1. Lasers in Surgery and Medicine
Automated 3D Bone Ablation With 1,070 nm Ytterbium-Doped
Fiber Laser Enabled by Inline Coherent Imaging
Chenman Yin, MASc, Sacha W. Ruzzante, and James M. Fraser, PhD
Ã
Department of Physics, Engineering Physics and Astronomy, Queen’s University, Kingston, Ontario, Canada K7L 3N6
Background and Objective: Laser osteotomy bears
well-identified advantages over conventional techni-
ques. However, lack of depth control and collateral
thermal damage are barriers to wide clinical implemen-
tation. Flexible fiber delivery and economical benefits of
ytterbium-doped fiber lasers make them desirable for
laser osteotomy. In this work, we demonstrate auto-
mated bone ablation with a 1,070 nm industrial-scale
fiber laser to create 3D target structures with minimal
thermal side-effects.
Materials and Methods: Fresh and dry ex vivo cortical
bone samples are ablated using 50–100 ms laser pulses of
15–30 mJ. In situ inline coherent imaging monitors
ablation dynamics with micron precision and on microsec-
ond timescales. Ablation depth is extracted by on-the-fly
processing of ICI data, enabling feedback control of depth
(via laser pulse number). Final ablated morphology,
measured by an ex situ stylus profiler, is compared to the
target shape. Histological examination is performed to
quantify the thermal side-effects of laser ablation.
Results: Percussion drilled hole depth is highly variable
for fixed laser parameters (880 Æ 151 mm on fresh bone and
1038 Æ 148 mm on dry bone) due to nondeterministic
ablation. ICI-enabled depth control is implemented to
achieve precise ablation of complex 3D features. The RMS
deviation between target and ablated morphology is
12.6 mm. The heat-affected zone is found to be 5–10 mm
on fresh and dry bone.
Conclusions: An ytterbium-doped fiber laser is utilized for
cortical bone ablation with limited thermal side-effects. In
situ real-time ICI measurement enables characterization of
bone ablation dynamics. Furthermore, ICI closed-loop
feedback realizes depth-controlled ablation on heteroge-
neous bone. This proof-of-principle study shows great
promise for ICI-guided laser osteotomy. Lasers Surg. Med.
ß 2015 Wiley Periodicals, Inc.
Key words: 3D morphology; closed-loop feedback; corti-
cal bone; in situ monitoring; histology; laser surgery;
osteotomy
INTRODUCTION
Laser osteotomy bears a number of advantages over
conventional surgical techniques that employ mechanical
tools like rotating drills or oscillating saw blades. Due to
their high degree of spatial coherence, laser beams provide
excellent transverse precision since they can be focused to
micron-scale spot sizes, much smaller than the dimensions
of mechanical cutting pieces. Mechanical tools, which rely
on direct contact with bone, can cause profound vibrations
that not only cause discomfort to the patient but also
induce micro fractures in adjacent tissues [1]. In addition,
deposition of metal shavings is unavoidable due to the
large friction and shear stress between bone surface and
the cutting tool. Previous studies have shown that metal
shavings hinder the healing process and distort post-
operative investigations [2]. In comparison, the non-
contact nature of laser osteotomy overcomes the risks
associated with vibrations and metal shavings. It also
permits flexible choice of cutting path without the
limitations posed by the size and geometry of a mechanical
instrument. Despite all the advantages that laser osteot-
omy can offer, it is still not commonly implemented in
clinical settings. Two main barriers to its use in surgery
are: (1) difficulty in precisely controlling the depth of
incision [3,4] and (2) thermal damage of surrounding tissue
by heat diffusion [5,6]. A third consideration is the cost,
size, and complexity of current laser sources and beam
delivery techniques. This work demonstrates real-time
depth control with automatic feedback to ablate predefined
3D incisions with an industrial-scale fiber laser (ytter-
bium-doped, wavelength 1,070 nm). Histological examina-
tion of laser ablated features shows that with careful laser
parameter choice, surrounding tissue suffers minimal
thermal side-effects.
Bone is known to have complicated structures with
considerable inhomogeneity and porosity, making accu-
rate depth control assuming invariant ablation rates
impractical. Typical bone consists of a dense outer layer
called cortical (or compact) bone, which sits on top of porous
Conflict of Interest Disclosures: All authors have completed
and submitted the ICMJE Form for Disclosure of Potential
Conflicts of Interest and none were reported.
Contract grant sponsor: Natural Sciences and Engineering
Research Council of Canada (NSERC); Contract grant sponsor:
Canadian Foundation for Innovation (CFI).
