Implantable Drug Delivery Chips Enabled by Radio-Controlled … · 2015-10-09 · Implantable Drug Delivery Chips Enabled by ... administration. However, this constant release profile

1Recent Trends in Targeted Drug Delivery | www.smgebooks.comCopyright Takahata K.This book chapter is open access distributed under the Creative Commons Attribution 4.0 International License, which allows users to download, copy and build upon published articles even for commercial purposes, as long as the author and publisher are properly credited.

Implantable Drug Delivery Chips Enabled by Radio-Controlled Smart Microactuators

ABSTRACTMicro-electro-mechanical systems for drug delivery applications have attracted significant

interests that have led to extensive investigations. This chapter discusses approaches to realizing implantable micro devices that can be wirelessly controlled for localized and precise drug release using external radiofrequency (RF) fields along with two different device prototypes developed to incorporate radio-controllable active micro valves and pump. Both devices, packaged in the form of chips with approximate sizes of 10×10×2 mm3 or less, are wirelessly operated via RF-to-thermal energy conversion using inductor-capacitor circuits that serve as wireless resonant heaters, controlling temperature-sensitive microstructures integrated within the devices for valving and pumping. The first device to be discussed utilizes photo-patterned microstructures of a thermoresponsive hydrogel as active microvalves that are integrated with the wireless heater for controlled drug release via diffusion. A novel wireless shape-memory-alloy actuator with a built-in resonant heater is developed and embedded in the other device to operate a positive displacement pump, demonstrating forced and temporal release when the actuator is triggered by an external RF field. This second device is measured to exhibit a release volume of 219 nL/pump, suggesting that the device’s reservoir has a capacity of ~350 individual ejections. These prototype chips validate the effectiveness of wireless RF powering for fully controlled, long-lasting drug delivery, a key step towards enabling patient-tailored, targeted local drug delivery through highly miniaturized implants.

Jeffrey Fong and Kenichi Takahata*Department of Electrical and Computer Engineering, University of British Columbia, Canada

*Corresponding author: Kenichi Takahata, Department of Electrical and Computer Engi-neering, University of British Columbia, 2332 Main Mall, Vancouver, BC V6T 1Z4, Canada, Email: [email protected]

Published Date: May 10, 2015

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Keywords: Drug delivery; Micro-electro-mechanical systems; Implantable devices; Wireless; Resonant heating; Hydrogels; Shape memory alloys; Microvalves; Micropumps

INTRODUCTIONTraditionally, the most common methods of drug delivery include oral, topical, inhalation,

and injection. Besides topical application where the effect is generally localized, these methods typically utilize systemic administration, meaning that the medication is delivered through the circulatory system and affects the whole body. In systemic delivery, high dosages are required at the point of entry in order for the desired therapeutic level to be reached at the targeted location. This undesirable situation is necessary not only because the drug is diluted throughout the body but also because of the first pass effect for oral delivery and the diffusion barriers of certain tissue [1]. Although drugs can be tailored to target specific organs or cells, serious systemic side effects may still occur, damaging otherwise healthy tissue. For example, chemotherapy targets rapidly dividing cells such as cancer cells but also targets cells in bone marrow and hair follicles, which results in immunocompromised conditions and hair loss, respectively [2,3]. Localizing drug delivery to prevent the previously mentioned problems could be performed using implantable devices. Passive devices rely on diffusion to continuously release the drug at a controlled rate. Gliadel wafers, Zoladex implants, and Retisert are commercially available examples of such devices to treat gliobastoma, prostate cancer, and chronic non-infectious uveitis, respectively [4–6]. These devices can produce better control over drug concentration compared to bolus administration. However, this constant release profile may not be ideal for treating such issues.

