. Example mathematical models are
provided below.
[0118] In the illustrated example, the property determiner 36 is
implemented by a signal processor. More generally, the property
determiner 36 can be implemented by hardware, firmware, software,
or any appropriate combination thereof. For software
implementations, there is provided a computer readable medium
having computer executable instructions stored thereon for
execution on a processor for implementing functionality described
herein.
[0119] Referring now to FIG. 15, shown is a flowchart of a method
of determining a property of an interface between a medium and an
implant that is at least partially embedded therein. The property
of the interface might for example be a measure of the integrity of
the interface. The measure of the integrity of the interface might
for example be a stiffness of the interface. This method can be
implemented by a signal processor, for example by the property
determiner 36 of the signal processor 30 shown in FIG. 14. More
generally, this method may be implemented in any appropriate
apparatus.
[0120] At step 15-1, the apparatus maintains a mathematical model
for impacting an impact body against an abutment connected to the
implant. Example mathematical models are provided below. At step
15-2, the apparatus receiving a signal generated from a motion of
the impact body during impact with the abutment. At step 15-3, the
apparatus determines a system property based on the signal. The
system property might for example be a natural frequency of the
system. According to an embodiment of the invention, at step 15-4
the apparatus analytically determines the property of the interface
by applying the system property to the mathematical model.
Therefore, the apparatus solves for the property of the interface
based on the mathematical model and the system property that has
been determined.
In Vitro Experimental Model
[0121] Referring now to FIG. 16, shown is a schematic drawing of
the in vitro experimental model for impact testing. The in vitro
experimental model was developed to assist in the development of a
measurement protocol and for validation of analytical and numerical
models. The model is approximately the size of an oral
implant-abutment system. It includes an aluminum post fixed with
acrylic into the centre of a disk of FRB-10 (Measurements Group
Inc, Raleigh, N.C., USA).
Finite Element Analysis (FEA) Model
[0122] Finite element analysis (FEA) models have been used by
various researchers to attempt to find relationships between
natural frequencies and the surrounding conditions of the implant.
One approach involved modal analysis of the implant-tissue system
to investigate how bone type and bone density affects resonant
frequency. In the current work, FEA is used to produce a more
thorough dynamic model of the implant-abutment by including the
impact of the Periotest rod with the implant.
[0123] Referring now to FIG. 17, shown is a schematic of an FEA
model for an impact test. The finite element model created to
simulate this in vitro model used ANSYS 7.1 (ANSYS Inc, Canonsburg,
Pa., USA) on a personal computer. Containing approximately 15000
elements, the model includes the in vitro geometry and the
Periotest rod. Only one half of the structure is considered due to
symmetry resulting in decreased processing time used to arrive at a
solution. The model was meshed with tetrahedral elements with
mid-sided nodes. These quadratic elements are comprised of 10 nodes
having three degrees of freedom at each node: translation in the
nodal x, y, and z directions. Element properties and geometric
values were matched to those of the in vitro model, as listed in
Table 1 and Table 2. Convergence testing was performed to ensure
the mesh was adequately dense such that solutions did not change
more that 1% when the element size was halved.
[0124] To model the impact between the rod and the aluminum post,
contact elements were created between the two adjacent surfaces so
that the rod and the rest of the system can move independently of
each other without allowing the rod to penetrate the post. This was
done using a combination of 3-D eight node, surface to surface
contact elements which are used to represent contact and sliding
between three dimensional deformable surfaces and 3-D target
elements which overlay the solid elements describing the boundary
of the deformable body. Since the impact is direct, sliding is
assumed to be negligible thus friction coefficients were ignored to
save processing time. The Periotest rod was constrained to move
horizontally and was assumed to have an initial velocity of 0.2 m/s
towards the aluminum post to match the manufacturer's
specifications of the Periotest's performance. A transient analysis
was used to determine the motion of the system with a typical
sampling rate of once every 0.6 microseconds. In cases where
greater resolution was desired, this was increased such that the
highest resolved frequency was 20 times faster than the highest
desired frequency as recommended by ANSYS.
[0125] As mentioned previously, one of the goals of the model was
to simulate changes in the natural frequencies of the
implant-tissue-Periotest system due to changes in the status of the
interface. Specifically, these include loss of osseointegration,
loss of bone margin height and development of connective soft
tissue in the bone-implant interface. This was accomplished through
slight modifications to the interface region of the model. To
ensure a smooth transition of the elements from the relatively
small interface elements to the disk, the ANSYS element expansion
function was utilized to keep aspect ratios small while expanding
each consecutive element by 130% until the specified disk element
size was reached.
[0126] For a fully osseointegrated implant, the implant and
interface layer shared nodes along their common boundaries and thus
allowed no separation between the two. When a loss of
osseointegration was simulated, the implant and interface layer no
longer shared nodes along the common boundaries. Instead a layer of
contact elements were meshed between the two to allow separation in
the area of osseointegration loss but not penetration. The nodes
below this loss still coincided however. For the simulation of
reduced bone margin height, the height of the interface layer was
reduced to simulate receding bone around the neck of the implant.
From a mechanical viewpoint, the difference is that while in both
of these cases there is no possibility of generating tensile forces
between the implant and the surrounding tissue in the area of loss,
compressive forces can be generated in the non-osseointegrated
case. The development of connective soft tissue in the interface
layer was simply modelled as a reduction in the stiffness of the
entire interface layer.
[0127] To verify the FEA model, the implant-abutment system while
being impacted with the rod at the free end was modelled as a
Bernoulli-Euler beam fixed at one end with a point mass at the
other. The solution for this problem results in the transcendental
equation for the frequency parameter (.beta.L). For the beam
parameters given in Table 1 and Table 2, the first two values of
(.beta.L) are 0.5776 and 3.9311. The natural frequency of vibration
of this system can be determined using the following:
p = ( .beta. L ) 2 ( EI m b L 3 ) ( 1 ) ##EQU00002##
where p is natural frequency, E is Young's modulus of the
cantilever, I is the second moment of area about the neutral axis,
L is the length of the cantilever and m.sub.b is the mass of the
cantilever.
[0128] Simulations were for the typical implant-abutment systems
mentioned previously (extraoral prostheses: 4 mm implants with 5 mm
abutments and oral implants: 9 mm implants with 10 mm abutments).
In all instances it was assumed that the rod impacts the top of the
abutment. For simulation of the loss of osseointegration and bone
loss, it was assumed that this begins at the outer surface of the
hard tissue and propagates toward the base of the implant. In the
extraoral case, the dimensions of the model were altered to those
found in Table 1.
[0129] Referring now to FIG. 18, shown is a graph depicting a
typical transient analysis signal for the FEA model of FIG. 17.
Note that the transient analysis signal is produced assuming a
system with an infinitely stiff disk and interface layer resulting
in a rigidly fixed, 10 mm long cantilevered aluminum post. Even
though transient analysis signal is a displacement/time result, it
is very similar to that of the raw experimental result shown in
FIG. 4. For instance, the transient analysis signal shows two
natural frequencies. If the signals are assumed to be a combination
of harmonic functions, the acceleration signal is equivalent to a
scaled displacement signal. Therefore the transient signal can be
related to the experimental signal. As with the experimental
result, the simulated signal appears to have a fundamental
frequency with a higher frequency superimposed. When the contact
status between the Periotest rod and the implant was reviewed, it
was found that they remained in contact throughout the strike. The
transient signals were analyzed using the custom software to
determine the two natural frequencies, which appeared to combine to
produce the characteristic signal.
[0130] These FEA results were compared to the frequencies predicted
by Equation (1). The results for the FEA model and analytical
solution for the first natural frequency were 2728 Hz and 2711 Hz
respectively and 127 kHz and 126 kHz for the second natural
frequency. Therefore the FEA model was within 0.8% relative error
of the analytical solution for both cases. As this is below the
convergence criterion of 1%, the model yielded accurate
results.
[0131] The finite element simulation was also compared to the in
vitro model. The FEA simulation parameters were set to those found
in Table 1 and Table 2, and the results were compared to the
results of in vitro testing. With the rod impacting the top of the
post, the FEA produced results of 1726 Hz and 46 kHz for the first
two natural frequencies respectively, while the in vitro tests
produced 1790 Hz and 40 kHz (averaged from the three tests each on
the six identical post/disk systems). This equates to 3.5% relative
error for the first natural frequency and 13% for the second. Again
the status of the contact elements remained closed throughout the
entire strike.
[0132] Since no separation was found to occur during the strike in
the finite element analysis and the model predicted the second
natural frequency for both the analytical and in vitro cases, there
is strong evidence that the higher frequency in the experimental
signal is indeed the second mode of vibration of the
implant-abutment system.
[0133] Referring now to FIGS. 19A through 19D, shown are graphs
depicting changes in first and second natural frequencies of the
implant abutment as a function of increasing loss of
osseointegration and bone margin height. These results were
obtained from the FEA model shown in FIG. 17. FIGS. 19A and 19B
show graphs depicting changes in first and second natural
frequencies of the implant abutment as a function of increasing
loss of osseointegration and bone margin height for a 10 mm
abutment (oral implant). FIGS. 19C and 19D show graphs depicting
changes in first and second natural frequencies of the implant
abutment as a function of increasing loss of osseointegration and
bone margin height for a 5 mm abutment (extra-oral implant).
[0134] It has been assumed that the region of loss begins at the
outer surface (skin side) of the hard tissue and propagates towards
the base of the implant. The error bars for the first natural
frequency plots represent the difference between to adjacent data
points when calculating the contact time. For instance, there may
not be a data point exactly on the zero displacement axis, thus it
lies somewhere between the point before and after the axis
crossing. The error bars for the second natural frequency plots
represent the FFT resolution.
[0135] For the first (lowest) natural frequency (FIGS. 19A and
19C), both sizes of implant-abutments evaluated show measurable
changes for relatively small regions of loss. As it has been
reported that changes equivalent to 100 Hz are statistically
significant, a loss of approximately 0.2 mm would be detectable for
the shorter implants and 0.4 mm for the longer system. The
difference in loss (osseointegration vs. bone loss) is not
distinguishable until the height of loss has extended to
approximately 0.8 mm for the shorter implant and to approximately
1.9 mm for the longer (depicted as "h" in FIGS. 19A-19D). While the
second (higher) natural frequency (FIGS. 19B and 19D) show a
similar trend as the length of the loss zone increases, the
differences between loss of osseointegration and bone loss are not
as evident.
[0136] Referring now to FIGS. 20A through 20D, shown are graphs
depicting changes in first and second natural frequencies of the
implant abutment as a function of increasing interface layer
stiffness. These results were obtained from the FEA model shown in
FIG. 17. FIGS. 20A and 20B show graphs depicting changes in first
and second natural frequencies of the implant abutment as a
function of increasing interface layer stiffness, for a 10 mm
abutment (oral implant). FIGS. 20C and 20D show graphs depicting
changes in first and second natural frequencies of the implant
abutment as a function of increasing interface layer stiffness, for
a 5 mm abutment (extra-oral implant).
