Supplying energy to tetherless milli-and micro-scale robots is challenging. Energy storage is a function of volume, which decreases with the inverse cube of body length. For this reason, magnetism is the primary source of propulsion for milli-and micro-scale tetherless robots [1]- [12]. This propulsion is often achieved by making robots with spiral or helical shapes, embedding a permanent magnet, and applying a rotating magnetic field. This rotation is primarily designed for propulsion, but also turns the robot into a rotating tool. The tip can be utilized for abrasion, drilling, or cutting [8]- [13]. Still, the power that can be delivered is limited. This paper presents a complimentary means to add functionality to milli-scale robots by using induction heating. While induction heating may be most popularly known in higher-end stove tops, it is a mature technology used in diverse fields. This work uses a commercial, off-the-shelf induction heater to heat a tetherless robot and investigates related implementation issues. These devices may be used continuously/repeatedly, because energy for treatment and propulsion is provided externally.
Heat is used in various forms of medical treatment. For example, cold coagulation is a gynecological technique used to treat cervical intraepithelial neoplasia, where a probe is heated to 90 C to destroy abnormal cells [14]. Heat is also used in laser ablation to destroy undesirable cells such as cancer cells [15]. Cardiac ablation uses localized heat delivered by a catheter and radio frequency energy to create scars in the heart to treat arrhythmias [16]. The white cubical workspace is the magnetic manipulator. An inductive heater is mounted on a UR-3 arm, and positioned to administer heating. (Right) The instrumented swimmer with a 30-gram section of chicken thigh and skin. The swimmer and tissue were immersed in water in a test tube, and the swimmer tip cauterized the tissue when the inductive heater with a helical coil was applied for 90 seconds. See video overview at https://youtu.be/I0tHE13myHE.
However, these forms of heat delivery are tethered. Using tethers such as catheters comes with known detriments. Catheters rub against the walls of the arteries and can detach plaque (calcified fat deposit), which can travel within the bloodstream and cause a stroke [17]. In addition, steering guidewires can be challenging as there is friction which can cause the wire to buckle. A tetherless robotic agent may solve these problems.
Hyperthermia consists of increasing the temperature of cancer tissue to around 40 to 48 C to enhance the effect of other therapies [18]. There are two predominant ways of heating tissue for hyperthermia therapy: photothermal therapy and magnetic hyperthermia. In photothermal therapy, lightsensitive nanoparticles are injected inside a cancer tumor. Infrared light is then applied to increase the temperature of the nanoparticles and the surrounding tissue. In magnetic hyperthermia, magnetic nanoparticles are injected into a tumor and an alternating magnetic field is applied to increase the temperature of the nanoparticles. The increase in temperature is mainly produced by magnetic hysteresis and relaxation losses [19].
conductive and ferromagnetic tip epoxy resin swimmer's body permanent magnets In this paper, we propose using Miniature Magnetic Robotic Swimmers (MMRSs), each carrying a ferromagnetic material that can be heated wirelessly using induction. This preliminary work uses a 7⇥-scaled model of our magnet swimmer from [20]. Figure 1 shows a prototype device in which a 35 mm long magnetic swimmer is heated by a commercial induction heater mounted on a UR-3 robot. The device can denature chicken tissue that is immersed in water with the swimmer in a test tube. The relatively large size limits use to the GI tract, but the design could be scaled for the circulatory system as in [20].
The system consists of a Miniature Magnetic Rotating Swimmer (MMRS) and a magnetic manipulator. The MMRS is the end-effector of the robot, and its position is controlled by the magnetic manipulator, a set of six electromagnets and a computer vision system described in [21]. The induction heating is supplied by a commercial off-the-shelf induction heater (Hot Rod H7E) mounted on a UR-3 robot arm, and controlled by a LabVIEW cRIO.
The proposed MMRS consists of four parts: a nonmagnetic body, a permanent magnet, a ferromagnetic tip, and a heatresistant insulator.
The body is made using an SLA printer. It includes propeller fins. It has one cavity in the front to receive the ferromagnetic tip and one cavity in the back to receive a permanent magnet.