Active, reservoir-based drug delivery devices (DDDs) offer the greatest potential in many applications because they have higher drug loading compared to infused polymer matrix devices, they are refillable, and the dosages can be tuned for individual patients. They could be used to treat a wide range of conditions including brain tumors, chronic pain syndromes, infectious abscesses, chronic eye diseases, and diabetes [7–9]. Benefits over traditional forms of drug delivery comprise of reducing the number of surgeries required, using higher drug concentrations at localized regions for greater efficacy, reducing the side effects, and not having to rely on patients to stay on top of their medication. An active device with dimensions in the centimeter range containing the drug delivering mechanism, batteries for powering, and complex circuitry for operation management may be highly invasive and pose biocompatibility issues as an implant. Advances in microfabrication have resulted in significant interest towards building DDDs with micro-electro-mechanical systems (MEMS) technology [10–13]. These active devices with small form factors could be implanted directly at the targeted site with minimally invasive surgery. Some devices have been developed to selectively release individual drug filled reservoirs from an array of micro-reservoirs using electrothermal [14–16] or electrochemical [17–19] mechanisms. Each reservoir is sealed with a metal membrane and can be opened destructively by applying electrical power. The electrothermal mechanism is being pursued by the company MicroCHIPS for several different applications including contraception, osteoporosis, and multiple sclerosis [20]. Its osteoporosis


application was the first wirelessly controlled MEMS-based DDD to be tested in humans [7]. For implantable devices, the elimination of a wired interface is essential for practical usage. Wireless MEMS drug delivery devices have been reported [7,15,16,21,22]; these devices operate with active control circuitry for release triggering and control, as well as batteries in many cases to power the circuitry, making them bulky and more invasive. Using passive operation mechanisms for release control eliminates the need for active circuits and batteries and is therefore advantageous in terms of size/invasiveness, longevity, cost, and robustness of the implants.

Another approach based on MEMS is to use micromachined actuators as valves and/or pumps for the devices to enable non-destructive (reversible) drug release and dose control in a timely manner [23–27]. This type of device is promising for achieving effective, long-term therapies with temporal release patterns designed and adjusted for particular applications and patients. MEMS actuators have been utilized to develop active DDDs with a variety of mechanical micropumps [28], including electrostatic [29], piezoelectric [30], shape-memory alloy (SMA) [31], bimetallic [32], electromagnetic [33], and electrochemical [34]. Hydrogels have also been used for drug delivery vehicles [35,36]. Some hydrogels are capable of swelling and deswelling in response to environmental parameters such as heat. Poly (N-isopropylacrylamide), or PNIPAM [37,38] is a thermoresponsive hydrogel that exhibits a phase transition temperature called the lower critical solution temperature (LCST) above which they shrink and deswell the fluid. The LCST value can be modified using different material compositions of the hydrogel [37]. Responsive hydrogels offer promising opportunities to design valves and pumps for DDD applications.

This chapter discusses two approaches to producing implantable wireless DDDs that are functionalized with thermoresponsive “smart” microactuators. The wireless control of both microactuators is enabled through the same principle based on resonant heating induced using externally radiated radiofrequency (RF) electromagnetic fields. In one approach, microstructures of a thermoresponsive hydrogel are integrated with an RF resonant heater to serve as active microvalves that are wirelessly opened/closed to control the diffusion of the drug stored in the device’s reservoir to the surroundings. In the other approach, a novel wireless SMA cantilever actuator is developed and embedded in a device with microfluidic valves, achieving forced and directed drug release with higher controllability. The wireless device architecture based on the use of smart actuators, adopted in both approaches, removes the need for control circuitry and power sources within the implants, exploiting the advantages noted earlier. These devices are made of biocompatible materials in the form of chips with approximate sizes of 10 mm or less.

OPERATION METHOD AND APPROACHESThe DDDs developed utilize frequency-dependent induction heating to control the actuation

of thermoresponsive materials/structures. When an inductor-capacitor (LC) resonant tank is exposed to an ac electromagnetic field, ac current is generated due to an electromotive force induced by the field. When the frequency of the field, fM, matches the resonant frequency of


the LC circuit ( 1/ 2Rf LCπ= ), where L and C are the inductance and capacitance of the circuit, respectively), the reactance of the inductor and the capacitor cancels out and the power consumed by the circuit can be simply expressed as 2 /P v R= , where R is the resistance of the circuit and v is the induced voltage. This results in maximum power consumption and conversion into thermal energy [39]. The circuit essentially functions as a wirelessly controllable heater that can be switched on simply by tuning fM, rather than the field intensity. This mechanism can potentially be an accurate and reliable means of control for micromachined thermal actuators as the intensity of the field radiated to the implant can vary significantly depending on the device’s location, i.e., the field’s attenuation level. Another important feature provided by this technique is that the field strength necessary to produce a certain amount of heat is much less than that required in non-resonant induction heating [40]. This feature is favorable in terms of achieving reduced doses of electromagnetic radiation to the body.

