Monopolar Electrosurgical Thermal Management for ... - IEEE Xplore

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 59, NO. 1, JANUARY 2012

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Monopolar Electrosurgical Thermal Management for Minimizing Tissue Damage Robert E. Dodde*, Jacob S. Gee, James D. Geiger, and Albert J. Shih

Abstract—In this study, a novel thermal management system (TMS) is developed for the minimization of thermal spread created by a monopolar electrosurgical device, the most commonly used surgical instrument. The phenomenon of resistive heating of tissue is modeled using the finite-element method (FEM) to analyze the electrical potential and temperature distributions in biological tissue subjected to heat generation during monopolar electrosurgery. Ex vivo experiments are used to validate the FEM by comparing the model predicted and experimentally measured temperatures. The predicted FEM maximum temperature 1.0 mm adjacent to the electrode is within 1% of the experimentally measured maximum temperature using a standard monopolar pencil electrode. A TMS consisting of adjacent cooling channels produces coagulation volumes 80% that of standard monopolar procedures while maintaining comparable temperatures in the targeted tissue below the electrode. In vivo temperatures using a device incorporating a TMS at distances of 2 and 3 mm adjacent to the electrode edge are maintained below temperatures known to damage tissue. Index Terms—Electrosurgery, finite-element method (FEM), monopolar, thermal management.

I. INTRODUCTION ONOPOLAR electrosurgical instruments are widely used during surgical operations to incise, ablate, and dissect tissue while promoting hemostasis by transferring electrical energy to the tissue in the form of heat generation through resistive heating in the tissue. These instruments are vital in coagulating tissue: a widely used surgical technique that denatures proteins to minimize bleeding during surgical procedures [1]. Without proper coagulation, internal bleeding during a surgical procedure is a danger to the patient and obstructs the surgeon’s field of view. At frequencies over 100 kHz, the energy in radio frequency (RF) electrical current can be delivered safely to generate the heat necessary for coagulation [2]. However, a substantial amount of heat created by electrosurgical devices

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Manuscript received June 13, 2011; revised August 27, 2011; accepted September 1, 2011. Date of publication September 22, 2011; date of current version December 21, 2011. This work was supported in part by the National Science Foundation under Grant CMMI 0620756. Asterisk indicates corresponding author. *R. E. Dodde is with the Department of Biomedical Engineering, The University of Michigan, Ann Arbor, MI 48109 USA (e-mail: [email protected]). J. S. Gee is with Ethicon Endo-Surgery, Cincinnati, OH 45242 USA (e-mail: [email protected]). J. D. Geiger is with the Department of Surgery, The University of Michigan Medical School, Ann Arbor, MI 48109 USA (e-mail: [email protected]). A. J. Shih is with the Department of Mechanical Engineering, The University of Michigan, Ann Arbor, MI 48109 USA (e-mail: [email protected]). Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/TBME.2011.2168956

has been shown to spread throughout the tissue leading to unintended thermal damage and impairment of a patient’s postoperative quality of life [3]–[6]. Due to this, special attention must be given by surgeons to minimize unneeded thermal damage. A monopolar device consists of a pair of electrodes: an active electrode with a small surface area positioned at the surgical site and a larger dispersive electrode placed on the patient’s body to close the electrical circuit. The electrical resistance of tissue generates heat within the tissue due to the high current density created by the small contact area between the electrode tip and tissue. Unintended thermal damage of surrounding tissue is referred to as thermal spread [7]. This thermal spread has been linked to tissue damage in various surgical procedures and has been shown to occur at temperatures as low as 50 ◦ C [8], [9]. Concern with thermal spread has provided a research goal for the development of a new class of electrosurgical devices using a thermal management system (TMS) to control the tissue temperature surrounding the site of the procedure [10]. Numerous procedures where precise noninjurious dissection is required can benefit from the use of a TMS. Entry point skin incisions [11], dissection of pelvic organs through natural dissection planes [12], vessel harvesting [13] for coronary bypass procedures and nerve-sparing techniques [3] are examples of areas, where a TMS can provide the needed thermal spread minimization while not compromising performance. Other technologies have been proposed to solve the issue of thermal spread. These include other energy-based technologies, such as bipolar electrosurgery and ultrasonic coagulation, and the use of nonenergy-based technologies such as the use of knives and scissors for dissections, heat pipe electrodes, and saline irrigation. However, thermal spread has been shown to be a concern even when using advanced energy-based devices [6], [7]. Monopolar energy is often preferred by surgeons because it is faster than bipolar or ultrasonic and still provides a degree of hemostasis compared to using nonenergized devices. Heat pipe technology is directed more at tissue sticking and charring concerns and is more adept at maintaining a cool electrode than minimizing lateral thermal spread [14]. Saline irrigation during electrosurgery can occlude the field of view of the surgeon and requires significantly higher powers to be used due to the reduced surgical effect from lowered current densities [15]. Nonconductive irrigants can lead to hemolysis and hyponatremia [16]. The physical phenomenon of thermal spread in biological tissue is difficult to measure and predict. The finite-element method (FEM) is a necessary tool for understanding how thermal energy is generated and dissipates through tissue during monopolar electrosurgical coagulation (EC) [17]. With this