Ã
Correspondence to: James M. Fraser, PhD, Department of
Physics, Engineering Physics and Astronomy, Queen’s Univer-
sity, Kingston, Ontario, Canada K7L 3N6.
E-mail: james.fraser@queensu.ca
Accepted 30 November 2015
Published online in Wiley Online Library
(wileyonlinelibrary.com).
DOI 10.1002/lsm.22459
ß 2015 Wiley Periodicals, Inc.

2. cancellous (or spongy) bone usually occupied with bone
marrow. The typical composition of cortical bone is 13%
water, 27% organic matrix and 60% inorganic minerals by
weight [7,8]. The components of the organic matrix are
collagen, proteins, blood cells and lipids [9]. Inorganic
minerals, mainly a form of calcium phosphate known as
hydroxyapatite, are embedded in the collagen matrix [10].
The components exhibit differing density, optical absorp-
tion and heat conductivity, which are important properties
that affect laser bone ablation rate.
Various laser sources have been used to explore the laser
bone interaction from UV to IR wavelengths. Among
numerous candidates, Er:YAG (l ¼ 2.79 mm) and CO2
(l ¼ 9.6 mm) lasers are found to achieve efficient and
minimally invasive laser ablation on dental and bony
tissues [2,7,11]. The success of these two laser types is due
to the overlap of their wavelengths with water absorption
peaks. A widely accepted theory for laser bone ablation
mechanism is water microexplosion [12–14]. Strong
deposition of energy within a concentrated tissue volume
quickly heats up the water content in bone. This leads to
rapid vaporization of water and induces internal
microexplosions that blow off the surrounding tissues. In
contrast, ytterbium-doped fiber lasers (which have made
great advances in industrial settings due to their high
power, robust operation, and convenient fiber delivery) are
unpopular in bone ablation due to poor water absorption at
their center wavelength 1,070 nm. Therefore, laser bone
interaction at this wavelength is left under-explored.
However, ytterbium-doped fiber lasers have certain
advantages over Er:YAG solid-state and CO2 gas lasers
in surgical applications. The fiber beam delivery system
allows high flexibility in accessing hard-to-reach areas and
their high-power output has the potential for extremely
fast ablation. Due to their widespread adoption in
industry, considerable research and development has
made units smaller, more economical, and maintenance-
free. Ytterbium gain media have low quantum defect and
high wall-plug efficiency, which keeps operating costs low.
These are important considerations for clinical applica-
tions. A recent study determined that an ytterbium-doped
fiber laser is able to perform computer-assisted osteotomy
with high cutting efficiency and minimal damage to
surrounding muscle tissues [15]. In addition, the study
showed that healing time did not vary significantly
between laser and mechanical cutting techniques for
in vivo rabbit experiments, but precise control of the depth
is still an open research question.
Numerous studies have been done to parameterize laser
bone ablation in terms of wavelength dependence, ablation
rate, ablation threshold, and thermal side-effects. Most
quantitative results for ablation rate are drawn by dividing
the final depth of a hole or the total mass ablated by the
total number of laser pulses. The ablation threshold is
estimated based on the maximum number of laser pulses
incident on a location while zero hole depth is maintained.
Typical hole depth and morphology measurements are
obtained through optical profilometers [16], microbalance
weighing [17], and histology sectioning [18]. These
methods are limited by their lack of in-process depth
tracking information; only final depth and morphology can
be measured. Therefore, the ablation rate is statistically
interpolated from multiple holes ablated with increasing
laser pulses based on the assumption that ablation rate is
uniform across samples. Here we show that there is a
considerable amount of variability in bone ablation, and in
general this is not a good assumption. Perhavec et al. [19]
used an optical triangulation set-up to measure laser
ablation in hard dental tissue that allows multiple depth
measurements of the same hole during the formation
process. However, this approach can image only low aspect
ratio holes thus limiting potential clinical deployment.
In this work, a novel imaging diagnostic is exploited to
characterize laser bone ablation in situ. By depth tracking
the ablation front at high speed, we achieve automated
ablation of predefined 3D structures. Inline coherent
imaging (ICI), first applied to industrial laser machining
of metals and ceramics, utilizes low-coherence interferom-
etry to record sample morphology with micron precision
and microsecond timescales during laser processing [20].