[0137] The simulations for the development of a softer interface
layer, which could correspond to the development of connective soft
tissue or reduced stiffness during healing, are given in FIGS.
20A-20D for the two sizes of implants with similar error bars to
the previous plots. FIGS. 20A and 20C show the dramatic change in
the lowest natural frequency as the stiffness (modulus of
elasticity) of the interface layer changes. The region between the
dashed lines is an estimated range of modulus of elasticity for
soft connective tissue (scar tissue) to hard tissue (quality bone)
and it is evident that the lowest natural frequency can change in
the order of 50% and will therefore be easily detectable. The
higher natural frequency also shows a similar dramatic change as
the modulus changes. It should be remembered that for these
simulations, the change in stiffness occurs over the entire
interface simultaneously.
[0138] If the change in stiffness were only over a portion of the
engagement length, then the change in the natural frequency would
not be as large as shown. The situation depicted does simulate the
change in overall stiffness, which is expected to occur during the
healing after implant placement. After initial placement, the
supporting hard tissue is believed to first "soften" as it begins
to remodel. This is followed by a period in which the stiffness
increases as osseointegration occurs. This would mean that the
natural frequencies would fall slightly from their initial values
then increase as osseointegration occurs. If it does not and the
frequency does not increase this would be a signal that soft
connective tissue is developing instead of the osseointegrated bond
desired.
[0139] The interface model used for the simulations above was
verified by comparison to the in vitro experiment as well as an
analytical solution. The first and second natural frequencies
predicted numerically were within 1% of the analytical result. The
comparison to the in vitro results produced not only a close
comparison in its frequency content, but indicated that the
impacting rod remained in contact with the abutment and did not
"bounce" as had been previously speculated. Instead, the higher
frequency, which had been seen previously in similar tests, related
to a predictable second mode of vibration of the implant-abutment
system that was also excited by the impact.
[0140] The results of the simulation above indicate that clinical
changes in the integrity of the interface should be detectable from
the frequency response changes. The simulations indicate that with
either a loss of osseointegration or bone margin height for the
shorter implant of as little as 0.2 mm, the change in frequency
response is sufficient to be clinically detectable. In addition,
changes in the stiffness of the interface, such as might occur
after initial implant placement or through the development of
connective soft tissue, result in easily measurable frequency
changes. All of the simulations indicate that the use of an impact
test can produce clinically meaningful results using the lowest
natural frequency excited by the impact.
Four-Degree of Freedom Model
[0141] A four-degree of freedom analytical model has been developed
to interpret measurement results of an impact testing method based
on the Periotest handpiece. Model results are compared to a variety
of in vitro tests to verify model predictions and to gain an
understanding of the parameters influencing the measurements. Model
simulations are then used to predict how changes in the supporting
stiffness properties, material loss around the neck of the implant
and the presence of an implant flange will affect the
measurements.
Analytical Model Development
[0142] Referring now to FIG. 21, shown is a schematic of a
four-degree of freedom model for the impact system. The model has
an implant 51, partially embedded in a supporting material 55. An
abutment 52 is affixed to the implant 51. A striking rod 56 is also
present. The symbols used are shown in Table 3.
TABLE-US-00003 TABLE 3 Symbol Definitions Variable Definition
X.sub.1 Coordinate describing horizontal position of impacting rod.
X.sub.2 Coordinate describing horizontal position of specific point
on abutment. .THETA..sub.1 Coordinate describing angular rotation
of the abutment. .THETA..sub.2 Coordinate describing angular
rotation of the implant. K.sub.I Impact stiffness (K.sub.SYS and
K.sub.DEF in series. K.sub.T Torsional stiffness at the
implant-abutment joint. k Stiffness of bone-implant interface (per
unit length). O Position along abutment longitudinal axis that
crosses the line of impact. G.sub.1 Location of abutment center of
gravity. G.sub.2 Location of implant centare of gravity. L.sub.A
Length of the abutment. L.sub.I Length of the implant. L.sub.O
Vertical distance from the top of the abutment to point O. h.sub.1
Vertical distance from O to G.sub.1. L.sigma. Distance joint is
above the supporting material surface. b Radius of implant-abutment
system.
[0143] In this model, both the implant 51 and the abutment 52 are
treated as separate rigid bodies, while the impact rod 56 is
treated as a point mass with mass m.sub.R. The implant and abutment
are assumed to be connected by a pin and torsional spring of
stiffness K.sub.T.
[0144] During the time that the abutment and impact rod are in
contact, the dynamic response of the system is described using the
coordinates X.sub.1 (displacement of the impacting rod), X.sub.2
(displacement of a point, O, along the central axis of the abutment
at the same height as the striking rod), .theta..sub.1 (rotation of
the abutment) and .theta..sub.2 (rotation of the implant) as shown
in FIG. 21. The stiffness, k, of the supporting material 55 is
represented by a series of distributed horizontal and vertical
springs along the length of the supporting material 55. This
stiffness is assumed to be uniform and constant in both the
vertical and horizontal directions. The supporting stiffness, k, is
an equivalent stiffness. Although the impact rod and abutment are
modeled as rigid bodies, there is in fact some deformation, which
occurs during the impact. A spring of stiffness K.sub.I is
introduced between the rod and the abutment to account for these
deflections. Similarly, the torsional spring (K.sub.T) is used to
approximate bending or flexibility about the implant-abutment
joint. Although the abutment is modeled as rigid, FEM studies have
shown that bending can occur. To account for this, the torsional
spring is used to attempt to account for the relative motion
(bending) within the abutment. While simulating bending with two
rigid bodies (the implant and abutment) and a torsional spring is
quite simplistic, it does provide an estimation of the bending
while minimizing the added complexity of the mathematical
model.
[0145] To estimate the damping properties of the supporting bone,
proportional damping was used in which the damping matrix is
assumed to be proportional (proportionality coefficient .beta.) to
the stiffness matrix so that the equations of motion become
[M]{{umlaut over (x)}}+.beta.[K]{{dot over (x)}}+[K]{x}={0}.
Assuming proportional damping allows for normal mode analysis to be
utilized and simplifies the analytical solution.
[0146] Some implants have a flange, which is modeled at 54. For the
flanged extraoral implants an additional flange stiffness (K.sub.F)
was added. While the flange 54 may be providing support across its
entire surface, K.sub.F is represented as a single equivalent
stiffness at an effective distance r from the center of the implant
51 as shown in FIG. 21. The added spring K.sub.F provides forces in
either tension or compression representing a flange osseointegrated
to the supporting bone structure. Calculations including K.sub.F
assume the flange is bonded to the supporting surface and represent
a maximum contribution to the implant support. For implants without
a flange, K.sub.F was set to zero.
[0147] The equations of motion for the four degree of freedom
analytical model illustrated in FIG. 21 are detailed below.
[ M ] = { X 1 X 2 .THETA. 1 .THETA. 2 } + .beta. [ K ] { X . 1 X .
2 .THETA. . 1 .THETA. . 2 } + [ K ] { X 1 X 2 .THETA. 1 .THETA. 2 }
= { 0 } ##EQU00003##
where [M] contains constants which describe the mass properties of
each of the elements in the system and [K] contains constants which
describe the stiffness or flexibility of the various components of
the system. Both [M] and [K] are influenced by the geometry
(lengths, etc.) of the various components in the system.
[ M ] = [ M 11 M 12 M 13 M 14 M 21 M 22 M 23 M 24 M 31 M 32 M 33 M
34 M 41 M 42 M 43 M 44 ] ##EQU00004## with [ M 11 ] = m R ;
##EQU00004.2## [ M 12 ] = [ M 21 ] = [ M 13 ] = [ M 31 ] = [ M 14 ]
= [ M 41 ] = 0 ; [ M 22 ] = m A + m I ; ##EQU00004.3## [ M 23 ] = [
M 32 ] = - ( m A h 1 + m I ( L A 2 + h 1 ) ) ; [ M 24 ] = [ M 42 ]
= - m I L I 2 ; [ M 24 ] = [ M 42 ] = - m I L I 2 ; [ M 33 ] = J A
+ m A h 1 2 + m I ( L A 2 + h 1 ) 2 ; [ M 34 ] = [ M 43 ] = m I ( L
A 2 + h 1 ) L I 2 ; [ M 44 ] = J I + m I ( L I 2 ) 2 ;
##EQU00004.4## and [ K ] = [ K 11 K 12 K 13 K 14 K 21 K 22 K 23 K
24 K 31 K 32 K 33 K 34 K 41 K 42 K 23 K 44 ] ##EQU00004.5## with [
K 11 ] = K I ; [ K 12 ] = [ K 21 ] = - K I ; [ K 13 ] = [ K 31 ] =
[ K 14 ] = [ K 41 ] = 0 ; [ K 22 ] = K I + 2 k ( L I - L C ) ; [ K
23 ] = [ K 32 ] = - 2 k ( L I - L C ) ( L A 2 + h 1 ) ; [ K 24 ] =
[ K 42 ] = - k ( L I 2 - L C 2 ) ; [ K 33 ] = 2 k ( L I - L C ) ( L
A 2 + h 1 ) 2 + K T ; [ K 34 ] = [ K 43 ] = k ( L A 2 + h 1 ) ( L I
2 - L C 2 ) - K T ; [ K 44 ] = 2 3 k ( L I 3 - L C 3 ) + 2 kb 2 ( L
I - L C ) + K T + r 2 K F . ##EQU00004.6##
[0148] The implants and abutments were treated as solid cylinders
with mass moments of inertia of J.sub.I and J.sub.A for the
implants and abutments, respectively. Implant masses (m.sub.i) were
measured and found to be 0.1538 and 0.647 gm for the 4 mm and 10 mm
implant The abutment masses (m.sub.A), were measured as 0.228,
0.333, 0.448, 0.647 gm for the 4, 5.5, 7, 10 mm abutments. The
lengths L.sub.A and L.sub.I refer to the length of the abutments
and implants, with L.sub.C and h.sub.1 referring to the implant
height above the bone level (for cases of bone resorption) and the
distance between the center of gravity of the abutment and the
striking point, respectively. The radius of the implant and
abutment is b and the effective radius of the flange support is r.
Using the normal mode method, the equations of motion were
uncoupled using the concept of the modal matrix. The general
solution then takes on the form of the summation of each of the
uncoupled solutions such that
{ X 1 ( t ) X 2 ( t ) .THETA. 1 ( t ) .THETA. 2 ( t ) } = r = 1 4 -
v r p r t [ 1 p r { u } r T [ M ] { x . ( 0 ) } sin ( 1 - V r 2 p r
t ) ] { u } r . ( 2 ) ##EQU00005##
where v.sub.r is the damping ratio for each mode, p.sub.r is the
resonant frequency for each mode and {.mu.}.sub.r is a column
vector of the normalized modal matrix [.mu.] and {dot over (x)}(0)
is the initial velocity of the system before impact.