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Fig. 2. (a) Experimental setup showing the tissue, electrode, and thermistors during for ex vivo testing. (b) Underside of the fixture showing the cooling channel used for TMS experiments. Fig. 1. (a) Configuration of the proposed monopolar device incorporating a cooling channel for a TMS. (b) Valleylab E2504 monopolar pencil electrode used in coagulation experiments.

understanding, the design of monopolar surgical devices can be optimized allowing the surgeon greater control over thermal spread. Previous FEM research on monopolar electrosurgical procedures has been focused on RF tissue ablation [9], [18], [19]. This procedure is primarily used for tumor ablation, such as in the liver, or cardiac ablation. It is characterized by voltage inputs on the order of 10–20 V and procedural times on the order of minutes [20], [21]. EC is a different technique than RF ablation, one characterized by high voltage inputs on the order of 40–100 V and procedural times on the order of seconds [7]. EC procedures are typically performed numerous times during a procedure. To date, detailed FEM analysis of voltage and temperature distributions of the EC procedure is still new and not well studied. The continued creation and study of models that mimic surgical procedures are important to the understanding of how biological tissue responds to RF energy. This knowledge can lead to design improvements of surgical devices and procedures. In this study, a proof-of-concept device is studied for its ability to control lateral thermal spread from a monopolar electrode during EC. The COMSOL (Burlington, MA) multiphysics simulation software is used to perform FEM simulations of monopolar EC for standard procedures as well as for a novel device [see Figs. 1(a) and 2(b)] incorporating a TMS to investigate the voltage and temperature distributions in biological tissue. Ex vivo experiments are performed on bovine liver using a standard and TMS monopolar pencil. The measured temperatures are compared with the model’s simulation results to validate the FEM model. The effect of TMS during an in vivo experimental procedure is discussed. II. MATERIALS AND METHODS A. Monopolar FEM Formulation The analytical modeling of heat transfer in tissue, or bioheat transfer, to account for heat sources and sinks from metabolism and blood perfusion was the pioneering work of Pennes [22]. The bioheat transfer model of tissue includes coupled thermal and fluid (blood) transport phenomenon. Previous work has shown that solid elements can be used to model tissue, a

multiphase material consisting of both solid and liquid, with sufficient accuracy [23]. The linear bioheat transfer equation for tissue is the general heat equation for conduction with added terms for heat sources and sinks and can be expressed as [22] ρc

∂T = k∇2 T + ρb cb wb (T − Tb ) + qm + qg ∂t

(1)

where ρ, c, and k are tissue density (kg/m3 ), heat capacity (J/kg·K), and thermal conductivity (W/m·K), respectively, wb is the effective blood perfusion parameter (1/s), ρb is the blood density (kg/m3 ), cb is the blood heat capacity (J/kg·K), T is the local tissue temperature (◦ C), Tb is the blood inlet temperature (◦ C), qm is the metabolic heat generation rate of the tissue (W/m3 ), qg is the external induced heat generation rate due to electrosurgical heating of the tissue (W/m3 ), and t is time (s). For all cases, it was assumed that the metabolic heat source and blood perfusion were zero (qm = 0 and wb = 0) as the experiments were performed ex vivo. A quasi-static electrical conduction model was applied to solve the electric field in the tissue using Laplace’s equation [24]: ∇(σ(T )V ) = 0