In the first DDD to be discussed, PNIPAM microvalves are actuated with a separate resonant heater in order to control the drug release through micro holes created in a reservoir wall (Figure 1A). The device is constructed so that a rise in temperature above the LCST occurs only when the field frequency is aligned to fR of the heater for frequency-controlled actuation of the microvalves (Figure 1C).The PNIPAM microstructures serve as soft valves that are expected to achieve more robust sealing. A photosensitive PNIPAM is used to lithographically form the microvalve structures on the heater circuit, which are designed to plug the release holes in their inactive mode. To initiate drug release, the temperature is brought above the LCST by activating the wireless heater through the field-frequency tuning method; this causes shrinkage of the hydrogel microvalves, unplugging the release holes through which the drug diffuses out from the reservoir. The generated heat also contributes to enhancing the diffusion of the drug. Shifting fM away from fR deactivates the heater and the microvalve, closing the release holes, and terminating the drug release.

The same resonant heating principle is adopted in the other developed DDD to drive a SMA actuator for active pumping of the drug from the reservoir (Figure 1B). In this device, the SMA itself is micromachined to form an LC tank to function as a cantilever actuator with a built-in RF receiver or wireless heater – when the SMA structure is excited with RF and heated to exceed its threshold temperature, the austenite phase temperature (Ta), the actuator is activated for pumping and the drug stored in the device’s integrated reservoir is released.


Figure 1: Conceptual diagrams of the RF powered, implantable drug delivery devices equipped with (A) thermoresponsive hydrogel microvalves and (B) SMA-based actuator/pump (adapted

from Ref. [41] with permission from The Royal Society of Chemistry) both operated by resonant heating that triggers the actuation of the hydrogel valves or the pump using tuned RF

electromagnetic fields applied externally as illustrated in (C).

EXPERIMENTAL RESULTSDDD Chip with Active Hydrogel Microvalves [42]

Design and fabricated prototype

Figure 2A shows asample design of the wireless LC heater circuit that has a 6×5-mm2 spiral coil with a theoretical fR of 94 MHz. The heater circuit is fabricated on a polyimide film using a planar microfabrication process. The drug reservoir is created by bonding a thick polyimide component (Figure 2B) that has a reservoir cavity to the planar heater circuit, forming an enclosed reservoir whose bottom surface is occupied by the coil heater. Prior to the bonding, release holes are created in the thinned top wall of the cavity component. This thin wall with the release holes is used as a photo mask to implement selective polymerization of the PNIPAM solution injected into the reservoir. This process forms well-defined hydrogel microvalve structures that are self-aligned to the release holes. Details of the fabrication process can be found in [42]. After all the fabrication steps are completed, the reservoir is filled with a liquid-phase drug through one of the holes reserved for this filling, followed by the sealing of this hole. The fabricated components and the final device are shown in Figure 2C-2E.


Figure 2: (A) A layout of the wireless resonant heater circuit; (B) top and side views of the design of the drug reservoir component;(C) fabricated sample LC heater circuit with photo-

patterned PNIPAM microvalve structures; (D) polyimide reservoir component with the cavity and the release holes (inset image shows a close-up of one of the release holes); (E) fabricated

wireless DDD chip whose release holes are plugged with the PNIPAM microvalves formed inside the reservoir. With kind permission of Springer Science + Business Media [42].

Characteristics of the wireless heater and PNIPAM valve

The measured thermal response of a fabricated wireless heater with a 5-mm-sized spiral coil (fR = 96 MHz) to periodic excitations with different fM is shown in Figure 3A, displaying quick responses in heat generation to the presence of a field – approximately 80% of the total temperature rise (e.g., 19.2°C at fM = 96 MHz) was completed in 15 seconds upon excitation. It also shows that the highest temperature was achieved when fM = fR. This effect is clearly seen in


another measurement result in Figure 3B obtained using a heater with a 10-mm-sized coil (fR = 35 MHz), showing a strong temperature peak approximately at the device’s fR. Figure 3B also shows the frequency dependence of the lateral dimension of a fully swelled PNIPAM structure patterned on the coil area. It can be seen that the piece shrunk down to 62% of the initial swelled size at the device’s fR and that the obvious shrinkage occurred when the heater temperature exceeded around 30°C, the LCST of the particular hydrogel used. The tested heater-hydrogel components showed an active frequency range for hydrogel actuation of approximately ±1 MHz at their resonant frequencies.