[0149] The damping ratio v.sub.r for each mode can be determined
from .beta. as
.beta. = 2 ( v r ) p r . ( 3 ) ##EQU00006##
The value for .beta. was found by setting r=2 and choosing the
damping ratio v.sub.2 to match the in vitro measurements. Once
.beta. is known, Equation (3) can be used to solve for v for each
mode.
[0150] The acceleration response can be obtained by taking the
second derivative of Equation (2) to give
{ X 1 ( t ) X 2 ( t ) .THETA. 1 ( t ) .THETA. 2 ( t ) } = r = 1 4 (
3 v r 2 - v r 4 - 1 ) p r - v r p r t [ { u } r T [ M ] { x . ( 0 )
} sin ( 1 - v r 2 p r t ) ] { u } r . ( 4 ) ##EQU00007##
Knowing that the initial velocity of the system is
{ x . ( 0 ) } = { v o 0 0 0 } , ##EQU00008##
with v.sub.o assumed to be 0.2 m/s (according to the Periotest
manufacturer), Equation (4) can be solved to determine the
acceleration of the striking rod, {umlaut over (X)}.sub.1(t), which
can be compared to the measured accelerometer signal on the
rod.
[0151] Due to the nature of these equations a fundamental aspect of
the solution is that there will be four distinct frequencies
associated with the motion of the system. The frequencies can
alternatively be found by calculating the determinate of another
matrix, [K]-.omega..sup.2[M] and solving for those values of co
that make the determinate zero.
Determinate of [ K 11 K 12 K 13 K 14 K 21 K 22 K 23 K 24 K 31 K 32
K 33 K 34 K 41 K 42 K 43 K 44 ] - .omega. 2 [ M 11 M 12 M 13 M 14 M
21 M 22 M 23 M 24 M 31 M 32 M 33 M 34 M 41 M 42 M 43 M 44 ] = 0
##EQU00009##
This results in a (long) equation of the form
A.omega..sup.8+B.omega..sup.6+C.omega..sup.4+D.omega..sup.2+E=0
(5)
which is used to determine the four values of .omega..
[0152] During the impact test, the lowest natural frequency of the
system is determined from the accelerometer signal contact time.
Equation (5) is then used to determine the interface stiffness k
that would produce the same lowest natural frequency in the model.
This is how the model is used, in conjunction with the
accelerometer signal, to estimate the interface stiffness k.
[0153] The simulation of the impact response of this system is
calculated from the initial value problem in which all the
coordinates are initially zero with only the mass m.sub.R having an
initial velocity v.sub.o.
[0154] As the ultimate goal of this model is to understand how the
impact response is related to the supporting bone properties (and
changes in these properties), model stiffness (K.sub.T and
K.sub.I), and inertia properties were estimated prior to using the
model.
[0155] Model results were obtained through the use of a custom
Matlab program that solved the equations outlined above for the
model of FIG. 21. There will be four resonant frequencies p.sub.1
to p.sub.4. The lowest (fundamental) frequency is represented as
p.sub.1 with p.sub.2, p.sub.3 and p.sub.4 corresponding to the
higher frequencies in increasing order. To avoid confusion,
.omega..sub.1 denotes the measured first mode resonant frequency
determined from the impact responses. The governing equations for
the system shown in FIG. 21 were used in one of two ways: [0156]
the support stiffness k was specified and the Matlab model would
determine the natural frequencies (p.sub.1 to p.sub.4) and the
acceleration response of the impact in the time domain, and [0157]
the measured first mode frequency (.omega..sub.1) was given and the
Matlab model would determine the support stiffness k and the
acceleration response of the impact in the time domain. The
analytical model results were compared to measurements only during
the contact time and in all cases the model was checked to ensure
that only compressive forces existed between the impact rod and the
abutment during this interval, as this is the only interval over
which the model results are valid.
Analytical Model Parameters
[0158] To calculate the support stiffness and damping properties
for the in vitro implants the appropriate stiffness values for the
internal components in the system were first calculated. The
internal stiffness of K.sub.I, K.sub.T and K.sub.F were estimated
through a combination of in vitro experimentation to directly
determine stiffness values and comparison of model results to
specific in vitro measurements. Once the internal stiffness
components were determined the support stiffness and damping values
for extraoral and intraoral implants could be estimated.
[0159] The impact stiffness, K.sub.I, was estimated directly by
clamping a steel block on one side of the abutment while impacting
the opposite. The purpose of this is to attempt to isolate the
abutment from the support at the implant. The Periotest handpiece
was placed in the holding stand and a series of five measurements
were taken on 10, 7, 5.5 and 4 mm abutments which were connected to
the flanged 4 mm extraoral disk with 20 Ncm of torque. By assuming
the steel backing is rigid, the impact stiffness could be
calculated from the measured first mode frequency as
K.sub.I=(2.pi..omega..sub.1).sup.2m.sub.R (6)
The results for the different abutment sizes are shown in the
following Table 4.
TABLE-US-00004 TABLE 4 Calculated impact stiffness (K.sub.I) for
different length abutments. .omega..sub.1 (Hz) K.sub.I (N/m)
.times. 10.sup.6 10.0 mm Abutment 2624 .+-. 22 2.51-2.60 (2.56) 7.0
mm Abutment 2542 .+-. 16 2.37-2.43 (2.40) 5.5 mm Abutment 2690 .+-.
7 2.67-2.70 (2.68) 4.0 mm Abutment 2836 .+-. 13 2.96-3.01
(2.98)
The .omega..sub.1 values reported in Table 4 are the average of the
five readings with the standard deviation of the measurements for
each abutment. The impact stiffness (K.sub.I) is reported as a
range to reflect the measurement variation, with the K.sub.I from
the average .omega..sub.1 value shown in brackets. Table 4 shows
that for different length abutments there are variations in the
measured .omega..sub.1 values resulting in different values for
K.sub.I. These differences are likely due to geometric differences
which exist between the abutments. While the outer diameter for the
different length abutments are the same, the internal dimensions
and the connecting screw details differ. Due to these differences,
the K.sub.I value used with the analytical model was the average
value shown in brackets and was specific to each abutment.
[0160] The torsional stiffness K.sub.T was initially estimated
based on the assumption that the abutment behaves as a cantilever
beam such that
K T = 3 EI L T . ( 7 ) ##EQU00010##
The length of the cantilever, L.sub.T, is determined by
L.sub.T=L.sub.A-L.sub.O where L.sub.A is the length of the abutment
and L.sub.O is the distance from the top of the abutment to where
the Periotest rod strikes. Using E=110 GPa for titanium and
approximating the abutments as solid cylinders allows for the
calculation of K.sub.T.
[0161] Equation (7) assumes a rigid, fixed connection at the
implant/abutment joint. However, the screw connection between the
implant and abutment will not provide an ideal fixed connection. As
a result, the torsional stiffness (K.sub.T) determined in Equation
(7) will over-estimate the true torsional stiffness. Additionally,
Equation (7) does not take into account any shear effects, which
are likely to be important since the abutments are relatively
short. To account for the effects of shear and a non-ideal joint
connection between the implant and abutment the value of K.sub.T is
adjusted empirically as discussed in the following section.
[0162] As a test of the developed model, the simulated un-damped
acceleration response (.beta.=0) was compared to one of the
measured acceleration signals for the 10 mm implant with a 10 mm
abutment. The 10 mm abutment was chosen as the amplitude of its
second mode was found to be largest and would better illustrate
model and measurement results. The results of the comparison can be
seen in FIG. 22A, which shows the acceleration response of 16
impact measurements and the predicted acceleration response from
the model using K.sub.I and K.sub.T as given above. A support
stiffness k=1.8-2.1 (1.9) GPa was found by matching the model first
mode frequency to the measured frequency .omega..sub.1=1500.+-.14
Hz determined from the impact tests. The range of k reflects the
variation in measured .omega..sub.1 with the k determined from the
average measured .omega..sub.1 in brackets. To directly compare the
model acceleration response to the measurements, the model response
was "normalized" by approximately matching the model acceleration
amplitude with the measured accelerometer signal amplitude (this
normalization was used since the calibration of the Periotest
accelerometer was unknown and its magnitude depends on the initial
speed of the Periotest rod which is also unknown for each
individual strike).
[0163] While the relative amplitudes of the accelerations between
the first and second modes are predicted well (the amplitudes of
the peaks appear to match), there is a noticeable second mode
frequency difference between the measured response and the model
results shown in FIG. 22A. The model appears to under-predict the
second mode frequency. To improve the predictions from the model,
the value of K.sub.T estimated previously in Equation (7) was
reduced to
K T = 0.26 ( 3 EI L T ) ( 8 ) ##EQU00011##
to account for any additional flexibility or shear effects as
described previously.
[0164] With the reduced K.sub.T, a new support stiffness k=6.8-8.4
(7.5) GPa was determined for the 10 mm implant and 10 mm abutment.
The resulting (un-damped) signal, shown in FIG. 22B, shows much
better agreement with the measured signal. While the geometry of
the supporting material will affect the relationship between k and
the elastic modulus, the average determined k=7.5 GPa value
compares well to the modulus of elasticity of FRB-10 which is 9.3
GPa. The 0.26 correction factor shown in Equation (8) was used
throughout all subsequent simulations.
Model Damping Calculation
[0165] To estimate the damping coefficient, .beta., its value was
increased in the model until the decay in the second mode amplitude
approximately matched the measured response as shown in FIG. 23.
Again, a 10 mm intraoral implant with a 10 mm abutment was used. As
can be seen, the damped model response agrees very well with the
measurements. The value of the damping coefficient found was
.beta.=2.45.times.10.sup.-7 sec in this case. This damping
proportionality constant results in a damping ratio for each mode
as shown in Table 5. The determined value of the damping
coefficient (.beta.=2.45.times.10.sup.-7 sec) is used for all
subsequent model simulations of implants placed in FRB-10.
TABLE-US-00005 TABLE 5 Calculated damping ratio for each mode for
implants placed in FRB-10 Mode Damping Ratio (%) 1 0.1 2 1.5 3 4.8
4 6.3
Effect of Flange
[0166] To test the effect a flange has on implant stability, impact
measurements on a 3.75.times.4 mm flanged extraoral implant with a
10 mm abutment were performed. The flange on the implant was then
removed with a lathe and the tests with the 10 mm abutment were
repeated. The removal of the flange resulted in the average
measured first mode frequency decreasing from
.omega..sub.1=1536.+-.9 Hz to .omega..sub.1=1337.+-.12 Hz,
indicating that the flange was providing extra support to the
implant. The results are shown in FIG. 24 for the 4 mm flangeless
extraoral implant and in FIG. 25 for the 4 mm flanged extraoral
implant.
