UBC Theses and Dissertations

UBC Theses Logo

UBC Theses and Dissertations

A minimally-invasive MEMS drug delivery device for the treatment of prostate cancer Zachkani, Payam 2014

Your browser doesn't seem to have a PDF viewer, please download the PDF to view this item.

Notice for Google Chrome users:
If you are having trouble viewing or searching the PDF with Google Chrome, please download it here instead.

Item Metadata

Download

Media
24-ubc_2015_february_zachkani_payam.pdf [ 2.01MB ]
Metadata
JSON: 24-1.0135633.json
JSON-LD: 24-1.0135633-ld.json
RDF/XML (Pretty): 24-1.0135633-rdf.xml
RDF/JSON: 24-1.0135633-rdf.json
Turtle: 24-1.0135633-turtle.txt
N-Triples: 24-1.0135633-rdf-ntriples.txt
Original Record: 24-1.0135633-source.json
Full Text
24-1.0135633-fulltext.txt
Citation
24-1.0135633.ris

Full Text

A MINIMALLY-INVASIVE MEMS DRUG DELIVERY DEVICE FOR THE TREATMENT OF PROSTATE CANCER by  Payam Zachkani  B.Sc., University of Tehran, 2012  A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF  MASTER OF APPLIED SCIENCE in THE FACULTY OF GRADUATE AND POSTDOCTORAL STUDIES (Mechanical Engineering)  THE UNIVERSITY OF BRITISH COLUMBIA (Vancouver)  December 2014  © Payam Zachkani, 2014 ii  Abstract  We have developed a cylindrical shape magnetically-actuated MEMS drug delivery device for localized prostate cancer treatment. The device is small enough for implantation through a needle with minimally invasive procedures with potentially fewer side effects compared with full prostate removal. This method of implantation will be similar to brachytherapy, a standard procedure to implant radioactive seeds inside the prostate through a needle.  The drug delivery device consists of a drug reservoir, a PDMS membrane, a magnetic block and housing. Docetaxel (DTX), an anti-proliferative drug, is deposited in the reservoir in solid form. The reservoir is then filled with fluids to form a saturated drug solution. When an external magnetic field is applied, it attracts the magnetic block towards the positive field gradient and causes the membrane to deflect. As a result, DTX is discharged from the reservoir, through a laser-drilled aperture on the membrane and into the housing. The housing has a 10 mm long opening which allows the released drug to diffuse to the surrounding tissues while it would prevent the tissues from touching the thin membrane. We have achieved a 1.8 fold increase of the actuating distance and a 3.6 fold increase in the magnetic force compared to the state-of-the-art magnetically-actuated drug delivery devices under the same actuation parameters. We have also demonstrated device implantation with a needle into swine bladder tissue and successful drug release of the device in the tissue. iii  Preface  The research presented in this dissertation was carried out at the University of British Columbia under the supervision of Dr. Mu Chiao. This work was also presented in the following conference:  P. Zachkani, J. K. Jackson, F. N. Pirmoradi, H. M. Burt and M. Chiao, “A Minimally-Invasive MEMS Drug Delivery Device for the Treatment of Prostate Cancer,” Solid-State Sensors, Actuators and Microsystems, Hilton Head, South Carolina, USA, 2014.  All aspects of the work presented in this dissertation, including literature review, conceptual design, fabrication processes, device characterization, controlled drug release studies and tissue implantation was performed by the author of this thesis. Dr. Mu Chiao provided the author with editorial suggestions of this manuscript and his expert advice during the entire course of the project and helped the author during the conceptual design stage. Dr. Pirmoradi from Palo Alto Research Center (PARC) provided training and expert advice to the author. One figure in this thesis is used with her permission from her journal paper. Mr. Jackson from Dr. Burt’s group at the Faculty of Pharmaceutical Sciences played a pivotal role in developing the idea of using a magnetically-actuated device for prostate cancer. He provided the author with technical advice, editorial suggestions, equipment and materials. Dr. Burt provided the author with equipment. The author’s major contributions in this dissertation are as follows: 1. Developed a minimally-invasive magnetically-actuated drug delivery implant for the treatment of prostate cancer with a controlled-release profile on an on-demand basis. 2. Developed the conceptual design of the device and characterized the device in terms of magnetic and mechanical performance.  iv  3. Developed fabrication processes of two types of devices including an all PDMS device and a 3D printed device. 4. Demonstrated an improved performance of the device compared to the state-of-the-art by analytical and computational methods. 5. Obtained reproducible drug release profiles with consistent release rates. 6. Demonstrated device implantation into a swine bladder tissue through a needle and successful operation of the device in the tissue. v  Table of Contents  Abstract .......................................................................................................................................... ii Preface ........................................................................................................................................... iii Table of Contents ...........................................................................................................................v List of Tables ............................................................................................................................... vii List of Figures ............................................................................................................................. viii Acknowledgements .................................................................................................................... xiv Dedication .....................................................................................................................................xv Chapter 1: Introduction ................................................................................................................1 1.1 Prostate Cancer and Current Treatments ........................................................................ 