Figure 3: (A) Thermal response of a fabricated resonant heater with fR = 96 MHz to temporal wireless excitation with varying frequencies at a constant excitation power; (B) frequency

dependence of the heater temperature and the size of the PNIPAM hydrogel photo-patterned on the heater. With kind permission of Springer Science+Business Media [42].


Release tests

The release performance was tested using fluoresce in, a fluorescent dye (absorption maximum at 494 nm, emission maximum at 521 nm in water), for a device with fR =65 MHz. A fluorescent microscope was used to characterize the concentration of fluoresce in released from the device. Prior to the release test, leak tests were performed. For this, the device fully filled with fluoresce in was immersed in deionized (DI) water (total amount ~7.8 cc) and the fluorescence intensity was measured every 7 minutes over 12 hours at a location approximately 400 µm away from one of the release holes while no external field was present. As shown in the measurement result in Figure 4A, no distinct increase of the intensity was detected for this period. The intensity level observed in this leak test was almost identical to the background level measured with fresh DI water (corresponding to the first dot at the time of zero) before immersing the device. The release test was conducted as follows: after leaving the device in DI water with no external field for 2 minutes, an external field at fM =65 MHz was turned on for ~1 minute, then fM was shifted and kept at 20 MHz for ~2 minutes; this fM cycle of tuning in and out of 65 MHz was repeated while measuring the intensity with the same method used in the leak test. The RF output power was kept constant (at ~800 mW) during the entire period that the external field was present. Figure 4B shows a measured result from this release test. In the graph, the initial flat region corresponds to the period without the external field. The flat region was followed by periodic peaks that appeared when fM was tuned to 65 MHz. It can be seen that as fM was shifted to 20 MHz, the intensity dropped and returned to the base value. It is understandable that this decrease of intensity or dye concentration is because once the dye release was terminated, the released dye present at the measurement location quickly diffused away. Fluorescence images of the release hole used for the measurement are shown in Figure 5. The bright region in Figure 5B, 5D, and 5F indicates the dye that diffused out of the reservoir through a gap created between the release hole and the hydrogel microvalve due to its actuation at fM = 65 MHz. In Figure 5C and 5E, the bright region disappeared as shifting fM to 20 MHz caused a temperature drop that led to swelling of the hydrogel microvalve, closing off the release hole and terminating the dye release.


Figure 4: Measured fluorescence intensities observed with a fluoresce in-loaded device in DI water: (A) a leak test result measured with no external field; (B) a result of wireless temporal

release control using the frequency tuning technique, showing periodic peaks of intensity led by tuning fM in and out of 65 MHz (the letters in the graph correspond to the images in Figure 5).

With kind permission of Springer Science+Business Media [42].

Figure 5: Images of the release hole (with an angled view) captured during the temporal release test shown in Figure 4(B) using the fluorescent microscope, demonstrating dye release (in the bright region in (B), (D), and (F)) in response to field frequency tuning to 65 MHz. The timings

when the images were captured are indicated in Figure 4(B) with the corresponding letters. With kind permission of Springer Science+Business Media [42].


DDD Chip with SMA-Based Micropump [41]

Design and fabricated prototype

The second DDD was developed with a diaphragm pump containing two out-of-plane micro check valves, all made with thin films of Parylene C, a biocompatible polymer [43]. One microfluidic channel is used to carry the drug from the reservoir to the inlet valve of the pump chamber and another channel to carry it from the outlet valve to outside the device. Pumping in the device is achieved by a wireless SMA actuator designed for the targeted DDD. This actuator is based on a biocompatible nickel-titanium SMA, also known as Nitinol, which is designed to have a rectangular spiral-coil shape with an integrated capacitor to form an LC tank (Figure 1B). This Nitinol tank structure serves as a frequency-sensitive wireless heater activated by an external RF electromagnetic field. The stress layers of compressive SiO2 film are locally patterned on the coil so that one longitudinal end of the coil is lifted up at room temperature while the other end (the capacitor side) is fixed on the substrate of the device. Similar to the previous device, the drug release is controlled by wirelessly resonating the Nitinol LC-tank heater, which induces an RF electromotive force in the circuit and heats up most effectively as fM is tuned to fR of the circuit. When the temperature of the excited Nitinol coil exceeds the austenite phase temperature, Ta, the bent Nitinol coil returns to its remembered flat shape, causing cantilever-like actuation that squeezes the pump chamber present between the cantilever’s free end and the substrate to eject the drug out the nozzle through the outlet valve. As the RF field is tuned to a frequency outside of the actuator’s active range (Figure 1C) or simply turned off, the actuator cools down below Ta. This causes the Nitinol to return back to the cold (marten site phase) state where the material becomes compliant and the stress layer deforms the Nitinol structure to pull the free end up once again. By physically coupling this end of the actuator with the diaphragm pump chamber, the pull-up motion causes negative pressure inside the chamber that brings the drug in from the reservoir through the inlet valve. Therefore, repetitive activations of the actuator with these mechanisms lead to full pumping cycles. Because the actuator itself serves as an RF receiver with a built-in self-heating function, unlike the previous device, this device does not require a separate receiver antenna, minimizing the space allocation for the actuator while maximizing that for the drug reservoir for a given device size.