The imaging probe of ICI is built coaxially inline with the
ablation laser such that ablation and imaging take place
simultaneously. Leung et al. [21] previously demonstrated
that ICI during laser bone ablation provides forward-
viewing capability that enables manual detection of
subsurface features beyond the ablation front so the
operator can terminate laser exposure before over-
ablation. In the present work, we demonstrate that with
proper signal processing, ICI-enabled closed-loop feedback
control allows bone ablation of predetermined 3D features
even in heterogeneous bone. In addition, ICI allows
identification of a particular regime of light-matter
interaction that generated minimally heat-affected bone
ablation with a high-power 1,070 nm fiber laser. Fresh and
dry bone samples yield similar results with minor
heat-affected zone (5–10 mm as verified by histological
examination). The flexibility of the 3D features and the
limited thermal side-effects indicates the promising
potential for ICI-guided laser bone surgery.
MATERIALS AND METHODS
In this proof-of-principle study, features are ablated
on small segments of ex vivo bone samples with a static
ablation laser. Sample movement is enabled by transla-
tion stages (Aerotech PRO-115/165) controlled by
Aerotech 3200 software. Ablation is performed with a
kW-class continuous-wave 1,070 nm ytterbium-doped
fiber laser (IPG YLR-1000-IC). Pulse durations of
50–100 ms with 15–30 mJ are used. A dry gas jet (e.g.,
nitrogen or compressed air) is delivered to the sample
through a nozzle to provide protection of the optics. A
flow rate of roughly 10À4
m3
/s is used. The optical probe
of the ICI system is built coaxially inline such that
imaging and ablation beam foci are spatially coincident
on the sample (refer to sample arm quadrant of Fig. 1).
This allows bone ablation dynamics and sample depth to
be tracked in situ.
2 YIN ET AL.

3. Optical Design of In Situ Imaging System
The ICI system comprises four components: broadband
light source, spectrometer, reference arm, and sample arm.
A 50:50 fiber coupler connects the four sections of the
system as illustrated in Figure 1 (adapted from ref. [22]).
The imaging light source is a superluminescent diode
(SLD) with a center wavelength of 840 nm and spectral
bandwidth of 25 nm (Superlum BLM-S-840-G-I-30). The
spectrometer consists of a transmission grating (Wasatach
Photonics WP-HD 1800) and a high-speed Si CMOS line
camera (Basler spL4096-140 km) capable of frame rates up
to 312 kHz. Image processing is performed with custom-
designed LabVIEW software on a four-core PC (Intel Core
i7 CPU 920). The reference arm has a fixed path length
with no moving parts. To minimize dispersion mismatch,
the sample and reference imaging fibers are designed to
be the same length. Imaging light from the SLD and the
ablation beam are coaxially aligned using a dichroic mirror
and then focused onto the sample through an objective lens
with 150 mm focal length. The imaging and ablation beam
have focus diameters of 70 and 210 mm, respectively. Due
to the nature of interferometric techniques, ICI achieves
high sensitivity and a large dynamic range of over 60 dB.
This is particularly relevant for observing deep features
where the collected backscattered light intensity varies
considerably over the course of an incision.
The underlying optical principles of ICI are similar to
that of spectral-domain optical coherence tomography [23].
White-light interferometry is used to measure the optical
path length difference (OPD) between a sample arm and
reference arm. The spacing of the interference fringe
pattern corresponds to a spectral domain frequency, which
can be Fourier transformed to the depth domain to extract
the OPD, or depth of the sample. The axial resolution is
inversely proportional to the spectral bandwidth of the
light source [23]. In this work, the choice of broadband SLD
and spectrometer design enables an axial resolution of
20 mm and a single sided field of view of 4 mm. Raw data
from the spectrometer are processed by: background
spectrum subtraction, Gaussian spectral shaping, linear
interpolation, and fast Fourier transform (FFT) [24]. For a
specific location of the sample, two final outputs are
available: (1) the sample reflectivity profile (known as an
axial-line or “A-line”), which shows backscattered inten-
sity as a function of depth, including backscattering from
sidewalls, ablation front and subsurface features and (2)
sample surface tracking, through which the depth of the
ablation front is extracted from the sample reflectivity
profile. A-lines are displayed in logarithmic brightness
scale and the intensities are in dB units as calculated
relative to the noise floor. Depending on the needs of
applications, ICI images can be used in either time-
resolved mode or scanning mode. A time-resolved mode
image is created by taking a series of A-lines at the same
location to record the ablation dynamics (in particular
depth change). A-line number is converted to time given
the A-line acquisition rate. Scanning mode involves
scanning the imaging beam across the sample to
generate a cross-sectional view of the sample. A-line
number is converted to distance given the A-line
acquisition rate and scan speed. Three-dimensional (3D)
morphology images can be built up with multiple
adjacent cross-sectional datasets by scanning over a two-
dimensional (2D) region. In this study: (1) time-resolved
mode image is used to monitor ablation dynamics in
real-time. (2) Scanning mode images are used before and
Fig. 1. ICI system comprises four components: broadband light source, spectrometer, reference
arm, and sample arm (where the imaging light is coaxially aligned with the ablation laser beam).