[0167] If the previously determined K.sub.I and K.sub.T values for
a 10 mm abutment are used with the analytical model the supporting
stiffness (k) for the flangeless 4 mm implant was 7.3-8.1 (7.7)
GPa, which agrees well with the k=6.8-8.4 (7.5) GPa value found for
the 10 mm implant previously. Since k is represented as a stiffness
per unit length the two implants should have similar stiffness
values, as they are supported by the same material. The results of
16 impact measurements with a 4 mm flangeless implant with a 10 mm
abutment compared to the model results with k=7.7 GPa are shown in
FIG. 25A. The predicted acceleration response shows excellent
agreement with the flangeless implant measurement.
[0168] To determine the value of K.sub.F in the model, K.sub.T and
K.sub.I were as determined for a 10 mm abutment and k was set to
7.7 GPa for the flanged implant case. The effective distance that
K.sub.F was applied was taken as half the width of the 2 mm flange
plus the radius of the implant (1.875 mm) to give r=2.875 mm. The
value of K.sub.F in the model was then increased until the model
first mode frequency matched the average measured first mode
frequency of the flanged readings (.omega..sub.1=1536 Hz) which
occurred when K.sub.F=3.65.times.10.sup.7 N/m. A comparison between
the flanged measured results and model results is shown in FIG.
25.
[0169] In FIG. 25 the model results do not match the measured
signals as well as in previous tests. In particular, the predicted
higher frequency component does not agree as well as for the
flangeless implant. One possible explanation is that when the
implant was placed in the FRB-10 some of the epoxy used to secure
the implant ended up under the flange, bonding the flange to the
FRB-10 surface providing not only vertical but horizontal support
for the flange as well (which was not included in the model).
Model Acceleration Response
[0170] The damped model acceleration response shown in FIG. 23 is
in actuality a superposition of four different acceleration
responses, which are shown in FIG. 26. FIG. 26 shows four modal
acceleration components for a 10 mm implant with a 10 mm abutment
and k of 7.5 Gpa. Note that the third and fourth mode responses
have been magnified for clarity. The maximum amplitude of the
second mode response is approximately an order of magnitude smaller
than the maximum amplitude of the first mode response. The maximum
amplitude of the third and fourth modes is approximately three and
four orders of magnitude smaller, respectively. Only the first and
second modes make noticeable contributions to the overall response.
Additionally, FIG. 26 also demonstrates the effects of damping in
the model with the higher modes being damped out more quickly than
the lower ones.
Model Validation
[0171] While the model simulations in previous sections indicate a
very good agreement with the actual acceleration response, this was
for a limited number of specific tests. In order for the model to
be effective it should be able to accurately simulate a broad range
clinical situations. To this end, the fundamental frequency results
from measurements utilizing different implant-abutment parameters
were compared to the model results. Tests with different striking
heights, different length implants and with different abutment
lengths were conducted to evaluate the suitability of the model
under different geometric conditions while holding k constant for
each disk used.
[0172] The analytical model frequency results were calculated using
the previously determined stiffness values. For the FRB disk
containing the 4 mm implant k was 7.7 GPa and for the disk with the
10 mm implant k was 7.5 GPa. These values were held constant for
all the subsequent comparisons. Similarly, for all of the in vitro
results, K.sub.I for each length abutment was as listed in Table 4,
K.sub.T values were as calculated from Equation (8) and K.sub.F,
where appropriate, was 3.65.times.10.sup.7 N/m.
Variations in Striking Height
[0173] One technique used to validate the model was to compare the
model results with experimental results obtained from striking a 10
mm abutment at different heights above the surface of the FRB disk.
Measurements were completed at different striking heights along a
10 mm abutment for the 3.75.times.4 mm flanged extraoral implant
and 4.times.10 mm intraoral implant. The measurements were taken by
striking the top corner of the abutment and then lowering the
handpiece 1, 2, 3, and 4 mm. Five readings were taken at each
height.
[0174] The measured first mode frequencies (.omega..sub.1) are
compared to the predicted model first mode frequencies (p.sub.1) in
FIGS. 27A and 27B. FIG. 27A shows the comparison using a 10 mm
intraoral implant, while FIG. 27B shows the comparison using a 4 mm
extraoral implant. In each graph, five measured frequency values
are averaged and the error bars shown represent two standard
deviations of the repeatability and reproducibility of the
measurement system (.+-.24 Hz). The model results agree very well
with the measurements for both implants and demonstrate the
sensitivity of the impact method to variations in striking height.
To reduce measurement variation due to changes in striking height
measurement protocols should ensure that impacts occur at an easily
identifiable and repeatable position (such as the superior rim of
the abutment).
[0175] To compare the second mode frequency to the model results at
different striking heights, the model predicted acceleration
response for the 10 mm implant case was directly compared to
measurements as shown in FIG. 28. In FIGS. 28A through 28D, the
abutment is struck at the top of the abutment, 2 mm from the top of
the abutment, 3 mm from the top of the abutment, and 4 mm from the
top of the abutment, respectively. The measured second mode
frequency appears to match the model frequency quite well for the
different striking heights, with the exception of FIG. 28B where
the model under-predicts the second mode frequency. The model
predicted amplitude of the second mode frequency appears to be
smaller than the measurements in all but FIG. 28A.
[0176] While the first mode frequency in both the measurements and
model match very well, there is some discrepancy between the model
results and the measurements when comparing the second mode
frequency and amplitude. These differences may be due to
assumptions made to take into account the deformation of the
abutment. At different striking positions along the abutment
K.sub.I may have different values, as the rim of the abutment will
likely be less stiff than the wall of the abutment. There may also
be some errors introduced by the manner the model handles bending.
Modeling bending with a torsional spring in Equation (8) may be too
simplistic to provide a higher level of agreement. Although the
level of agreement in FIG. 28 is not as good as in previous
measurements, the level of agreement is still reasonable
considering the simplifying assumptions made in the model. While
there is some discrepancy between the predicted higher frequency
components in the accelerometer signal and measurements the lower
frequency or "contact time" shows excellent agreement in all cases.
It should be noted that clinically, only impacts at the top are
relevant.
Variation in Abutment Length
[0177] Since different abutment sizes are commonly used with
implants, it is important to compare the model results with
different sized abutments. Both the 3.75.times.4 mm flanged
extraoral implant and the 4.times.10 mm intraoral implant had 4,
5.5, 7 and 10 mm abutments connected with a torque of 20 Ncm. Each
implant was tested five times on the top rim of each abutment. As
in the previous section, model first mode frequencies (p.sub.1)
were compared to the measured fundamental frequencies
(.omega..sub.1). The results of this comparison can be found in
FIG. 29. FIG. 29A shows the comparison for a 10 mm intraoral
implant, while FIG. 29B shows the comparison for a 4 mm extraoral
implant. The measurement results in the figure are the average of
the five readings and the error bars are two standard deviations of
the repeatability and reproducibility of the measurement setup. The
results in FIG. 29 again show good agreement between the model
predicted fundamental frequency and the measured values for the
abutments tested. The agreement between the predicted values and
measurements provides evidence that the model correctly accounts
for the effect different length abutments have on the fundamental
frequency.
[0178] The predicted model acceleration response for the different
abutment sizes with a 10 mm implant are compared to the measured
results in FIG. 30. FIGS. 30A through 30D show the comparison using
a 10 mm abutment, a 7 mm abutment, a 5.5 mm abutment, and a 4 mm
abutment, respectively. The second mode amplitude and frequencies
match the measured values quite well in all of the cases shown.
Analytical Model Simulations
[0179] One use of the developed analytical model is to investigate
expected changes in the first mode resonant frequencies of
Branemark implant-abutment systems due to simulated changes in bone
properties. Three changes in bone structure were investigated with
the analytical model; changes in the supporting bone stiffness,
changes to the damping properties of the bone and marginal bone
height losses around the neck of the implant. The model was then
used to determine if it would be possible to predict the effect a
flange has on implant stability in vivo. The flange stiffness
determined previously represents a flange bonded to the support
surface (such as might occur when the flange was osseointegrated
with the bone surface). Simulations with this flange stiffness will
provide simulations for cases in which the flange is providing a
maximum amount of support.
Simulation of Changes to Supporting Stiffness
[0180] Changes in the supporting stiffness can be modeled by
changing the stiffness of the horizontal and vertical springs (k)
in the analytical model. The simulations of changes in k were done
over a range of implant-abutment geometries, a 4 mm extraoral
implant with 5.5 mm and 7 mm standard abutments and a 10 mm
intraoral implant with 5.5 mm and 7 mm standard abutments. For the
simulations, all impacts occur at the top rim of the abutments. The
stiffness k was varied from 0.75 to 15.0 GPa for each implant with
each different abutment. The variations in k between 0.75-15.0 GPa
represents the range of supporting stiffness used to produce first
mode frequencies equivalent to those measured in vivo in
patients.
[0181] The effects of varying k on the first mode frequency are
shown in FIGS. 31A and 31B. FIG. 31A shows the effect using a 10 mm
intraoral implant, while FIG. 31B shows the effects of using a 4 mm
extraoral flanged implant. For both implants, the effect of
increasing the abutment length from 5.5 to 7 mm lowered the
resonant frequency. For the flanged 4 mm implant two separate
simulations were done for each abutment, one simulation without a
flange and one with the flange value determined from the in vitro
simulations as described previously. The K.sub.F value determined
was for a flange with a thin epoxy layer bonding it to the FRB disk
surface, and was taken as a maximum possible flange contribution.
In FIG. 31B the upper curve for each abutment represents the
maximum flange effect and the lower curve shows the effect without
a flange. In a clinical situation the flange stiffness would
produce an effect between the maximum and minimum curves shown.
From FIG. 31, the 5.5 mm abutment has a slightly greater change in
frequency over the range of k than the 7 mm abutment for both
implants.
[0182] FIG. 31B also shows that the curves without a flange have a
greater frequency range than with a fully integrated flange. The
inclusion of a flange has the effect of reducing the sensitivity of
the resonant frequency to changes in the support stiffness k.
[0183] In FIGS. 31A and 31B, a steeper slope indicates a greater
frequency sensitivity to changes in k. For both implants the curves
start to plateau after approximately 5 GPa. This indicates that as
the supporting bone stiffness (k) continues to increase the
resonant frequency becomes less sensitive to these changes. For
cases in which the supporting bone stiffness is high, the
measurement system may be unable to quantify changes occurring in
the bone properties. However, for values of k in this upper range,
the implant is generally considered well integrated and not in
immediate danger of failing, so the changes which may occur in k
are less important. Fortunately, the impact test is much more
sensitive to changes in supporting bone properties for a poorly
integrated implant that may be in danger of failing.