(2)

where σ(T) is the temperature-dependent electrical conductivity (S/m) and V is the electric potential (V). At electrosurgical generator frequencies (300–550 kHz) capacitive coupling in tissue is negligible and only the resistive dissipation of the electrical energy needs to be considered [19]. The electrical and thermal properties of the tissue are available in [9], [21], [25]–[27]. The properties for all experimental materials used are shown in Table I. For the σ(T) of bovine liver, an increase of 2%/◦ C is used in accordance with Schwan and Foster [28]. For tissue temperatures exceeding 100 ◦ C, the electrical conductivity was dropped by a factor of 10 000 to simulate tissue desiccation as done previously by Haemmerich et al. [29]. It has been previously shown that other parameters within the model are also temperature-dependent, specifically k and c [30], [31]. However, the effects are small in comparison to that of σ(T) and are not considered in this model in order to reduce computational time [17]. In the multiphysics software COMSOL 3.5 a, the bioheat transfer equation is coupled with the Conductive Media DC Module to solve for (1) and (2) simultaneously. The equations

DODDE et al.: MONOPOLAR ELECTROSURGICAL THERMAL MANAGEMENT FOR MINIMIZING TISSUE DAMAGE

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TABLE I THERMAL AND ELECTRICAL MATERIAL PROPERTIES USED IN THE FEM

are coupled through the externally induced heat generation term (qg ) from (1). This is also the resistive heating of the tissue from the electrical energy and is defined as qg = J·E, where J is the current density vector (A/m2 ) and E is the electric field vector (V/m). For all simulations, the electric field E and temperature T are calculated with the COMSOL PARDISO direct solver using matrix row elimination [32].

0 V. For the TMS FEM model, the inside diameter of the cooling channels were given a constant temperature (T = 22 ◦ C) boundary condition. All other internal boundaries were given electrical and thermal continuity boundary conditions. Tissue was considered to be coagulated after reaching 50 ◦ C based on previous literature values [8], [9]. C. Experimental Equipment and Setup

B. 3-D FEM Analyses of Monopolar Coagulation FEM analyses were performed for a standard and TMS monopolar procedure. An overview of the TMS model is shown in Fig. 3(a) with a monopolar electrode in the center of the tissue and cooling channels adjacent to both sides of the electrode. In the standard monopolar model, the cooling channels are removed. Each case modeled a 5-mm-thick rectangular section of ex vivo bovine liver tissue (100 mm long, 50 mm wide). The stainless steel electrode is 25 mm long, 2.5 mm wide and 0.5 mm thick. The end of the electrode embedded in the tissue has a 1.25-mm radius and is embedded 1.5 mm into the tissue. For the TMS model, 1-mm diameter, 12-mm-long stainless steel cooling channels with 0.1-mm wall thickness were positioned 1.0 mm away from each side of the monopolar electrode and are embedded 0.5 mm deep into the tissue. As we are interested primarily in the solution at plane ABCD [see Fig. 3(b)], the curvature of the cooling channel is not included to decrease model complexity. Due to symmetry, only the upper right quarter of the geometry is modeled using the FEM as shown in Fig. 3(b). Details of the TMS FEM geometry are shown in Fig. 3(c)–(d). Both models were run at increasingly finer mesh sizes until the maximum temperature within the tissue changed by less than 0.2 ◦ C. The final mesh for the TMS monopolar has 36 207 second-order tetrahedral elements. The mesh for the standard monopolar FEM has 25 010 second-order tetrahedral elements. The increased number of elements in the TMS FEM model is primarily due to the meshing of the cooling channel. Symmetry plane boundary conditions were applied on planes ABCD and EAGC [see Fig. 3(b)]. Electrical insulation and constant temperature (T = 22 ◦ C) boundary conditions were applied on planes BFDH and FEHG. Electrical and thermal insulation boundary conditions were applied on planes EFAB and CDGH. Further details are included in Table II. A voltage equal to the experimentally measured root-mean-square (RMS) voltage was applied along the entire surface of the monopolar electrode. The average RMS value for both procedures was 43 V. The portion of the electrode outside of the tissue and not on a symmetry plane was given a free convection thermal boundary condition with h = 25 W/m·K [33]. The surface of the ground plate in contact with the tissue was given a boundary condition of V =