Figure is adapted from ref. [22].
AUTOMATED 3D BONE ABLATION 3

4. after ablation to display sample cross-sections. (3)
Morphology images of the final ablated 3D features are
used for comparison with the target design.
Data Processing and Depth Tracking
Automatic depth tracking of the ablation front requires
careful consideration of data acquisition and processing
such that the depth measurement is robust and accurate.
Due to finite spectrometer resolution, spectral domain
coherent imaging systems suffer from signal sensitivity
roll-off for increased OPD from zero-delay (zero-delay
corresponds to equal sample and reference arm length). In
addition, an artifact known as complex conjugate ambigu-
ity makes it infeasible to distinguish interfaces that have
equal OPD on different sides of zero-delay [23]. In this
work, imaging is performed strictly on one side of zero-
delay to avoid any ambiguity. The chosen side also allows
deeper features, for which less backscattered light enters
the collection optics, to appear closer to zero-delay and be
imaged with higher sensitivity.
Since both sidewalls (above the ablation front) and
subsurface features (below the ablation front) are imaged
together with the ablation front in a single A-line,
identification of the true ablation front is not clear cut.
In addition, like all imaging techniques that exploit
spatially coherent light, ICI suffers from speckle which
can dramatically change signal level due to subwavelength
surface variations [25]. By setting a larger imaging beam
diameter on the common lens (compared to the ablation
beam), the imaging beam focus at the sample (diameter of
70 mm) is much smaller than the ablation beam focus
(diameter of 210 mm) such that sidewall signals are
minimized. Even with the presence of subsurface features
in ICI images, the ablation front depth can be tracked by
the brightest interface in each A-line. This straightforward
tracking algorithm works because the largest change of
refractive index occurs at the interface between air and the
ablation front, yielding the most intense backscatter at this
depth. A small fraction of A-lines with weak backscatter
from the ablation front, most likely due to speckle, are
called “dark” A-lines. They are filtered out by proper
thresholding to prevent false depth tracking. A threshold
of 15 dB (in logarithmic brightness scale) is typically used,
which is found to be three standard deviations above the
noise floor and well below the typical brightest interface
intensity of A-lines ($25 dB) acquired on bone surfaces.
Also note that with the high imaging rate (up to 200 kHz),
filtering out the dark A-lines (typically less than 2%) does
not degrade the integrity of the ablation front depth
tracking.
Closed-Loop Feedback Control
Since sample depth can be extracted from ICI images as
the sample is being ablated, real-time depth feedback is
possible. For the simplest 1D case of percussion hole
drilling, a target depth can be defined in our custom
software and used to gate prespecified waveform of laser
pulses based on real-time depth measurements (i.e., the
laser fires if measured depth is shallower than target
depth). The ablation process is terminated once the target
depth is reached. To achieve automated ablation of a
predefined 3D morphology, the region to be ablated is
divided into a 2D grid with equally sized pixels and each
pixel has an assigned target depth. The ablation region is
scanned pixel by pixel and the laser fires at each pixel
based on ICI depth measurement. Due to finite depth
removal by a laser pulse (which is typically smaller than
the specified total removal depth), multiple scans are
generally required to ensure that the target depth is
reached for each individual pixel. The sample morphology
is plotted upon completion of each scan and the deviation
from the target shape is calculated. Dark A-lines may lead
to under-drilling since the laser is set to not fire in the case
of no detection of a bright interface. To help mitigate the
issue, an oversampling technique is introduced to check
the depth repeatedly. The pixel size is intentionally chosen
to be 50 mm  50 mm (i.e., smaller than the imaging beam
diameter of 70 mm) such that sufficient spatial overlap
exists between adjacent measurements. In addition, the
scanning pattern evaluates depth at each pixel each scan,
irrespective of information obtained from previous scans.
The oversampling technique is not optimal for speed, but it
is necessary to eliminate the chance of under-drilling and
ensure the accuracy of ICI depth feedback control.
Sample Preparation
Both fresh and dry ex vivo cortical bone samples are used
in this study. Fresh bone more closely resembles actual
surgical situations whereas dry bone produces more
consistent results. Dry bone is also much easier to preserve
before denaturing which makes it suitable for long-term
comparison studies. Both conditions are of research
interest to the laser osteotomy community [26,27]. Fresh
bone samples used for this work are taken from a porcine
tibia a few hours after the animal is slaughtered.