Simulation of Changes in Damping Properties
[0184] Many studies utilizing the Periotest often erroneously refer
to the device as measuring the damping characteristics of the
interface. To estimate the effects due to changes in damping the
damping coefficient (.beta.=2.45.times.10.sup.-7 sec) used in the
model was doubled, then quadrupled for a 10 mm intraoral implant
with a 10 mm abutment as shown in FIG. 32.
[0185] FIG. 32 demonstrates that as the damping coefficient is
increased the amplitude of the second mode frequency is affected,
however, there is virtually no change in the contact time. Devices
that utilize contact time or resonant frequency measurements (such
as the Periotest and Osstell) are therefore very insensitive to
changes in damping when implant systems are considered.
Simulation of Bone Loss
[0186] One of the mechanisms with which an implant can fail is from
crestal bone loss around the head of the implant. It has been
suggested that, in some cases, implant failure may be the result of
a "positive feedback" loop in which bone loss at the top of the
implant leads to more bone loss and this continues until implant
failure. If implants can be identified as having bone loss early
enough, preventative measures may save the implant. As such, the
ability to measure implant bone loss would be of clinical value. To
this end, the model was used in a number of simulations to help
determine how bone loss may manifest itself in the impact
measurements.
[0187] For the simulations, bone loss starts at the top of the
implant and progresses toward the base. For the bone loss
calculations, two implant-abutment geometries were used, a 4 mm
extraoral implant with a 5.5 mm abutment and a 10 mm intraoral
implant with a 5.5 mm abutment. In the simulations the engagement
length was reduced 5 mm in 0.5 mm increments for the 10 mm implant,
and 2 mm in 0.5 mm increments for the 4 mm implant. This was done
for k values of 1, 5, and 10 GPa.
[0188] The simulations of the changes in the first mode frequency
due to bone loss around a 10 mm intraoral implant is shown in FIG.
33A and for a 4 mm extraoral flanged implant in FIG. 33B. The k=1
GPa curve in FIG. 33A shows substantially a linear relationship
between bone loss and first mode frequency. At higher support
stiffness values the relationship between the amount of bone loss
and first mode frequency is nonlinear and there is a smaller
overall change in frequency corresponding to the bone loss. For k=1
GPa the first mode frequency changes by approximately 800 Hz
(amounting to a change of about 80 Hz per half-millimeter of bone
loss) while for k=10 GPa the change is approximately 500 Hz. The 4
mm implant curves shown in FIG. 33B are substantially linear,
however, there is an initial rapid decrease in stability during the
first 0.5 mm of bone loss. This decrease in the first 0.5 mm was
caused by the removal of the flange stiffness K.sub.F as material
is removed from under it. This is more extreme than what likely
occurs in practice, as the K.sub.F value used was larger than would
be expected clinically. The removal of the flange was less
significant in the 10 GPa case than when k=1 GPa. This is due to
the underlying stiffness k being higher in the 10 GPa case, thus
K.sub.F provides proportionally less stability than it does for the
1 GPa case. This indicates that as the supporting bone becomes
stiffer, the effect of K.sub.F becomes less significant.
[0189] After the initial loss of K.sub.F, the first mode
sensitivity to bone loss shown in FIG. 33B decreases as the
supporting stiffness increases. There is a change of about 100 Hz
per half-millimeter of bone loss for the k of 1 GPa and 75 Hz per
half-millimeter of bone loss for a k of 10 GPa. The 4 mm extraoral
implants have a greater change in frequency per half-millimeter of
bone loss as compared to the longer intraoral implants (100 Hz
compared to 80 Hz in the first mode for a k of 1 GPa). This is not
entirely unexpected, as it indicates that shorter implants are more
sensitive to the loss of bone along their lengths than a longer
implant.
Simulation of Flange Loss
[0190] In the previous section, it was shown that the loss of the
flange reduces the stability of the 4 mm extraoral flanged implant.
It would be useful if the model could predict the effect of a
flange in vivo based on the measured impact accelerometer response.
The flange value used was K.sub.F=3.65.times.10.sup.7 N/m as
determined in the model verification section. As discussed
previously, this K.sub.F value represents a maximum flange
contribution case with the flange fully bonded to the supporting
surface. Clinically, the value of K.sub.F would likely fall between
either no flange support or the maximum K.sub.F value. To this end,
the model acceleration response for a 4 mm extraoral implant with a
10 mm abutment was compared with and without a flange at two
different first mode frequencies as shown in FIG. 34. The two
frequencies were chosen to represent a stable implant measurement
(1500 Hz) and a less stable implant measurement (1300 Hz). FIG. 34A
shows the model predictions for stable implants which have higher
measured first mode frequencies (1500 Hz). There is substantially
no difference between the results with and without a flange,
indicating that for more stable implants, the inclusion of a flange
has a negligible effect on the model output response. However, for
less stable implants (1300 Hz), as shown in FIG. 34B, there is a
noticeable difference between the higher frequency component in the
response for the flange and no-flange signals. This suggests that
by comparing the measured results to the model predictions, it may
be possible to determine the degree to which a flange is
contributing to the overall stiffness of the system for less stable
implants.
Conclusions
[0191] An analytical four-degree of freedom model was developed to
aid in interpreting the response of different implant-abutment
geometries during impact measurements. The model relates the
resonant frequencies of the system to the supporting bone stiffness
which was represented as a stiffness per unit length k (GPa). The
analytical model includes a number of internal stiffness components
to represent local deformations during impact and
bending/flexibility about the implant-abutment joint. However, a
correction factor of 0.26 was applied to the bending/joint
flexibility equation. The correction factor is a likely result of
the combination of the simplifying assumptions made to incorporate
bending into the model, a non-idealized joint, and the complete
absence of shear effects in the current analysis. While the 0.26
value was determined from matching model results to measurements
for one specific geometry, it was held constant throughout all
subsequent simulations on different length implants, different
striking heights, and different abutment lengths. With this one
modification a very high level of agreement with the measurements
was obtained over a variety of geometric conditions.
[0192] Once validated, the model could evaluate the supporting
material properties for implants based on un-filtered (raw)
accelerometer impact measurements. Model estimates of the
supporting stiffness in vitro found the average support stiffness
of FRB-10 modeling material to be 7.5-7.7 GPa which is comparable
to the 9.3 GPa modulus of elasticity.
[0193] Model results were compared to measured in vitro cases over
a range of implant-abutment geometries. The predicted response
showed good agreement with a number of in vitro measurements
demonstrating that the internal stiffness components in the system
could not be ignored and had to be included to accurately reflect
the system dynamics. The model internal stiffness parameters were
determined based on tests on a limited number of Nobel Biocare
implant/abutment systems. As there are presently a large number of
different implant/abutment designs available, these parameters may
have to be evaluated for these different implant systems. The
agreement between the analytical model acceleration response and
the in vitro testing indicated that the high frequency component
found in the accelerometer signal was a second mode of vibration of
the system.
[0194] Model simulations were then used to predict the effect of
changes in the stiffness (k) on the first mode resonant frequency
measurements. The model simulations demonstrated that for support
stiffness values greater than approximately 5 GPa the first mode
frequency becomes less sensitive to changes in the supporting
stiffness. This indicates that due to the stiffness inherent in the
implant/abutment system, there is an upper limit to the support
stiffness that the impact measurement can effectively distinguish.
However, for these values the implant is generally considered
healthy, so the changes which may occur in k are of lesser
importance. Model simulations were then used to show that damping
changes affect the amplitude of the accelerometer signal,
particularly the second mode, while having little influence on the
implant system's resonant frequencies. Current dynamic mechanical
testing methods that measure contact time or resonant frequency
(such as the Osstell and Periotest) are relatively insensitive to
changes in the damping properties. The effects of bone loss from
the top of the implant were modeled. Both the 10 mm intraoral
implant and 4 mm extraoral flanged implant were found to be
sensitive to bone loss. The sensitivity to bone loss decreased for
both implants as the support stiffness increased. The 4 mm
extraoral flanged implant was also shown to be more sensitive to
bone loss than the longer 10 mm intraoral implant.
[0195] Finally, the model was used to predict the effect the flange
has on implant stability and to determine if it would be possible
to use the model as a diagnostic tool in evaluating the effect of
the flange in vivo. For implant systems with higher first mode
frequencies, where the implant is considered healthy, the model was
not able to distinguish between the flange and no-flange condition.
However, if the implant is less stable the model does show
significant differences in the predicted measurement responses
between the flange and no-flange conditions. For these "less stiff"
cases it may be possible to estimate how much stability is being
provided by the flange and how much is due to the supporting
bone.
[0196] The developed analytical model, in conjunction with the
impact measurements, can allow direct estimation of the bone
properties that support implants. Model simulations show the impact
testing technique to be sensitive to bone loss and stiffness
changes that would correspond to poorly integrated implants (ones
that may be in danger of failing). Similarly, for implants with
very stiff support, little useful quantitative data can be obtained
about the bone supporting the implant, as the stiffness of the
other components of the system dominate the response. However, such
implants are generally considered healthy.
[0197] Note that simpler models (e.g. three-degree of freedom
model) can easily be derived from the four-degree of freedom model.
For instance, in the event that the abutment 52 and the implant 51
are part of the same rigid component (i.e. they are not separate
components), then a simpler model could be derived by assuming
K.sub.T to be infinite. In practical implementations, K.sub.T can
be given a very large value instead of an infinite value. Other
derivatives of the four-degree of freedom model are possible.
Section III: Impacting Particulars
Introduction
[0198] Adherence to a strict clinical protocol is used to yield
reproducible results. One of the advantages of the use of an impact
technique--its flexibility--is also a disadvantage in that used
incorrectly it may give inconsistent or spurious results that have
no clinical value. This is believed to be one of the reasons for
the large variations in results reported in the literature. It
appears that one of the major factors causing inconsistent results
is uncontrolled clinical variables.
[0199] The repeatability and reproducibility of the current
measurement scheme when measuring the same implant/abutment system
were discussed earlier with reference to FIG. 13. This figure also
highlights the fact that the results are even more consistent for
an individual test (small error bars for any given column), and
suggests that when the impacting rod is re-aligned even in a
controlled laboratory setting, variability is added to the results
(difference between columns). This highlights the importance of a
strict protocol to maximise the precision of the measurement.
Method of Conducting an Impact Test
[0200] Referring now to FIG. 35, shown is a flowchart of an example
method of conducting an impact test. This method includes steps
carried out by a person, such as a dentist or a clinician, for
using the impact test for example on a patient. Note that this
method can be applied to a plurality of different implant/abutment
systems, or to a single implant/abutment system to assess the
integrity of the implant interface over time.