1 1.2 Controlled Drug Delivery ............................................................................................... 2 1.2.1 Passive Drug Delivery Implants ................................................................................. 3 1.2.2 Active Drug Delivery Implants ................................................................................... 4 1.2.2.1 Reservoir-based Implants.................................................................................... 5 1.2.2.2 Micropumps ........................................................................................................ 6 1.2.3 Challenges of the Active Implants .............................................................................. 6 1.2.3.1 A Magnetically-actuated Battery-less MEMS Drug Delivery Device ............... 7 1.2.3.2 Minimally Invasive MEMS Drug Delivery Implant for the Treatment of Prostate Cancer ................................................................................................................. 10 1.3 Thesis Overview ........................................................................................................... 12 Chapter 2: Design, Fabrication and Characterization .............................................................14 2.1 Design ........................................................................................................................... 14 vi  2.2 Fabrication .................................................................................................................... 18 2.2.1 PDMS Device ........................................................................................................... 18 2.2.2 3D Printed Device ..................................................................................................... 22 2.3 Actuation Setup ............................................................................................................. 26 2.3.1 In-vitro Testing Setup ............................................................................................... 27 2.3.2 Ex-vivo Testing Setup............................................................................................... 28 2.4 Characterization ............................................................................................................ 29 2.4.1 Magnetic Force ......................................................................................................... 31 2.4.2 Membrane Deflection ............................................................................................... 33 2.4.3 Release Time ............................................................................................................. 35 2.4.4 Mixing Time ............................................................................................................. 38 Chapter 3: Results and Discussions............................................................................................40 3.1 MB Controlled Release ................................................................................................. 40 3.2 DTX Controlled Release ............................................................................................... 43 3.3 Tissue Implantation ....................................................................................................... 45 Chapter 4: Conclusions and Future Work ................................................................................47 4.1 Summary ....................................................................................................................... 47 4.2 Future Work .................................................................................................................. 51 References .....................................................................................................................................54 Appendices ....................................................................................................................................61 Appendix A : Actuation Setup .................................................................................................. 61 Appendix B : COMSOL Simulation Verification .................................................................... 62 vii  List of Tables  Table  2.1 Theoretical forces on the magnetic block and the corresponding deflections of the membrane. ..................................................................................................................................... 34 Table  2.2 Estimated release time in different magnetic fields. ..................................................... 38 Table  4.1 Comparison between the magnetic force and the actuation distance between the current device and the device from  [34]. ................................................................................................... 51 Table B.1 Comparison between the simulated deflection and the approximate analytical solution for the magnetic PDMS membrane in  [34] with Φ = 6mm and t = 40 μm. .................................. 62  viii  List of Figures Figure  1.1 Controlled drug delivery compared to conventional drug delivery. Image created based on the image from  [13]. ........................................................................................................ 