The permanent magnet is rectangular with a length of 9 mm and a 3 mm ⇥ 3 mm cross section. This magnet interacts with an external rotating magnetic field to produce a torque on the swimmer and makes it rotate. The propeller fins convert rotational movement into propulsive force.
The ferromagnetic tip is the part that heats up via induction (the magnets produce negligible heat). This tip must be electrically conductive. An external coil applies an AC magnetic field that induces currents in the tip and produces heat via Joule losses. The coil also heats up due to the large amount of current that circulates in it. The tip is attached to the swimmer using a heat resistant epoxy resin (J-B Weld HighHeat™). This resin acts as a thermal insulator and allows reaching higher temperatures at the tip by shielding the majority of the insert from the cooling effects of the surrounding liquid. Figure 3 show the dominant heat flows.
A model of the inductive effect was created using the software Finite Element Method Magnetics (FEMM) [22]. The model is axisymmetric (see Figure 4). The heated element simulated is an iron cylinder with a length of 18 mm and a diameter 1.2 mm. Two different induction coils with pancake geometries were simulated. The first one (Pancake Coil A) has 5 turns, an external diameter of 52 mm, an internal diameter of 12 mm and a thickness of 3 mm. The second simulated coil (Pancake Coil B) has the same internal diameter, the same thickness, an outer diameter of 92 mm and 10 turns. Coil C is a helical coil with a height of 12 mm and four turns. Coil A and coil C correspond to the coils used in our experiments (section V) while coil B was simulated to test how increasing the dimensions of the pancake size and evaluating heat transfer could be applied in a medical lab.
The outer boundary of the simulation is a sphere having a diameter of 400 mm. A Dirichlet boundary condition was set with the value of the potential set to 0. This condition keeps the magnetic flux from crossing the boundary and is valid if the boundary is sufficiently far from the objects studied. This type of boundary typically needs to be at least 5 times larger than the objects studied.
The amplitude of the current through the coil was measured to be 150 A at a frequency of 40 kHz. These values were used in the FEMM models for analysis.
The power transfer was calculated as a function of the distance d between the coil and the element being heated. Results are shown in Figure 5. Coil A is able to produce significant amount of power for small values of d (8.5 W at d = 0 mm). The transferred power quickly decreases when the distance increases. At a distance of 10 mm, only 1.5 W is transferred, which corresponds to a decrease of 82%. Coil B, which has a larger diameter, has a decrease of power of 0 5 10 15 20 25 30 35 40 45 0 5 1 0 1 5 2 0 Power transferred [W] Distance, d [mm] Coil A: Pancake Coil B: BIG Pancake + Coil A: Pancake ID = 12mm, OD = 52mm, Turns = 5 X Coil B: Pancake ID = 12mm, OD = 92mm, Turns = 10 O Coil C: Helical Height = 12mm, Turns = 4 Fig. 5.
Power as a function of distance calculated using FEMM for three different coil geometries. Coils A and C were used experimentally. Simulations use a 18 mm long, 1.2 mm diameter iron rod, with the coil driven at 40 kHz.
68% at a distance of 10 mm showing that a larger diameter coil has a greater range. Coil C is helical and has a decrease in power of 63% at a distance of 10 mm, demonstrating a tradeoff between coil geometry and power transfer efficiency.
Another plot is provided in Figure 6 where the power transferred is plotted as a function of the frequency of the current. This is shown for both pancake coil A and helical coil C. This plot shows that the power increases linearly when the frequency increases. The experimental system works at 40 kHz, a relatively low frequency compared to some commercial devices able to reach 1.2 MHz [23].
Figure 7 plots the power transferred as a function of the length of the iron rod. The value increases asymptotically with the length. The extremity of the rod opposed to the coil gets further away from the magnetic source as the length increases. As a result, less power is transferred to this section which explains the asymptotic behaviour.