Connective tissue and periosteum are stripped off to reveal
the cortical bone surface. The fresh bone sample is
wrapped in wet cotton towel and stored in a sealed
container until use within 4 hours of collection. After fresh
bone experiments, the remaining porcine tibia is processed
and turned into dry bone by first boiling and then
air-drying. A comparison of fresh and dry bone ablation
characteristics is conducted using the same porcine bone
sample. A number of experiments are also conducted with
dry bovine bone. Bone type is reported for each experiment
in this study.
Ex Situ Verification and Histological Examination
In order to quantify the accuracy of ICI depth feedback
control, laser ablated 3D structures are examined using a
stylus profiler (Bruker DektakXT) and compared with
the designed morphology by computing the root mean
square (RMS) deviation. RMS deviation is a measure of
the difference between two datasets (calculated as the
square root of the mean differences squared). In addition,
the thermal side-effects of laser ablation in bone are
4 YIN ET AL.

5. assessed using the following histological procedures.
Immediately after ablation, bone specimens are fixed
with formalin solution for 48 hours. The specimens are
then decalcified with HCL solution, dehydrated with
alcohol solutions and embedded in paraffin. A number of
5 mm thick sections are cut out parallel to the axis of the
laser beam. The sections are then stained with haematox-
ylin and eosin (H&E) and examined under a light
microscope. The extent of the heat-affected zone can be
quantified by the color change in the histology sections.
RESULTS AND DISCUSSION
To quantify laser ablation rate and ablation threshold,
the laser beam is set incident on both fresh and dry porcine
tibia for percussion hole drilling. Figure 2 shows a typical
time-resolved ICI image of percussion drilling where the
bottom of the ablated hole descends with increasing pulse
number. Laser exposure begins at 0 second (from the top of
the figure) and ends at 3 seconds. Hundred micro seconds
laser pulses of 30 mJ are fired at 100 Hz. This time-resolved
image of hole formation shows three stages: no ablation
when energy deposited is below threshold (between
0 second and $1 second), sudden onset of ablation, and
lastly bone removal with fairly uniform ablation rate.
Subsurface features are visible in this dataset (Fig. 2a), yet
depth tracking by brightest interface successfully extracts
the depth of the hole bottom (Fig. 2b). Depth tracking by
brightest interface is therefore utilized to significantly
reduce the amount of data that needs to be processed to
extract the bone ablation characteristics. Three character-
istics of percussion drilling can be obtained: (1) The hole
depth is taken as the depth difference of the sample before
and after ablation. (2) The ablation threshold is deter-
mined from the total energy deposited until an abrupt
change of slope of the bone surface. (3) The ablation rate is
calculated from the slope of the linearly descending
ablation front. These results show that initially (due to
poor optical absorption at 1,070 nm) the bone ablation
process is not efficient, but the abrupt onset of ablation is
likely due to a highly localized thermal modification to the
bone. The underlying mechanism of this process is not
clear but is exploited in the remainder of this work.
Depth tracking of the ablation front by brightest
interface is performed on 10 percussion drilling datasets
obtained at adjacent locations on fresh porcine tibia in
the cortical layer (Fig. 3a). The tibia piece is later boiled
and dried and the same experiment is repeated on the dry
specimen (Fig. 3b). Using the above procedure, the
average ablation depth over ten holes for fresh bone is
found to be 880 mm with a standard deviation of 151 mm.
The average ablation depth for dry bone is larger at
1,038 mm with a standard deviation of 148 mm. The onset
of ablation is identified visually in post-processing
showing the average ablation threshold to be 3.0 Æ 1.3 J
and 1.9 Æ 0.8 J, respectively, for the fresh and dry bone.
Linear regression is performed to each dataset from onset
Fig. 2. Time-resolved ICI image of percussion drilling of dry porcine bone with 100 ms pulses of
30 mJ at 100 Hz. a: Laser and imaging beams are incident from the top of the figure and the surface
depth is initially $400 mm. Laser exposure starts at 0 second and stops at 3 seconds. Flat slope from
0 second to 1 second indicates no ablation. After onset at $1 second, the negative slope indicates
fairly uniform ablation rate. b: Identical data as above with brightest interface depth tracking
overlaid.