[0201] At step 35-1, the person impacts an impact body against an
abutment affixed to the implant. According to an embodiment of the
invention, at step 35-2 the person ensures that the impact body
impacts against a superior rim of the abutment. If the person
always impacts the impact body against the superior rim of the
abutment, then there is consistency in using the impact test. Note
that the superior rim of the abutment is typically easy to identify
and therefore the person can achieve success in consistently
impacting against this portion. Alternatively, the person ensures
that the impact body impacts against another portion of the
abutment, provided that the person consistently impacts against the
same portion.
[0202] For any given implant system, impact test readings will by
meaningful (and comparable) if that implant system is consistently
struck at the same spot. That way impact test results over time can
be compared to see how the interface is changing/progressing. Note
that the "same spot" does not need to be the same for different
implant systems (but when abutments are involved the superior rim
is a logical choice). For different artificial teeth systems, the
consistent spot may be slightly different among these systems,
depending on the details of each system. However, within a given
implant system, in order to compare results over time the same
consistent spot should be used.
[0203] In some implementations, as indicated at step 35-3, the
person ensures that the impact body impacts against the abutment at
an angle between about 1.degree. and about 5.degree. above a plane
perpendicular to an axis of said abutment. In some implementations,
as indicated at step 35-4, the person ensures that the impact body
is initially positioned between about 0.5 mm and about 2.5 mm from
the abutment.
[0204] Note that the method described above assumes that each
implant is at least partially embedded in a medium and has an
abutment connected thereto. It is to be understood that the "each
implant . . . having an abutment connected thereto" does not
necessarily mean that the abutment and the implant are formed of
separate members. In some implementations, the abutment and the
implant are formed of a same continuous member. In this manner,
although the abutment and the implant are referred to separately,
they are still part of the same continuous member. In other
implementations, the abutment and the implant are formed of
separate members.
[0205] In some implementations, when the abutment and the implant
are formed of separate members, they are threaded attached. There
are many ways that the abutment and the implant can be threaded
attached. In some implementations, they are threadedly attached
with a torque applied to the abutment that exceeds about 10 Ncm.
Other implementations are possible.
[0206] The present invention "ensures" that the impact test is
performed in a manner that can yield accurate results. Previous
approaches do not ensure that the impact body impacts against a
superior rim of the abutment. Rather, they typically provide no
guideline, which can result in inaccurate results. The present
invention includes specific guidelines for adherence in order to
achieve acceptable results. These specific guidelines come from
results of experimentation, details of which are provided
below.
Experimentation
[0207] An experimental apparatus was used to evaluate several
clinical variables that potentially could affect the readings.
These variables include: [0208] handpiece distance from abutment,
[0209] abutment torque, [0210] striking height (position along the
abutment where contact is made), and [0211] angulation of
handpiece. To evaluate the effect of these variables, one variable
was changed while attempting to hold all other variables constant.
Measurements were done by striking the top rim of the abutment in
each of these cases. Handpiece Distance from Abutment
[0212] Referring now to FIG. 36, shown is a chart depicting natural
frequency as a function of the distance of the Periotest handpiece
from the abutment. For these readings, measurements were taken at
distances of 0.5, 1.0, 1.5, 2.0, and 2.5 mm from the 4 mm
implant/5.5 mm abutment system. Five readings were taken at each of
these distances then the handpiece was re-aligned and the readings
were repeated for a total of three separate trials.
[0213] The mean value for the 0.5 mm distance was 2121.+-.25 Hz
while the reading at 2.5 mm was 2116.+-.36 Hz. It should be noted
that for the 2.5 mm readings the Periotest did not produce a PTV
value, however a resonant frequency was obtained from the moving
average filtered acceleration data. The Periotest instructions
recommend that the handpiece be held a distance of 0.5 to 2.0 mm
from the object being measured. The distance of the handpiece from
the abutment was shown to have little influence on the resonant
frequency. As long as the initial distance from the handpiece tip
to the abutment tip was between 0.5 and 2.5 mm there were
practically no differences noted.
Abutment Torque
[0214] Referring now to FIG. 37, shown is a chart depicting natural
frequency as a function of abutment torque. For these readings, a
5.5 mm abutment was torqued to the 4 mm implant system at 5, 10,
15, 20, and 25 Ncm. Five consecutive measurements were done at each
of these values. Torque values were measured with a TorsionMaster
Testing System (MTS Systems Corp, Eden Prairie, Minn., USA). Three
separate trials at each of the torque values were conducted.
[0215] The 5 Ncm torque (which was noticeably loose) had the lowest
resonant frequency reading of 1615 Hz and the largest standard
deviation of 248 Hz. The torque applied when mounting a standard
5.5 mm abutment has little effect on the resonant frequency for
torques above 10 Ncm. The torque applied to the abutment when
mounted to the fixture had a large effect on the resonant frequency
until the torque exceeded approximately 10 Ncm. For torques below
this value, which are rarely encountered clinically, the reduced
stiffness of the joint caused a large reduction in the resonant
frequency of the system. This effect has been reported previously
based on PTV values. For torques greater than 10 Ncm, the resonant
frequency remained substantially unchanged. The most consistent
results (lowest standard deviation in the readings) occurred at a
torque of 20 Ncm. It should be noted that this threshold torque
(over which no change occurred) was for an Entific system implant
and standard 5.5 mm abutment. The effect of varying torque on other
implant/abutment systems could vary depending on the details of the
thread surfaces and length of the abutment (length of threaded
screw).
Vertical Striking Height
[0216] Referring now to FIG. 38, shown is a chart depicting natural
frequency as a function of striking height. This shows a very
significant effect that striking height has on the resonant
frequency. While there was very little change in the frequencies
when the handpiece was moved up to 1.5 mm from its initial
position, there was a noticeable difference between the 1.5 mm and
2 mm positions, and beyond.
[0217] For these readings, a 10 mm abutment replaced the 5.5 mm
abutment used in previous measurements, since a 10 mm abutment
allowed for a greater variation of the striking height.
Measurements were taken striking the top of the abutment and then
lowering the handpiece distances of 0.5, 1, 1.5, 2, 3, 4, 5, and 6
mm. Five readings were taken at each height. The handpiece was
re-aligned and the readings were repeated for three separate
trials.
[0218] The position at which the impacting rod strikes the abutment
(striking height) can have a very pronounced effect on the resonant
frequency. FIG. 38 shows that a 3 mm variation results in a change
to the resonant frequency of 194 Hz (13%). However, FIG. 38 also
shows that there was effectively no change in the resonant
frequency when the rod is moved up to 1.5-2.0 mm from its initial
position. This is due to the fact that the impacting rod is 2 mm in
diameter, and since it was hitting the rim of the abutment in its
original position, it could move up to 2 mm (depending on its exact
initial position) before it started striking a point below the top
rim of the abutment. As long as some portion of the Periotest rod
struck the rim of the abutment little variation in the results
occurred.
[0219] As the effect of striking height on the resonant frequency
is considerable, it is recommended that the impacting rod always
strike the superior rim of the abutment, a point that is clinically
easy to identify and a point that allows a .+-.1 mm variation when
centred, without significantly changing the results.
Angulation of Handpiece
[0220] Referring now to FIG. 39, shown is a chart depicting natural
frequency as a function of handpiece angulation. For these
readings, five consecutive readings were done at 0.degree.,
1.degree., 2.degree., 3.degree., 4.degree., 5.degree., 10.degree.,
15.degree., and 20.degree. such that 0.degree. corresponds to when
the handpiece is perpendicular to the abutment. Measurements were
done on the 4 mm implant with the 5.5 mm abutment. This process was
then repeated for three separate trials.
[0221] A handpiece angulation from 0.degree.-20.degree. caused the
resonant frequency of the system to change from 2178.+-.19 Hz to
2236.+-.10 Hz. The results at 0.degree. are noticeably different
from the 1.degree. readings, while the results are more consistent
between 1.degree. and 5.degree.. Note that the Periotest
instructions recommend an angulation of +20.quadrature. from the
horizontal. This range is significantly greater than 1.degree. to
5.degree..
[0222] When kept within a 1.degree. to 5.degree. range, no
substantial differences were evident. There was, however, a
noticeable difference between the 0.degree. and 1.degree.
measurements. This difference is likely due to the fact that when
the handpiece is nominally perpendicular to the striking surface it
is not certain which part of the 2 mm diameter rod is striking the
abutment. If the lower edge of the rod strikes the abutment this
results in a higher frequency reading than if the top part of the
rod strikes (effectively there is a change in striking height as
the rim of the abutment is not being contacted). To eliminate this,
a slight angulation of the handpiece is advisable. As angulation
increases to 10.degree. and beyond there is a trend of increasing
resonant frequency.
[0223] Thus, the inconsistent and insensitive results reported when
using the Periotest for measuring implants may result from both the
techniques used to analyse the accelerometer signal and from
clinical variations that occur during measurements. Utilising a
moving average filtered signal and a stricter measurement protocol,
it is believed that the impact technique can provide a reliable and
sensitive diagnostic means to monitor implant stability.
Section IV: Patient Study
Introduction
[0224] In vivo testing was done in conjunction with the
Craniofacial Osseointegration and Maxillofacial Prosthetic
Rehabilitation Unit (COMPRU), located at the Misericordia Community
Hospital, Edmonton, Alberta, Canada. All testing was approved by
the University of Alberta Health Research Ethics Board and patients
signed an informed consent form prior to taking part in the
study.
[0225] The patient study group included 12 patients (8 males and 4
females) with a mean patient age at time of implant placement of 53
years (range 27-75 years). Patients enrolled in the study were
treated with bone anchored hearing aid (Baha) implants which were
left to heal for 3 months before the patients received their
hearing processors. To have been considered for the study the
patients: [0226] had to be 18 years of age or older, [0227] had to
meet audiological criteria for selection into the Baha program,
[0228] had to be able to maintain a skin penetrating abutment,
[0229] could not have any condition that could jeopardise
osseointegration (e.g. malignancy in the temporal region, radiation
therapy of the temporal region, undergoing chemotherapy), and
[0230] had to be able to understand and read English. Following a
one-stage procedure, 12 flanged extraoral implants (3.75 mm, SEC
002-0, Entific Medical Systems, Toronto, Ontario, Canada) were
placed (one per patient). The implants for 11 of the patients were
4 mm in length while one patient received a 3 mm implant. Implants
were installed on either the right or left side, based on the
audiological recommendation.
Clinical Protocol
[0231] An in vivo protocol was developed prior to patient
measurements based on previously completed in vitro measurements
(submitted for publication, Swain, R. et al., International Journal
of Oral & Maxillofacial Implants, 2006): [0232] The handpiece
would be aligned so that the impacting rod would strike the
superior rim of the abutment. [0233] The handpiece should be held
with a slight angulation (1 to 5 degrees) from a line perpendicular
to the longitudinal abutment axis, [0234] To ensure the
measurements were taken in a consistent azimuthal direction, the
handpiece was oriented parallel to the longitudinal axis of the
patient (i.e. handpiece pointed towards the patient's feet when
lying flat).