Our setup is shown in Fig. 8 and our procedures were as follows. A flat-tip swimmer was placed inside a test tube with 35 g of water and energized for three minutes using either a pancake coil or a helical coil, then left to cool. For each trial of each experiment, the swimmer, the water, and the coil were brought to room temperature before starting. Temperature 0 5 10 15 20 25 30 0 0.5 1 1.5 2 2.5 0 1 0 2 0 3 0 4 0 Power transferred, coil C [W] Power transferred, coil A [W] Iron rod length [mm] + Coil A: Pancake O Coil C: Helical Fig. 7. Power as a function of the iron rod length calculated using FEMM for coils A and C at d =10 mm, for a 1.2 mm diameter iron rod. Coil driven at 40 kHz.
was measured at four locations in our setup using type-E thermocouples (Oega TT-E-36-200). The thermocouples (red circles in the illustration) were placed in the water, on the iron tip of the swimmer, taped on the outside surface of the test tube (facing the coil when the pancake coil was used), and wrapped around the coil of the inductive heater.
The four thermocouples were connected to a LabVIEW cRIO with a NI-9213 high-density thermocouple input module, and data was saved at 10 Hz.
The relationship between the energy supplied to the thermal swimmer and its proximity to the heating element was easier to study using the pancake coil. For these experiments, the inductive heater was fixed so that the center of the pancake coil was at a distance d from the center of the ferromagnetic tip of the swimmer. The distances tested were d = 12.5 + [0, 5, 10] mm to account for the radius of the 25 mm O.D. of the glass test tube. For the helical experiments, the coil was placed around the test tube at the height of the swimmer, with an 8 mm gap between the test tube and the coil.
ASTM F444-88 specifies that temperature should be kept below 49 C to avoid burns to tissue. The default setup with the induction heater violated this temperature, so we repeated these experiments, but with ceramic fiber insulation (2600F) around the test tube. This was trimmed to the nominal thickness desired, and can be modeled in a clinical setting as insulation between the patient and the induction coil. For the pancake coil, this insulation was friction fit into the plane within the wooden frame between the test tube and the coil. For the helical coil, this insulation was wrapped around the test tube.
As an alternate safety measure, we used forced convection employing a blower with airflow at a nominal speed of 5.7 m/s to cool the coil and test tube.
The experimental results are categorized according to the coil used, and subdivided by the separation distance and the methods used to control the radiated heat transferred to the test tube from the coil: {no insulation, ceramic insulation, forced convection}. In a surgical setting, the patient must be protected from undesired tissue heating [24]. Surprisingly, the ceramic insulation was counter-productive. In all tests using insulation, the insulation prevented the coil from passively air-cooling. This had a positive feedback effect on the coil temperature because increased temperature increased the resistance of the copper coil until the current limit of the inductive heater was reached. In one test with insulation, the coil heated to 656 C, causing it to glow red.
Six tests were conducted with the pancake coil, all shown in Fig. 9. In each test, the coil was the hottest, the next hottest was the test tube facing the coil, then the swimmer, and the
g Coil Test tube Swimmer Water ins. mm t s C t s C t s C t s C Pancake N 0 73 271 110 50 167 25 113 17 N 5 123 145 108 40 156 11 111 10 I 5 138 337 93 71 155 11 107 8 N 10 105 135 104 38 162 7 111 9 I 10 92 529 95 66 161 5 95 5 B 10 78 41 92 8 145 4 187 2 Helical N 8 125 350 128 76 10 73 160 10 I 8 130 167 134 24 9 73 155 14 B 8 76 34 133 7 9 63 134 7 TABLE I 10% TO 90% RISE TIME AND TOTAL CHANGE IN TEMPERATURE FOR THE NINE EXPERIMENTS IN FIGS. 9 AND 10. N STANDS FOR NO INSULATION, I FOR CERAMIC INSULATION, AND B FOR BLOWER WITH FORCED AIR. d = g + 12.5 MM RADIUS OF THE TEST TUBE.
water was the coolest. The swimmer responded quickly to the induction heating, but in all tests except with the blower, the convection from the coil to the test tube warmed the swimmer more than the induction heating.