The design of this DDD is shown in Figure 6. The entire device is built around the Nitinol actuator (with a size of 9.2×2.7 mm2) in the form of a ~10×10×2-mm3 chip. To maximize drug loading, all space except for the one for the actuator was set aside for the reservoir, leading to a loading capacity of 76 μL. The diaphragm pump chamber was made to be 1.75-mm wide, 2.94-mm long, and ~170-µm tall. The Nitinol-coil cantilever produces a cold-state displacement of 215 μm. The mechanical coupling between the cantilever’s free end and the chamber’s top surface is made by inserting and bonding a polyimide spacer between them. The exterior walls of the DDD encase the entire chip and the interior wall separates the fluid in the reservoir from contacting the actuator. Because of the interior wall, a channel underneath it is required to connect the


reservoir to the pump chamber. The channels, with the smallest width being 100 μm, are created by wet etching through a 127-µm-thick sheet of polyimide. The top of the channels are sealed with Parylene C and the bottom side by bonding on another sheet of polyimide. A 305-μm-thick polyimide piece is used to seal the top of the chip. A 0.8×0.8 mm2 square hole is created in the top cover for refilling purposes (plugged during operation). All components in contact with the human body and the drugs are made of polyimide or Parylene C, each of which is a commonly used biocompatible material. The microfluidic components are constructed through photolithography-based processing, including micromachining of the Parylene C and polyimide layers for the formation of the check valves, fluidic channels, and pump chamber. The packaging case is prepared by mechanical micromachining of 1.52-mm-thick and 305-μm-thick polyimide sheets. Microfabrication of the spiral-coil cantilever actuator is achieved through bulk micromachining of a Nitinol sheet (Alloy M, Ta = 65°C, Memry Co., CT, USA) in combination with photolithography processes for monolithic integration of a thin-film capacitor to form the LC-tank circuit. The SiO2

stress layers are also lithographically defined on selected parts of the Nitinol-coil structure to cause its cold-state deformation as described earlier. Details of the fabrication processes are given in the supplementary information of [41]. The fabricated device components and the completed chip are shown in Figure 7. The fabricated LC-tank heater/actuator was measured to have an inductance, capacitance, and fR of approximately 75 nH, 9.9 pF, and 185 MHz.

Thermal characteristics of the fabricated prototype

Wireless heating of the Nitinol actuator integrated in the DDD chip was performed using a tuned RF field produced with output powers of up to 1.1 W. To experimentally assess the thermal impact of the actuator’s heating on the surrounding materials and the contained drug agent, the reservoir and all the channels (including the pump chamber) of the device were filled with distilled water.


Figure 6: (A) 3D model and dimensions of the DDD chip (excluding the SMA actuator) and a close-up view of the valves; (B) a cross sectional view of the chip showing structural layers used for its construction. Adapted from Ref. [41] with permission from The Royal Society of

Chemistry.

Figure 7: (A) Front side (left) and back side (right) of the microfluidic channels and Parylene pump chamber created in/on the polyimide substrate; (B) (left) inlet valve with the valve disc on top, (middle) outlet valve with the valve seat on top, and (right) scanning electron microscope (SEM) image of an outlet valve formed on the polyimide channel structure; (C) completed DDD chip without(left) and with(right) the top cover; (D) close-up SEM image of

the integrated Nitinol actuator coupled with the pump chamber. Adapted from Ref. [41] with permission from The Royal Society of Chemistry.