A Valleylab E2504 Monopolar pencil was used for all experimental procedures as shown in Fig. 1(b). The electrode was powered by a Valleylab SurgiStat generator. The voltage output from the generator was measured using an Agilent (Santa Clara, CA) 10076A 100:1 high voltage probe connected to a National Instruments (Austin, TX) NI PXI-5114 oscilloscope card as part of an NI PXI-1033 chassis. Data was acquired and analyzed using an NI LabVIEW program. Tissue temperature was measured using Alpha Technics (Irvine, CA) microthermistors with 0.48 mm outside diameters and 0.25 s thermal response times. Thermistors have been shown to work effectively near electrosurgical devices, due to their relative immunity to electromagnetic interference and their stability and high sensitivity in the targeted temperature range (30–100 ◦ C) [34]. Temperature readings were acquired using an NI-PXI 6221 I/O card and an NI LabVIEW program. A polycarbonate fixture was used to position the thermistors at set distances in relation to the electrode. The overall dimension of the fixture is 37 × 20 × 12 mm and consists of a center cutout for the electrode and three 0.5 mm diameter holes at specific distances, 1.0, 2.0, and 3.0 mm, from the center. The microthermistors were fixed in these holes to measure the temperature inside the tissue at a set depth of 2.0 mm for comparison to the FEM temperatures. Fig. 2(a) shows the monopolar electrode with the attached fixture and thermistors. A 5 mm thick cut of bovine liver was used for all ex vivo EC experiments. Configuration of a proposed TMS monopolar device using a cooling channel for continuous flow of coolant is shown in Fig. 1(a). Due to the difficulty in manufacturing this device, a proof of concept TMS device was created by securing a cooling channel to the bottom of the standard monopolar fixture as shown in Fig. 2(b). Room temperature (22 ◦ C) distilled water was pumped through the channels with a peristaltic pump. D. Experimental Procedures As shown in Fig. 2(a), the monopolar electrode with the attached thermistor fixture was placed on the tissue such that the electrode penetrated into the tissue 1.5 mm deep. The electrode was activated in the “coag” mode at a power level of 3 for duration of 3 s (standard monopolar) or 2.5 s (TMS monopolar). The

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TABLE II BOUNDARY CONDITIONS FOR PLANES MARKED IN FIG. 3(B)

device was held in the tissue for an additional 10 s to measure the dissipation of thermal energy through the tissue. For TMS experiments, coolant flow was initiated prior to energizing the electrode and continued to flow throughout the entire procedure. The procedure was repeated a minimum of three times. The device was cleaned between each procedure to remove any charred tissue. In vivo testing was performed using a canine spleen model in accordance with the University Committee on Use and Care of Animals, Ann Arbor, MI. Seven coagulations were performed using each device following the experimental procedures for the ex vivo tests.

III. RESULTS A. Experimental Validation of the FEM Coagulation Model

Fig. 3. (a) Schematic of TMS experimental geometry showing symmetry planes (black arrows). (b) Detail of upper right quadrant of (a) showing the final FEM model geometry. (c) Geometric detail showing relationship between the electrode, tissue, ground plate, and cooling channel. (d) Cross-sectional view of dashed box in (c) detailing geometry of electrode and cooling channel. Note that the standard monopolar geometry is identical except the cooling channel is removed. Units in millimeter.