AUTOMATED 3D BONE ABLATION 5

6. of ablation to the end of laser exposure at 3 seconds. The
average ablation rates, calculated based on the linear
regression slopes, are 117 Æ 15 mm/J and 134 Æ 3 mm/J for
fresh and dry bone, respectively. Large variations in
ablation depth are observable in each bone sample, most
likely due to inhomogeneities in bone, which result in
different ablation onsets and rates. For example, blue
data in Figure 3a (fresh bone) show that ablation is
interrupted several times, indicating structures more
resistant to ablation. The red data in Figure 3a (fresh
bone) show a drop in the ablation front on the order of
100 mm, which is likely due to a pore in the bone. In
comparison, the dry bone has more uniform ablation with
no signs of obvious interruptions or cavities, possibly
because the boiling process removes blood vessels and
other organic compounds. Fresh bone has a shallower
average depth due to its smaller average ablation
threshold and rate. This is most likely due to the
presence of water (approximately 13% of fresh bone’s
total weight [8]) and intact organic components (such as
collagen and protein).
From the previous results, it can be concluded that the
bone ablation process is overall nondeterministic. In some
surgical applications, it is desirable to drill to a certain
depth regardless of the initial surface conditions (rough-
ness and tilt). Here we show drilling with fixed laser
parameters does not result in consistent depth. Figure 4a is
a scanning mode ICI cross-sectional image demonstrating
the variability of drilling with constant laser parameters
on dry bovine bone. Five holes are drilled with two hundred
30 mJ pulses of 100 ms at 100 Hz. The standard deviation is
found to be 203 mm from the mean for the open-loop
ablation. In comparison, Figure 4b shows five holes drilled
using the same laser pulse parameters with the addition of
ICI closed-loop feedback, where drilling is terminated once
a target depth of 2,000 mm is reached. The hole depths
show much better consistency, with a standard deviation of
39 mm from the mean (more than 5 times better than open-
loop ablation). The open-loop holes are drilled on the same
bone sample, very close to each other, on a fairly flat
surface. In a clinical setting, without these preconditions it
is likely that the precision for a system without closed-loop
feedback would be much worse. It is also worth pointing out
that there is no a priori way of determining the local
ablation rate of a sample of bone, so depth feedback control
is necessary if cutting is performed above critical
structures or tissues like in brain or spinal surgery.
Surgical procedures often require a 2D incision along a
straight (or curved) path to a desirable depth. In this
experiment, a 5 mm long incision is cut to a target depth of
1,000 mm using ICI closed-loop feedback control. The laser
pulse used to cut the trench is 100 ms with an energy of
30 mJ. In Figure 5a, the virgin bone surface is imaged by
ICI, and the yellow line indicates the target depth of the
incision. With the scan speed of 5 mm/s, an A-line is
acquired every 50 mm (smaller than the imaging spot size
to ensure spatial oversampling and prevent under-
drilling). If the brightest interface in an A-line is shallower
than the target depth, the laser fires a pulse. If the target
depth has been reached or exceeded, no laser pulse is fired.
In Figure 5b, the final trench shows excellent agreement
with the target shape and verifies the accuracy of
ICI-enabled depth control.
For further verification, an independent stylus profiler is
deployed to measure sample morphology ex situ. The stylus
profiler has high axial resolution ($nm) but due to the finite
transverse sizeofthe stylus,itcannotbeusedfor thenarrow
features shown in Figures 4 and 5. To avoid limitations of
the profiler, a wide feature with a triangle wave pattern is
designed. The target depth is specified every 50mm
indicated by blue dots in Figure 6. The feature is ablated
on dry bovine bone with 100 ms pulses of 30 mJ and a scan
speed of 5 mm/s using automatic closed-loop feedback
control. The ablated shape is measured by the profiler ex
situ (red dots in Fig. 6). A standard problem of using ex situ
analysis to verify final morphology is co-registration of the
final sample with the target shape (i.e., tilt and translation
of one dataset relative to the other). Co-registration of the
Fig. 3. Tracked laser drilling of (a) fresh and (b) dry porcine bone
for 10 holes drilled with three hundred 100 ms pulses of 30 mJ at
100 Hz (dot colour corresponds to different holes). The virgin
surfaces are placed at zero depth for comparison purposes. Laser
exposure starts at 0 second and ends at 3 seconds.
6 YIN ET AL.

7. two datasets is achieved in three steps: downsampling of
profiler data, tilt compensation, and horizontal location
matching. The profiler data are originally sampled every
0.825 mm, while the target map spacing is 50 mm, so the
profiler data is downsampled by averaging. The regions
shown in gray in Figure 6 are used for co-registering the two
datasets. Tilt compensation and horizontal translation of
the profiler dataset are performed to minimize the RMS
deviation for data in the gray regions. The RMS deviation of
the central zone is found to be 12.6mm indicating excellent
agreement, which serves as a clear ex situ verification of ICI
depth feedback control. An overall vertical offset error in
the final morphology cannot be assessed due to the
co-registration procedure. It is unlikely given excellent
agreement demonstrated by Figure 5.