Use of Calibration Block
[0235] Referring now to FIG. 40, shown is a photograph of a
calibration block used during in vivo measurements. To ensure that
the in vivo measurement values were as precise as possible, the
calibration block was used. The block includes four aluminum posts
with lengths of 4, 6, 8 and 10 mm threaded 4 mm into a rectangular
piece of Photoelastic FRB-10 plastic (Measurements Group Inc,
Raleigh N.C., USA). Epoxy was applied to the post threads during
installation to provide a uniform interface and to prevent any
loosening of the posts over time. The FRB-10 block was then mounted
in a stainless steel base. For each aluminium post, there is a
known natural frequency or other system property for the impact
test.
[0236] Measurements were taken by the clinician on each of the four
posts prior to the patient measurements as shown in FIG. 40. The
clinician was instructed to align the impacting rod so that it
would strike the superior rim of the post and with an angulation
between 1 to 5 degrees. The calibration measurements included at
least one impact measurement per post, with the measurement values
being compared to the values engraved on the calibration block. The
calibration block served two important purposes, it provided a
method for evaluating if any longitudinal changes occurred in
handpiece output, and it focused the operator on the proper
measuring technique prior to the patient measurements.
In Vivo Measurements
[0237] In vivo measurements involved impact measurements with
different abutment geometries at one patient visit as well as
longitudinal patient readings over the course of one year after
initial installation. To reduce any measurement inter-operator
variability only one clinician conducted the measurements at all
but implant installation. Due to the scheduling of the surgeries it
was not always possible for the same clinician to be present during
implant installation. In these cases, either another experienced
clinician or the surgeon performed the calibration and
measurements.
Measurements on Different Length Abutments
[0238] Two different length abutments were utilised in the study to
test the consistency of the proposed measurement method and
analytical model results for different implant-abutment geometries.
Measurements were completed using standard 7 and 5.5 mm abutments
(Nobel Biocare, Toronto, Ontario, Canada) for 10 of the twelve
patients at the one year patient visit (the multiple abutment
measurement was missed for 2 patients). Three impact measurements
were completed on each of the abutments, which were affixed to the
implant with a torque of 20 Ncm. After 3 impact measurements on the
5.5 mm abutment the 7 mm was connected to the implant and three
additional measurements were taken.
Longitudinal Impact Measurements
[0239] The in vivo longitudinal study involved three impact
measurements for each patient at implant installation and then at
1, 2, 3, 6 and 12 month scheduled patient visits. The measurements
were completed during the patient's regularly scheduled visits to
minimise additional time commitments. The impact measurements were
taken using 5.5 mm standard abutments (Nobel Biocare, Toronto,
Ontario, Canada) coupled to the implants with a torque of 20
Ncm.
Impact Accelerometer Signal Analysis
[0240] The impact signals utilised were from the Periotest
handpiece, which had been modified to permit improved signal
processing to be used with the accelerometer signal. Each separate
impact measurement consisted of a series of 16 impacts (therefore 3
measurements would consist of 48 total impact events). The
accelerometer signals were collected with an Instrunet
analog/digital model 100 sampling system with a sampling rate of
167 kHz connected to a Toshiba Satellite A10 laptop computer.
[0241] The impact signals were used in conjunction with an
analytical model to determine the interface stiffness and damping
properties in vivo. The interface stiffness value, k, is calculated
for each measurement and is reported in units of GPa. Damping
properties are represented as a damping ratio. To examine the
support an implant flange provides in vivo, measurement results
(which included the flange) were compared to analytical model
results with and without a flange for each of the patients.
[0242] The in vivo impact measurements collected in this study are
interpreted using an analytical model, which provides a
quantitative measure of the bone/implant interface stiffness and
damping. In addition, the analytical model provides a means of
evaluating the support the implant flange provides.
Impact Signal Analysis with and without a Flange
[0243] While the theoretical impact response with and without a
flange was compared to the measured impact response for all
patients, a representative example of this comparison for 16
measured impact responses (as each measurement consists of 16
impacts) is shown in FIGS. 41A and 41B. FIG. 41A shows the
comparison with a flange, while FIG. 41B shows the comparison
without a flange. If the number of peaks in the measured impact
signal are compared to the simulations with a flange (FIG. 41A) and
without (FIG. 41B), the predicted response without a flange can be
seen to more closely resemble the measurements.
[0244] The interface stiffness, k, was calculated with and without
a flange providing support for the measurements on the different
abutment geometries. The percent differences between the stiffness
values calculated for the 5.5 and 7 mm abutments with and without a
flange for 10 of the patients are shown in FIG. 42. Note that
multiple abutment measurements were missed on patients 1 and 9 and
therefore data for these patients has been omitted. The percent
difference is defined as the difference between the interface
stiffness for the 7 and 5.5 mm abutments divided by the interface
stiffness determined using the 5.5 mm abutment. As shown the model
estimations with a flange tended to have larger differences in
interface stiffness for the different abutment geometries. In
addition, the differences for the flanged case are generally biased
while the no flange results yield both positive and negative
differences.
[0245] The calculated interface stiffness without the effects of a
flange for the two different length abutments for each patient are
shown in Table 6. The patient number and gender is included in the
first column. The interface stiffness is shown in the table as a
range, with the stiffness value corresponding to the average
measurement shown in parenthesis after the range. The percent
difference between the interface stiffness for the two abutments is
shown in the last column.
TABLE-US-00006 TABLE 6 In Vivo interface stiffness values for a 5.5
mm and 7 mm abutment on the same patient 5.5 mm Abutment 7.0 mm
Abutment Patient k (GPa) k (GPa) % Difference 1 (F) -- -- -- 2 (F)
3.7-4.0 (3.8) 4.1-4.5 (4.3) 13% 3 (M) 2.1-2.5 (2.3) 2.2-2.7 (2.4)
4% 4 (M) 4.5-5.0 (4.7) 3.5-5.6 (4.4) -8% 5 (M) 9.2-56.8 (16.1)
12.4-17.1 (14.4) -11% 6 (M) 2.1-3.0 (2.5) 2.5-3.5 (2.9) 19% 7 (M)
8.8-13.1 (10.6) 5.4-14.8 (8.3) -21% 8 (F) 6.3-7.1 (6.7) 5.7-6.5
(6.1) -8% 9 (M) -- -- -- 10 (M) 6.0-8.9 (7.2) 6.2-7.1 (6.6) -8% 11
(F).sup. 5.5-6.1 (5.8) 5.1-5.9 (5.5) -6% 12 (M) 5.7-8.9 (7.0)
7.4-8.1 (7.7) 10%
The theoretical impact response is compared to the measured impact
responses for a representative patient (Patient 4) in FIGS. 43A and
43B. FIG. 43A shows the comparison using a 5.5 mm abutment, while
FIG. 43B shows the comparison using a 7 mm abutment. The predicted
impact response can be seen to match the measured responses quite
well for both abutment geometries.
Longitudinal Changes in Interface Stiffness
[0246] As shown in FIG. 44, the mean interface stiffness for all
patients at time of implant placement was 5.2 GPa with a similar
measurement of 5.5 GPa after one month. The mean stiffness
increased to 7.3 GPa between the one month and two month
measurements before stabilising for the remaining measurements. No
implants failed during the course of the study. Individual patients
showed distinctly different patterns from the mean as demonstrated
by the results for Patients 1 and 5 (see FIG. 44). While the
initial interface stiffness for the two patients are similar, the
stiffness decreased in the first month for Patient 1 while it
increased in the first month for Patient 5. The stiffness then
decreases at the second month measurement for Patient 5 while
increasing in the second month for Patient 1. Both patients see an
increase in interface stiffness between 3-6 months and end at
significantly different stiffness values at 12 months (16.1 GPa for
Patient 5 compared to 6.7 GPa for Patient 1).
[0247] The longitudinal interface stiffness estimated by the model
for all 12 patients are shown in Table 7. As the analytical model
takes into account changes in system geometry, the stiffness values
shown for Patient 2 can be directly compared to the other patients
although the implant length was different (i.e. 3 mm implant as
compared to 4 mm implant for the other patients).
TABLE-US-00007 TABLE 7 Interface stiffness values (GPa) based on
impact measurements at installation, 1 month and 2 months for 12
patients fitted with Baha implants Patient Installation 1 Month 2
Month 3 Month 6 Month 12 Month 1 (F) 3.1-3.3 (3.2) 2.0-2.2 (2.1)
3.3-3.7 (3.5) 2.7-3.3 (3.0) 4.2-8.4 (5.8) 6.5-6.9 (6.7) 2*(F)
4.8-5.2 (5.0) 1.8-1.9 (1.9) 3.1-3.9 (3.5) 3.2-3.4 (3.3) 4.0-4.1
(4.1) 3.7-4.0 (3.8) 3 (M) 1.5-2.2 (1.8) 2.0-2.2 (2.1) 2.7-3.2 (2.9)
3.8-4.6 (4.2) 3.7-3.9 (3.8) 2.1-2.5 (2.3) 4 (M) 3.1-5.7 (4.1)
4.5-5.2 (4.8) 4.9-5.6 (5.3) 3.0-3.7 (3.3) 6.5-6.6 (6.5) 4.5-5.0
(4.7) 5 (M) 3.6-4.4 (4.0) 8.1-9.2 (8.6) 4.5-4.9 (4.7) 5.2-6.1 (5.7)
12.7-18.8 (15.2) 9.2-56.8 (16.1) 6 (M) 7.6-11.1 (9.1) 1.6-1.8 (1.7)
2.1-2.4 (2.2) 1.5-1.6 (1.5) 3.1-3.4 (3.3) 2.1-3.0 (2.5) 7 (M)
6.5-11.8 (8.6) 10.8-14.1 (12.3) 10.0-13.7 (11.6) 28.6-29.1 (28.8)
12.7-20.0 (15.6) 8.8-13.1 (10.6) 8 (F) 2.5-3.2 (2.8) 2.3-2.7 (2.5)
2.9-3.1 (3.0) 3.6-3.7 (3.7) 4.2-4.4 (4.3) 6.3-7.1 (6.7) 9 (M)
8.7-11.2 (9.9) 4.6-13.1 (7.2) 30.6-38.9 (34.3) 10.8-20.1 (14.3)
6.6-7.5 (7.1) 18.5-25.9 (21.6) 10 (M) 6.6-9.4 (7.8) 11.8-12.4
(12.1) 6.5-6.7 (6.6) 6.3-11.8 (8.4) 11.1-17.8 (13.8) 6.0-8.9 (7.2)
11 (F).sup. 2.6-4.0 (3.2) 3.6-4.6 (4.1) 3.3-3.5 (3.4) 4.2-4.3 (4.3)
7.3-8.2 (7.7) 5.5-6.1 (5.8) 12 (M) 3.4-3.6 (3.5) 5.6-7.6 (6.5)
6.1-6.9 (6.5) 6.9-7.2 (7.1) 6.0-8.6 (7.1) 5.7-8.9 (7.0) *Patient
with a 3 mm implant (all other patients have 4 mm implants)
[0248] Examples of the model predicted impact responses compared to
the measured impacts over the one year time period for Patient 1
are shown in FIGS. 45A to 45F. Results for Patient 5 are shown in
FIGS. 46A to 46F.