Fig. 4(a) shows the comparison of experimentally measured tissue temperatures with FEM predictions during monopolar EC in ex vivo bovine liver. The temperatures at three points, 1.0, 2.0, and 3.0 mm away from the electrode, are presented. The FEM accurately predicts general trends for thermal profiles during active electrosurgical heating compared to the experimental measurements. The percent difference between maximum temperature values at the three distances are 0.6%, 3.8%, and 8.8%, respectively. While the FEM is able to model the postsurgical cool-down quite accurately, the peak temperature at 3.0 mm was not attained in the FEM. This could be attributed to the position of the thermistor experimentally or alteration of tissue parameters caused by the denaturation of proteins identified in [23]. After validating the FEM model for EC, the concept of a TMS was analyzed. Fig. 4(b) shows the thermal profiles created by placing cooling channels 1.0 mm adjacent to both sides of the electrode. Experimental temperature measurements at 2.0 and 3.0 mm show decreases in the maximum temperature of 21.9% and 14.9%, respectively, when compared to the standard monopolar procedure. Experimental temperature measurements could not be measured at the 1.0-mm distance because of conflict with the location of the cooling channel. However, the finite-element analysis of the TMS EC procedure predicts a lower maximum temperature and quicker temperature decay when compared to that of the standard EC procedure. FEM determines the time it will take the 1.0-mm temperature to decay 10% from its maximum temperature to be reduced from 1.8 to 1.15 s with the introduction of cooling channels.


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Fig. 6. Cross-sectional view of temperature profiles on the ABCD plane for FEM of standard monopolar procedure.

Fig. 4. Comparison of thermal profiles for ex vivo experiments and FEM of the EC procedure. (a) Standard monopolar and (b) TMS monopolar. Measurement 1 mm from the electrode edge was not acquired experimentally for TMS case due to position of cooling channel.

Fig. 7. Cross-sectional view of temperature profiles on the ABCD plane for FEM of TMS monopolar procedure.

Fig. 5. Cross-sectional view of FEM voltage distribution at t = 2.5 s for (a) standard monopolar and (b) TMS monopolar.

B. Voltage and Temperature Comparison Between Standard and TMS Monopolar In order to compare the standard and TMS monopolar devices, a separate FEM was run for the TMS monopolar device using the same voltage input as the standard monopolar device. The resulting potential distribution at t = 2.5 s for a standard and TMS monopolar procedure is shown in Fig. 5. The high conductance cooling channel reduces the resistance to ground for current following pathways near the tissue surface. This results in an increased voltage drop and increased current density in those pathways assuming the volumetric tissue resistance is the same. This increased current density generates more resistive heating near the tissue surface as seen in the higher tissue

temperature readings between the electrode and cooling channel for the TMS device as compared to the standard monopolar device [see Figs. 6(b) and 7(b)]. Figs. 6 and 7 show the temperature distribution within the tissue for the standard and TMS procedure, respectively. Maximum tissue temperatures in the tissue at 3 s are comparable but slightly higher for the TMS monopolar procedure (110.1 ◦ C versus 112.3 ◦ C). However, tissue is heated more evenly along the electrode using the TMS device due to the increased current density near the surface of the tissue as seen by comparing Figs. 6(b) and 7(b). Additionally, the TMS device increases the cooling rate of tissue as seen in the lower temperatures in Figs. 6(d)–(f) and 7(d)–(f). Uniform heating of tissue and accelerated cooling are both beneficial for increased surgical efficiency and minimized thermal spread. A time-dependent plot of coagulation volume is shown in Fig. 8. At 3 s, the volume of coagulated tissue created with the standard monopolar pencil is calculated to be 27.2 mm3 . The

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IV. DISCUSSION AND CONCLUSION

Fig. 8. FEM-determine coagulation size (T > 50 ◦ C) over time for standard and TMS-monopolar procedures. At t = 3 s, the TMS monopolar device produces a coagulation with a volume 80% that of a standard monopolar device.

Fig. 9. Summary of experimental in vivo test results. Maximum mean temperatures recorded at radial distances from the electrode with 95% confidence interval bars (n = 7). Plot highlights that there is a statistically significant decrease in maximum temperatures due to the addition of the TMS at distances of 2 and 3 mm adjacent to the electrode. ∗No confidence interval available for this result. Note that the 1-mm measurement for the TMS Monopolar trial is not available due to the location of the cooling channel.