The helical coil was more effective than the pancake coil at heating the swimmer. Three temperature sets of measurements were performed with the helical coil, and are summarized in Fig. 10 and Table I. With no insulation the coil heated to a steady state of 231 C in 125 s, while the swimmer reached a steady state 110 C in 10 s. The test tube temperature increased almost linearly to 53 C over the 180 s of actuation. This is too hot for patient safety.
Adding 8 mm of ceramic insulation made the problem worse, because as with the pancake coil, the insulation deprived the coil from the ability to passively cool on one side. The coil heated to a steady state of 461 C in 130 s, while the swimmer reached a steady state 110 C in 9 s. More dangerously, the test tube heated almost linearly to 110 C.
Using forced air convention was more effective, and the coil heated to a steady state of 64 C in 60 s, while the swimmer rapidly warmed to 93 C in 9 s, and the temperature steadily climbed to 102 C. The test tube peaked at 33 C, warming almost linearly the whole time, as did the water, which warmed to 32 C. Despite this, the pancake coil appeared to be more practical, with the flat shape allowing one to reach areas of the body that a helical coil may be too small to wrap around (such as the torso).
In water, the swimmer tip's maximum temperature is regulated to around 100 C as the nearby water boils. If the swimmer is not submerged in water, the swimmer rapidly heats up. Figure 11 shows an example with the swimmer heated by the helical coil for 15 s, long enough for the iron tip to glow red. It appears that if the tip is inserted into the correct tissue that high temperatures are also possible, as shown in Fig. 1, where the chicken skin and fat was charred black, even through the swimmer and chicken tissue were submerged in water.
To demonstrate that this method has the potential to apply enough heat to kill tumors, swimmers were used to leave burn marks on chicken pieces. The swimmers have differentshaped tips as shown in Fig. 2. In this case, the swimmers were lodged into chicken cubes underwater, and then heated up for 90 seconds using the inductive heater. Results are shown in Fig. 12.
The main advantage of using this system is the ability to apply hyperthermia multiple times in different places. This compares favorably to using a small swimmer to deliver medication, as the payload of a swimmer is limited. In contrast, a swimmer could be energized multiple times during the same procedure. Before After Fig. 12. Denaturing chicken submerged with swimmer underwater. For each experiment a 30 mm inner-diameter helical coil was placed around a 25 mm outer-diameter test tube containing a 30-gram cuboid of chicken thigh and 35 g of water. The coil was energized for 90 seconds and changed the tissue to opaque white. See Fig. 8 for the test setup. The middle chicken cuboid was least effected, with a small circular change in the lower right.
This paper presented a tetherless magnetic swimmer that can be heated wirelessly by induction. When heated using a helical coil, the swimmer in water heated to 100 C in around 10 seconds. This swimmer could be maneuvered magnetically and heated repeatedly without performance degradation. A potential use was demonstrated by denaturing chicken thigh tissue.
While these results are promising, there are challenges that still need to be addressed. The heating power decreases quickly as the distance between the coil and the swimmer increases. Cardiac ablation surgeries will require distances in the range of 5 to 10 cm. At this distance, the device presented could not transfer significant power. These tests are preliminary, and the system was not optimized.
Simulations showed that increasing the diameter of the induction coil could increase the range of the device. In addition, the power transferred increases when the frequency of the current in the induction coil increases. The presented system works at 40 kHz, a frequency that could be increased in future work. Systems working at 1.2 MHz are commercially available [23]. The helical coil had a higher efficiency than the pancake coil. Alternate coil geometries transfer energy with more efficiency, and selection will depend on surgical parameters [25]. Incorporating a cooling system into the coil, and optimizing materials will increase efficiency. Future work on safety and optimization would be key, with tests on limiting variability through higher frequencies, multiple conformal coils, or actively cooled coils.
Bellaire High School, TX USA kalyanigifford@gmail.com
University of Houston, TX USA {jgarciag, vjmontan, jleclerc, atbecker}@Central.uh.edu