Extending 2D control into the third dimension, a raster
scan pattern is implemented to allow a predefined 3D
shape to be ablated on bone. Figure 7a shows an example
of a raster scan pattern over a 10 Â 10 grid with equally
sized pixels. Position synchronized output (PSO) is
generated when the laser head reaches the center of a
pixel (indicated by a red dot in Fig. 7a). The PSO triggers
ICI to measure the depth and ICI-enabled feedback
controls pulsing of the laser. PSO plays an important
role to ensure that the laser fires at fixed distance intervals
even when the sample scan speed changes at feature
boundaries. This is an important engineering design since
the feature will be distorted if the laser fires at fixed time
intervals with a varying scan speed. The larger and
more complicated morphology of a 3D feature increases the
time required for ablation. In the current implementation
in which there is considerable time for the laser head to
travel between pixels (10 ms), time for heat dissipation
between laser pulses is included, resulting in a duty cycle of
1% (100 ms pulses out of 10 ms). The duty cycle (and thus
ablation speed) could be dramatically increased while
maintaining the heat dissipation time between adjacent
pixels by employing scanning mirrors and a more
sophisticated raster scan pattern. Efforts are made in
the current work to optimize the laser pulse sequence for
improved ablation efficiency while maintaining minimal
thermal side-effects. Using ICI, it is determined that a
pulse waveform of two 50 ms pulses with 150 ms inter-pulse
delay is 20% more efficient than 100 ms single pulse for dry
bone ablation. The double pulse waveform is therefore used
for ablating the 3D features described below. In order to
demonstrate the flexibility of the 3D laser ablation process,
a complicated feature is designed to be ablated on bone. A
Bessel shape is computed using the MATLAB built-in
Bessel function of the first kind with order one and
cylindrical symmetry (Fig. 7c). The designed feature has a
size of 6 mm  6 mm and depth range of 2000 mm, specified
by a 120 Â 120 matrix. Each matrix element corresponds
to a pixel target depth with an area of 50 mm  50 mm.
Fig. 4. Cross-sectional view of percussion drilled holes with 100 ms pulses of 30 mJ at 100 Hz in dry
bovine bone. a: Five holes drilled with 200 pulses. b: Five holes drilled using ICI closed-loop
feedback to control the pulse number until reaching a consistent depth.
AUTOMATED 3D BONE ABLATION 7

8. The maximum scan speed is set at 5 mm/s and PSO is
generated at every pixel during raster scanning. The
Bessel shape is ablated using ICI closed-loop feedback
control into dry bovine bone as shown in Figure 7b. The
final feature is measured by ICI after ablation has finished
(Fig. 7d). The RMS deviation between the designed model
and the ablated feature is 108 mm. It can be noted from
Figure 7d that most deviation occurs at boundary of the
feature. The finite ablation beam focus size ($200 mm in
diameter) limits the precision of ablating abrupt edges
with steep slopes (898 by design). To explore this further,
the RMS deviation is calculated for two separate regions:
(1) the center 116  116 pixels (5.8 mm  5.8 mm) that
correspond to the Bessel shape (with maximum slope of
748) and (2) the outer two pixels on each side (100 mm,
about half the size of laser focus) that correspond to the
boundary edges. For the central region the RMS deviation
is 59 mm, showing good agreement with the designed
model. For the outer pixels the RMS deviation is 357 mm,
showing our technique has limited precision in creating
abrupt features. By considering only the outer two pixels
we find that the median slope is 818, which can be
interpreted as an estimate of the upper bound for steep
edges with these laser and raster scan parameters.
Techniques that would likely increase this bound include
beam shaping (e.g., generating a top-hat intensity distri-
bution at the sample surface), trepanning and/or dynamic
control of the cutting angle enabled by robotic deployment.
Nevertheless, the complexity of the Bessel shape and the
good agreement of the central region demonstrate the
flexibility of ICI-guided ablation.
To quantify the thermal side-effects of the laser ablation
of bone, histological examination is performed on 3D
features that are ablated on both fresh and dry porcine
tibia. The heat-affected zone (HAZ) is identified by tissue
color change under a light microscope. Figure 8a shows the
Fig. 6. Comparison of profiler measured morphology (in red) and
designed morphology (in blue). The RMS deviation between the
two datasets is 12.6 mm. Gray areas indicate the data region used
for co-registration.
Fig. 5. A trench drilled with 100 ms pulses of 30 mJ at 100 Hz in dry bovine bone to the target depth
of 1,000 mm. a: Bone virgin surface with indicated target depth in yellow. b: Ablated trench showing
excellent agreement with the prespecified target shape.