coagulation volume created with the TMS monopolar electrode is 21.7 mm3 or 80% of the size of the standard procedure. The uniform heating of the tissue by the TMS device initially creates a larger coagulation volume (as seen for times 0.75–1.5 s in Fig. 8). However, over time the cooling channels are able to minimize the coagulation volume. C. In Vivo Thermal Spread Reduction With Monopolar TMS Maximum tissue temperature readings for in vivo testing of each device performing EC procedures on a canine spleen are shown in Fig. 9. The TMS monopolar device exhibits maximum temperature decreases compared to the standard monopolar measurements of 39% and 37%, respectively, at 2.0 and 3.0 mm distances from the electrode. There were statistically different maximum temperatures due to the effect of the TMS. The temperatures with TMS are all below temperatures that have been associated with tissue coagulation (>50 ◦ C). Conversely, without the TMS the temperatures would damage tissue even 3.0 mm from the electrode. As in the ex vivo experiments, temperature measurements could not be measured at the 1.0 mm distance for the TMS monopolar device.

This study demonstrated the use of a TMS to reduce thermal spread and the accuracy of FEM to study tissue temperature and electrical voltage distributions during monopolar electrosurgical procedures. Monopolar EC devices were shown to lead to significant thermal spread that puts tissue at risk even 3 mm from the electrode edge in in vivo cases. A functioning FEM of a complex process was shown to be beneficial when improving the process and designing a new monopolar device. The FEM determined the affects to key outputs, such as maximum temperature and coagulation volume, with the addition of heat sinks into the model. A simple concept that increased the heat transfer of the resistively generated heat out of the tissue through adjacent cooling channels was shown to impact the thermal profile of the tissue both in magnitude (maximum temperature) as well as temporally through an increased rate of temperature decline. As thermal damage is temperature- and time-dependent, increasing the temperature decay rate should decrease thermal spread [8]. Maximum temperatures using the TMS monopolar device are comparable to the standard monopolar device. However, temperatures at 1.0, 2.0, and 3.0 mm adjacent to the electrode are seen to be lower, thus lowering the lateral collateral damage compared to the standard monopolar device. While including a thermal sink adjacent to the monopolar electrode would necessitate a larger size device, the development of an energy-based surgical instrument incorporating such a TMS will also lead to an improved patient’s post-operative quality of life by reducing the amount of collateral thermal damage to nearby tissue. Other energy-based technologies, such as ultrasonic and bipolar instruments, could also benefit from the inclusion of such a TMS. With the completion of this proof-of-concept study, monopolar TMS devices specific to targeted procedures will be developed and tested in future works involving this technology. Clinical variables such as the motion of the electrode, generator modality, and orientation and size of the cooling channels can each be tested in procedures where a TMS is believed to be beneficial, such as in natural abdominal plane dissections, vessel harvesting, entry point incisions, and nerve sparing procedures.

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Robert E. Dodde was born in Grand Rapids, MI, in 1975. He received the B.S. degree in biology and environmental science from Aquinas College, Grand Rapids, MI, in 1999, and the M.S. degree in mechanical engineering and the Ph.D. degree in biomedical engineering from University of Michigan, Ann Arbor, MI, in 2005 and 2011, respectively. He is currently a Postdoctoral Fellow at The University of Michigan, Ann Arbor, where he is involved in research in the fields of soft-tissue modeling and bioimpedance.

Jacob S. Gee was born in Salt Lake City, UT, in 1983. He received the B.S. degree in mechanical engineering from Kettering University, Flint, MI, in 2007, and the M.S. degree in biomedical engineering from University of Michigan, Ann Arbor, in 2008. He is currently an Associate Design Engineer at Ethicon Endo-Surgery, Cincinnati, OH, where he is involved in medical device design.

James D. Geiger was born in Toledo, OH, in 1960. He received the B.S. degree in biology from The University of Michigan, Ann Arbor, in 1983, and the M.D. degree from Case Western Reserve University, Cleveland, OH, in 1987. He is currently a Professor of Surgery at The University of Michigan Medical School, Ann Arbor, where he is involved in research in the fields of electrosurgery and medical devices.

Albert J. Shih was born in Tainan, Taiwan. He received the B.S. and M.S. degree in mechanical engineering from National Cheng Kung University, Tainan, Taiwan, in 1984 and 1986, respectively, and the Ph.D. degree from Purdue University, West Lafayette, IN, in 1991. He is currently a Professor of mechanical engineering and biomedical engineering at The University of Michigan, Ann Arbor, where he is involved in research in the fields of design and manufacturing of medical devices.