An infrared (IR) image of the device (without the top cover) captured after 5 seconds of excitation is shown in Figure 8A (this time period was selected to be the same as in the wireless release tests described later, which was revealed to be enough time to complete the release of a single pump). The temperature reading at spot-1 (77.1°C) indicates that the center of the coil exceeded Ta (= 65°C) which is necessary to actuate the Nitinol-coil cantilever. Spots-2 and -3 show the temperatures of water in the reservoir and near the outlet nozzle, respectively, suggesting that even the outlet region adjacent to (~1.0-mm away from) the actuator was still close to room temperature. Exterior heating is another important factor in terms of the device design and packaging. Any heating on the exterior surfaces should be within a level that does not cause thermal damage to tissue (~43°C or less [44]). In light of this aspect, temperature distribution on the polyimide casing of a fully packaged chip was monitored through thermal imaging while exciting the chip wirelessly under the same conditions. Figure 8B displays an IR image of the chip’s backside after the 5-second activation, showing that the maximum temperature over the entire surface of the monitored side was 42.3°C (at spot-1 in the image). The location of this highest temperature is reasonable as it is where the anchor of the actuator is located and thus should experience the greatest level of heat flux from the actuator. This temperature is around the safe-level threshold and may need to be lowered. This could be achieved by simply using a Nitinol composition with a lower Ta and/or by further optimization of the device packaging.

Figure 8: IR images of (A) the integrated actuator and the surrounding structures of an open chip excited with 850-mW RF output power showing temperature readings of the actuator’s innermost coil turn (spot-1), water in the reservoir (spot-2), and water at the outlet (spot-3), and (B) the fully packaged chip excited in the same manner showing the temperature

distribution of the polyimide casing with a highest temperature of 42.3°C appearing around the actuator’s anchor region. The excitation antenna was located beneath the device in both cases.

Adapted from Ref. [41] with permission from The Royal Society of Chemistry.


Release tests

As a preliminary test of wireless pumping and release for visual verification, a fabricated chip, whose reservoir and pump chamber were filled/primed with a colored dye solution, was immersed in distilled water (gently agitated with a magnetic stirrer) and excited with a tuned RF field radiated from an antenna located outside of the water-filled container. The RF signal was repeatedly activated with a cycle of 5-seconds on and 13-seconds off. Ejection of the dye solution into the surrounding water was clearly observable (Figure 9A). With no agitation, cumulative release of the dye near the outlet nozzle was evident (Figure 9B). These tests qualitatively validated the device’s functionality in a liquid ambient.

For quantitative analysis of wireless release, another chip filled with an electrically conductive test agent (30-35% nitric acid) was activated to eject it into the water environment while tracking the electrical conductivity of the water. The same RF power and repetition cycle were used, except for an off time of 85 seconds (for uniform dispersion of the ejected agent and stabilization of the ambient conductivity). In this test, the position of the chip was adjusted to bring its outlet nozzle slightly above the water level but close enough to it so that a droplet coming out from the nozzle would be pulled away into the water. Figure 10A plots the cumulative amount of released agent (in moles, converted from the recorded conductivity data) as a function of time along with its incremental molar amount. The volume released with each pump was analyzed from the result in Figure 10A and plotted in Figure 10B. The result clearly demonstrates that tuned RF activation of the chip was effective in performing incremental release of the agent in a fully wireless manner. The graph in Figure 10B indicates a good stability of release volume under temporal operation (except the first activation, which might have been affected by the initial priming condition in the pump and the outlet channel). The average volume per pump was determined to be 219 nL. This measured value, in addition to the reservoir capacity (76µL), suggests that the chip can be pumped ~350 times before requiring a refill. As an example of its clinical relevance, human parathyroid hormone fragment (1-34)


Figure 9: Images from the preliminary test of controlled wireless release through a fabricated DDD chip, with colored dye, immersed in distilled water (A) with and (B) without flow. The chip was periodically excited by radiating an RF field (tuned to fR of the device) with an output power

of 1.1 W.Adapted from Ref. [41] with permission from The Royal Society of Chemistry.

(hPTH (1-34)) is used for anabolic treatment of osteoporosis and is shown to reduce the occurrence of bone fractures when drug doses of 20-40µg are administered daily [45,46]. With 40-µg daily doses and a solubility of 80 mg/mL [7], the developed DDD chip could be used for 5 months before refilling; in comparison with the Micro CHIPS’s device reported for this application [7], the above period is much longer than the current Micro CHIPS’s capacity of 20 days and is provided by a DDD with a >80× smaller device volume. It should also be noted that the capacity of the developed DDD chip can be easily increased by simply enlarging the drug reservoir (e.g., doubling its height) while maintaining its miniaturized form with a lateral size of ~10 mm.