8 YIN ET AL.

9. histology cross-section of the inset photograph of an
inverted pyramid feature (with a size of 3 mm  3 mm
and a depth range of 1.2 mm). The feature is ablated onto
dry porcine tibia using ICI depth feedback control and the
same laser parameters exploited for the Bessel shape.
There is little charring and carbonization visible on the
surface of the ablated feature suggesting the thermal side-
effects are minor. Figure 8b shows a magnified region of
the feature (indicated by black box in Fig. 8a), confirming
that the heat-affected zone is small (about 5–10 mm). These
are much better than previously reported results for the
same laser source [26] due to our particular choice of laser
parameters, and are similar to HAZ reported for water
microexplosion driven CO2 laser ablation [28]. The low
duty-cycle likely allows heat to dissipate from the
surrounding tissue, thus regions around the target volume
never exceed a temperature threshold required for
carbonization. Abrupt edges are found to have larger
HAZ compared to smooth features, due to the tails of the
Gaussian-profile laser beam, which heat the surrounding
bone without ablating it. Beam shaping techniques can be
used to reduce this effect. Thermal damage can be
minimized by designing features with gradual changes
in height. The small HAZ identified for the 1,070 nm fiber
laser shows promising potential for future clinical
deployment.
CONCLUSIONS
Lasers are emerging as surgical tools due to their high
transverse flexibility and noncontact cutting nature.
Innate heterogeneity of bone makes laser ablation nonde-
terministic even for fixed laser parameters. This is
problematic because cutting depth control is extremely
important in clinical operations where critical soft-tissue
lies underneath ablated bone structures. Inline coherent
imaging is a diagnostic tool that is combined coaxially with
the ablation laser such that in situ real-time depth
measurements can be obtained during bone ablation
processes. The bone ablation characteristics are examined
by performing ICI-monitored bone percussion hole drilling
Fig. 7. a: Raster scan pattern used in 3D ablation of the Bessel shape on dry bovine bone. b:
Photograph of the Bessel shape ablated with ICI closed-loop feedback control. c: Designed model for
the Bessel shape with size of 6 mm  6 mm and depth range of 2,000 mm. d: Final ablated feature
measured by ICI, showing good agreement with the model in the central region and larger
deviations at the abrupt boundary edges.
AUTOMATED 3D BONE ABLATION 9

10. experiments. Ablation rate and threshold are determined
from the ICI depth tracking data of the hole formations. It
is found that the threshold energy required to initialize the
onset of ablation is variable, and once ablation starts the
rate is fairly constant. Automatic closed-loop feedback is
implemented to control the ablation depth; 1D (holes), 2D
(trench), and 3D (Bessel shape) features are produced to
demonstrate the effectiveness of the feedback system. ICI
feedback control is further verified by ex situ measure-
ments with an independent stylus profiler. These results
have great significance for surgery applications where
laser cutting depth is critical. Histological examination is
also performed on laser ablated 3D features and the heat-
affected zone is quantified as 5–10 mm. This is the first
work to report minimally heat-affected bone ablation with
an ytterbium-doped fiber laser at 1,070 nm. This result is
unexpected given the general acceptance of the water
microexplosion mechanism and the low-water absorption
at 1,070 nm. The limited thermal side-effects in combina-
tion with the ICI-enabled high-precision depth control
suggest that ytterbium-doped fiber lasers can be deployed
as promising surgical tools with advantages of robustness,
high-power output, and low operating costs. The deploy-
ment of scanning mirrors will benefit clinical applications
to allow quick and flexible movement of the laser beam. In
this regard, future work will fully exploit scanning mirrors
to achieve faster overall ablation due to improved duty
cycle, while maintaining comparable thermal side-effects.
Future research will also exploit microsecond-resolved ICI
with mathematical modelling to identify the cause for
improved ablation with double pulse waveforms, and the
underlying physical processes responsible for bone abla-
tion at 1,070 nm.
ACKNOWLEDGEMENTS
Chenman Yin led efforts on conducting experiments,
analysis of results, wrote the majority of the manuscript.
Dr. James Fraser supervised the research team, provided
guidance, coauthored the manuscript. Sacha Ruzzante
performed data analysis, coauthored the manuscript. Yang
Ji made important contributions to LabVIEW software
automation. Dr. Paul Webster designed, built the ICI
imaging system, Laser Depth Dynamics Inc. provided
software support. Oliver Jones performed the histology
sectioning, provided insights on interpretation of histology
images. Dr. Donald Maurice kindly supplied the fresh
porcine bone sample. Funding for this work is provided by
Natural Sciences, Engineering Research Council of
Canada, the Canadian Foundation for Innovation.
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