Longitudinal Changes in Interface Damping
[0249] In addition to the stiffness properties the damping of the
supporting bone was estimated for each patient at each scheduled
visit by utilising the analytical model to interpret the impact
responses. Across all Baha patient measurements the amount of
damping present in the supporting bone was found to vary
longitudinally, however, the amount of change and overall magnitude
of the damping was very low, with the damping ratio for the first
mode ranging between 0.04-0.43%.
Discussion
[0250] While in vitro measurements with a flange bonded to the
supporting materials surface provided support for the implant the
results displayed in FIGS. 41 and 42 indicate that the model
simulations including flange support do not provide as consistent
interface stiffness results as those which assume the flange offers
no support. When model results were compared to measurements across
all the patients (as was done for one patient shown in FIG. 41) the
results without the flange tended to be in much better agreement
with the measurements. These results are reinforced in FIG. 42
where the differences in the model estimations for the same implant
with different abutment geometries tended to be larger when a
flange stiffness was included in the simulations. Further to this,
the differences for the simulations without a flange oscillate
between positive and negative values while the differences with a
flange appear to have a bias. This bias appears systematic when the
flange is assumed to offer stability to the implant and indicates
that the flange may not be providing significant support in vivo.
Since the underlying interface stiffness doesn't change when the
different abutments are placed on the implant the differences
plotted ideally should be zero. While the results without a flange
in FIG. 42 match the ideal case better than the results including
the effects of a flange, some differences do exist. From the data
shown in Table 6, the largest percent difference was 21% across the
patients and the smallest was 4%. For all but one patient (Patient
2) the difference in the estimated interface stiffness values for
the different abutment geometries could be explained by the
measurement variation (the range of stiffness values for the two
abutment geometries overlap). This indicates that, overall, the
analytical model provides an effective means of removing apparent
changes in measurements due to changes in geometry.
[0251] As the results from FIGS. 41 and 42 indicate that the flange
does not appear to provide significant support to the implant in
patients, the following discussion will concentrate on model
results without any flange contribution to the implant stability.
The validity of this assumption is further demonstrated when the
predicted impact response is compared directly to measurements such
as those shown in FIG. 43. The magnitudes and frequency components
of the signals are relatively well predicted, providing evidence
that the interface stiffness values estimated from the model are
realistic.
[0252] The longitudinal results shown in FIG. 44 indicate that
while an average interface stiffness can be determined across all
the patients at each time interval, the bone response due to
implant placement can vary significantly between different
individuals. Considering specifically Patients 1 and 5, while both
start with similar interface stiffness, and while the average
stiffness across all the patients only shows a 5% increase from
installation to one month (5.2-5.5 GPa), the interface stiffness
for Patient 1 decreased 34% and the interface stiffness increased
for Patient 5 by 118%. The changes in interface stiffness for all
the patients is summarised in Table 7. The difference between the
one month response between patients may be due to differences in
individual healing rates and the corresponding rate of bone
modelling/remodelling at the implant interface. Additionally the
longitudinal implant stability may be sensitive to the implant
installation procedure. Slight differences in the drilling and
tapping procedures may change how the bone responds to implant
placement (this may be especially important in the short term).
[0253] Between the one month and two month measurements the average
patient interface stiffness shown in FIG. 44 shows the largest
increase (33%) then appears to stabilise for the remaining
measurements. This contrasts sharply with the individual patient
results shown, where the largest change in interface stiffness
between measurements occurred between the three and six month
measurements (with Patients 1 and 5 having a 97% and a 169%
increases in this time interval). Seven of the twelve patients
tested had more than a 20% increase in the interface stiffness
between the three and six month patient measurements. The remaining
patients either had little change or more than a 40% decrease in
interface stiffness. The magnitude of interface stiffness changes
indicates that for many patients the bone-implant interface may
still be undergoing significant physiological changes between the
three and six month time interval. Increases in stability may be
the result of increased mineralization of new bone and increased
direct bone contact at the interface. Changes occurring at the
interface during this time period are further complicated by
implant loading. Patients received their processors at the three
month visit. In addition to any changes already in progress, the
stresses caused from the load applied to the implant may have
triggered an adaptive response in the bone around the implant.
[0254] Between the six and twelve month patient measurements there
was a 16% and 6% increase in the bone-implant interface stiffness
for Patients 1 and 5. The difference between six and twelve month
stiffness for these two patients is considerably less than that
found between the three and six month values. The difference
between the installation and six month interval measurements is
greater than the difference between the six month and twelve month
values for ten of the twelve patients tested (as shown in Table 7).
This may indicate that for these patients the majority of the
stiffness changes at the implant interface occurred within the
first six months. This falls in the 4-12 month interval cited by
Roberts in which secondary mineralization of new bone and increased
direct bone contact at the interface occurs and the remodelling of
the non-vital interface and supporting bone is completed.
[0255] When the predicted model impact response is compared
directly to patient measurements as in FIGS. 45 and 46, the
response appears similar to the measurements. The longitudinal
changes in the impact signal reflects changes in the bone
properties occurring at the implant interface. Agreement between
the predicted impact results and measurements demonstrates the
analytical model's ability to accurately evaluate the interface
properties. While overall agreement between measurements and
predicted responses is quite good, the patient measurements
occurring at the twelve month intervals tended to have better
agreement with model predictions than earlier measurements. This is
likely due to the assumption in the model that the interface
stiffness is uniform along its length. At implant placement the
interface may differ along the length of the implant depending on
the gaps between the implant threads and the surrounding bone. The
greater levels of agreement at later stages seems to indicate that
the interface becomes more uniform over time.
[0256] Although the total number of patients included in the study
is not large, there are some trends in the interface stiffness data
that are noteworthy. The average stiffness at the twelve month
Measurement for the male patients was 9.0 GPa and 5.8 GPa for the
females, with the top five twelve month stiffness values belonging
to male patients. These interface stiffness values compare well to
the Young's modulus of 13.4 GPa for cortical bone and 7.7 GPa for
trabecular bone used in finite element simulations of the human
skull. Overall, 67% of the patients had their lowest interface
stiffness within the first month. Five of the patients had their
lowest interface stiffness value at implant installation, with
another three at the one month mark. By the third month, all but
one patient had recorded their lowest interface stiffness value.
From the stiffness values determined, the initial three month
healing period appears to be when the implants are least stable. It
has been suggested that the woven bone lattice that forms at the
implant interface occurs within the first 0.5 months and that the
woven bone cavities then fill with high quality lamellae gaining
strength for load bearing within the first 0.5-1.5 months. The
lower interface stiffness values during this time frame may
correspond to the less stiff woven bone lattice and increases in
stiffness after this point indicating the placement of the high
quality lamellae.
[0257] Based on the tests completed, the current practice of
processor connection and implant loading after three months appears
reasonable. Loading implants during the period of initial
instability may have negative consequences. There is a transition
from primary mechanical stability (stability of old bone) to
biologic stability (stability of newly formed bone). During this
transition, there is a period of healing in which the initial
mechanical stability has decreased but the formation of new bone
has not yet occurred to the level suitable to maintain implant
stability. It has been suggested that, at this point, a loaded
implant would be at greatest risk of relative motion and would be
(at least theoretically) most susceptible to failure of
osseointegration.
[0258] Along with changes in the stiffness, it is believed that the
damping properties of the bone changes as the implant
osseointegrates. Some studies completed with the Periotest refer to
the device as measuring the damping characteristics of the
interface. While damping appears to be present in the measurements,
the largest damping ratio across all patients during the testing
was 0.43% (with damping ratios ranging from 0.04-0.43%). Damping
ratio measurements below 0.43% indicate that there is very little
damping present in bone supporting the Baha implants. There is such
a small amount of damping present that if the damping is neglected
entirely in the model it would have a negligible effect on the
interface stiffness results presented. The low damping ratios
calculated emphasise that the longitudinal changes in the measured
impact response of the in vivo implants tested are caused primarily
from changes in the interface stiffness and not from changes in the
damping properties of the supporting bone.
Conclusions
[0259] The in vivo tests utilising the impact test and the
analytical model provided longitudinal interface stiffness and
damping values for twelve patients fitted with Baha implants. In
vivo testing with two different abutment geometries demonstrated
that the impact technique and analytical model can account for
changes in implant system geometry. Model simulations with and
without a flange indicated that for the patients in the in vivo
study, the implant flange does not appear to significantly
contribute to the implant stability.
[0260] Longitudinal model results show good overall agreement with
the measured impact responses for the patients and provide a direct
measure of the bone-implant interface stiffness and damping
properties. The changes in interface stiffness values
longitudinally varied significantly between patients, indicating
that the bone response to implant placement is highly
individualistic. Further research could be completed to investigate
some of the possible causes for this variation. While the
longitudinal changes in the supporting stiffness varied
significantly between the patients, the male patients tended to
have higher interface stiffness. The average bone-implant interface
stiffness determined at the twelve month measurement was 9.0 GPa
for the male patients and 5.8 GPa for the females. Additionally,
the initial three month period appears to be when the implants have
the lowest interface stiffness. The minimum interface stiffness
values for 11 of the 12 patients occurred during this interval.
[0261] The interface damping properties were determined to be quite
low, with the highest estimate being 0.43%. While the damping ratio
for healthy Baha implants placed in the mastoid appears quite low,
a failing implant may have considerably different damping
properties particularly if scar tissue develops at the interface.
For this reason, further study on the damping ratio in failing
implants would be useful.
[0262] While the above discussion of the preferred embodiment of
the invention was made in the context of tests conducted using a
Periotest device, it is to be understood that the invention may be
utilised with other impact-type implant integrity testing devices,
as will be understood by persons skilled in the art. For example,
the impact rod of the Periotest device may be replaced by other
impact bodies such as bars or hammers. The means of accelerating
the impact rod towards the implant may use electromagnets, springs,
or other means.
[0263] The method of conducting the impact test is described as
using a Periotest device on an abutment threadedly attached to an
implant. It is to be understood that the test can be conducted on
an abutment which is attached to the implant by other means, by
being integral with the implant for example, as would be the case
for natural dentition.
[0264] Numerous modifications and variations of the present
invention are possible in light of the above teachings. It is
therefore to be understood that within the scope of the appended
claims, the invention may be practised otherwise than as
specifically described herein.
* * * * *