The temporal release test was also conducted with another chip immersed in water (Figure 10C


and 10D). The volume of agent released by each pump (Figure 10D) shows an apparent tendency of gradual decrease followed by saturation with subsequent pumping. This outcome seems to be led by the condition that a small amount of water from the surroundings was drawn back into the pump chamber during operation, as suggested by Figure 10C that indicates a consistent decreasing trend in the ejected molar amount (or in the conductivity gain of the ambient). This decrease in conductivity gain was then used as a compensation factor in the calculation of the ejected volume. The result, also shown in Figure 10D with a different color, clearly suggests that the actual volume of release was relatively constant (around 400-450 nL/pump; this level is larger than the case in Figure 10B as a different device with a taller pump chamber was used). A possible reason for this backflow is that the retracting (upward) force generated by the SiO2 stress layer on the Nitinol structure was not sufficiently large to expand the pump chamber fast enough; if the pump chamber expands without providing the required pressure difference, the outlet valve could be slightly open and allow for backflow.

Figure 10: (A) Cumulative and incremental molar amounts of released test agent per pump quantified from the recorded conductivity change of the surrounding water by repeated

ejections from the RF-excited DDD chip without submerging its outlet nozzle;(B) released volume per pump calculated from the result in (A);(C) and (D) show the same set of data

obtained with a chip entirely immersed in water; the calculated release volume in (D) also shows the case compensated with the effect of consistent reduction in the molar amount

observed in (C) to estimate the actual volume released per pump. Adapted from Ref. [41] with permission from The Royal Society of Chemistry.


If this is the case, the test agent coming from the reservoir is mixed with the surrounding fluid coming through the outlet valve, and thus the subsequent pump would be diluted. This hypothesis is supported by the experimental result from the previous test that showed a relatively consistent release volume over repetitive pumping when the device was positioned right above the water level, hence any backflow was physically prevented. Nevertheless, the experiment indeed demonstrates that controlled temporal ejection from the drug reservoir could be radio-controlled with the developed prototype in liquid via resonant RF power transfer to the Nitinol actuator/pump integrated in the chip. To prevent the backflow, the valves could use more compliant valve tethers or slanted tethers that apply a force on the valve disc against the seat to ensure valve closure [47]. Another potential method is to incorporate active valves such as the PNIPAM valves [42] used in the other DDD discussed earlier. These valves could be synchronously controlled with the Nitinol pump using the same RF excitation principle by assigning them wireless heaters with different resonant frequencies and by modulating the RF field corresponding to these frequencies [48] – this could offer not only complete outlet closure but also full controllability of valve operation within the device.

CONCLUSIONSTwo different implantable microdevices operated by wireless RF powering have been

developed towards realizing targeted local drug delivery. Both prototypes utilized resonant tank circuits that heated up when exposed to RF electromagnetic fields with field frequencies matching their resonant frequencies. One of the devices used the generated heat to actuate PNIPAM hydrogel microvalves to allow the fluid inside the reservoir to diffuse out into the surroundings. The other device used the heat to change the phase of a SMA actuator in order to compress a flexible pump chamber and force the drug out of the device. Being functionalized by these responsive smart actuators activated externally, neither device required any control circuitry or battery in their construction, a feature inherently advantageous for miniaturization of implantable DDDs. The wireless delivery functions of both devices were experimentally demonstrated to show temporal release control using microfabricated prototypes loaded with test agents. The frequency sensitivity of wireless heat generation permits release control by field frequency tuning. This ability will potentially enable precise and safe radio control of temporal drug delivery while offering immunity to ambient/irrelevant RF fields. It will also provide an opportunity to realize more complex fluidic control with microvalves and pumps that are integrated together within an implant and synchronously controlled by frequency modulation of the external field.

ACKNOWLEDGMENTThis work was partially supported by the Natural Sciences and Engineering Research Council

of Canada, the Canada Foundation for Innovation, the British Columbia Knowledge Development Fund, and the Canadian Microelectronics Corporation. K Takahata is supported by the Canada Research Chairs program.


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