3 Figure  1.2 Schematic design and principle of operation of the device from  [34] (a) before actuation, (b) drug release after the application of the magnetic field. ........................................... 8 Figure  1.3 (a) Schematic design and the principle of operation of the device (b) device implantation into the prostate through a needle. ........................................................................... 11 Figure  2.1 Optical images of the devices showing (a) latitudinal cross-section of the PDMS device without the housing (b) size comparison between the 3D printed device (without housing) and a Canadian 25 cent coin and (c) size comparison between the 3D printed device (without housing) and the tip of a pen. ........................................................................................................ 17 Figure  2.2 Fabrication process of (a) PDMS device and (b) 3D printed device. .......................... 25 Figure  2.3 SEM images of the PDMS device showing (a) longitudinal cross-section of the device without the housing and (b) closer view of the membrane, aperture and the magnetic block and (c) the laser-drilled aperture on the membrane. ............................................................................ 26 Figure  2.4 Magnetic flux density and magnetic field strength of the permanent magnet as a function of distance from the surface of the magnet..................................................................... 28 Figure  2.5 Magnetization curve of the magnetic film versus the applied magnetic field. ............ 31 Figure  2.6 Deflection of the membrane; (a) simulated deflection in a 92.9 mT field, (b) actual deflection of the membrane in 3 different magnetic fields. .......................................................... 34 Figure  2.7 Comparison between deflections from simulation and experimental measurements. The error bars represent one standard deviation from the measured values. ................................ 35 ix  Figure  2.8 Displaced volume under actuation and pumped-in solution after removing the magnetic field................................................................................................................................ 36 Figure  2.9 Maximum travelled distance of the discharged solution in (a) 135.7 mT, (b) 65.8 mT and (c) 32.9 mT magnetic fields. .................................................................................................. 37 Figure  3.1 (a) Cumulative MB release profile over 3 days in a 32.9 mT magnetic field and (b) average MB release rate per actuation interval in a 32.9 mT magnetic field. .............................. 42 Figure  3.2 Cumulative DTX release profile in a 135.7 mT magnetic field over 11 days............. 44 Figure  3.3 (a) Device implantation in swine bladder tissue with a needle, (b) control sample with no actuation after two hours and (c) released MB after two hours of actuation in a 206.3 mT magnetic field................................................................................................................................ 46 Figure A.1 Actuation setup. .......................................................................................................... 61    x  List of Abbreviations and Symbols  %w/v weight (grams) to volume (mililiter) percentage %w/w weight to weight percentage ° degree °C degree Celsius a membrane radius ABS acrylonitrile butadiene styrene ADT androgen deprivation therapy B magnetic flux density BSA bovine serum albumin c concentration Ci curie cm centimeter D diffusion coefficient DCM dichloromethane DPM disintegration per minute DTX docetaxel E elastic modulus emu electromagnetic unit F force H magnetic field strength xi  ID inside diameter IPA isopropanol IR infrared IRD3 implantable rapid drug delivery device kA kiloampere M magnetization MB methylene blue MEMS microelectromechanical systems mg miligrams ml mililiters mm milimeters MPa megapascal MS saturation magnetization mT militesla NdFeB neodymium iron boron ng nanogram nm nanometer OD outside diameter Oe oersted p uniformly distributed load Pa pascal PAA poly(acrylic acid) xii  PBS phosphate buffered saline PDMS poly(dimethylsiloxane) Pe Péclet number pH measure of acidity or basicity RP radical prostatectomy rpm revolutions per minute RT radiation therapy s second SEM scanning electron microscope SQUID superconducting quantum interface device T tesla t thickness TB trypan blue td release time tm mixing time u flow velocity UV ultraviolet v volume ΔV stroke volume μ0 permeability of free space μl microliter μm micrometer xiii  μN micronewton ν flow velocity π Pi (3.1415926) Φ diameter ω0 membrane center displacement  xiv  Acknowledgements  First and foremost, I would like to thank my supervisor, Dr. Mu Chiao, for his invaluable advice and never-ending support throughout the entire project. I am very thankful for the opportunity to work in his lab and under his wise supervision. I am grateful for being able to work with Mr. John Jackson, an exceptional scientist who not only provided me with priceless guidance on the pharmaceutical aspects of my project, but also was a source of inspiration to me. My sincerest gratitude goes to my friend, colleague and mentor, Nazly Pirmoradi for her time, training and support. I would like to thank Dr. Helen Burt from the Faculty of Pharmaceutical Sciences, Dr. Hongshen Ma and Dr. Karen Cheung from the Department of Electrical Engineering for allowing me to use their lab space and equipment and Dr. Ryozo Nagamune for being a member of the defense committee.  I would like to thank my colleagues David Plackett, Kevin Letchford and Leon Wan at Dr. Burt’s group and my fellow lab mates Farzad Khademolhosseini, Kevin Ou, Aurora Chen, Hongbin Zhang, Colin Chen, Hadi Mansoor, Eric Zhao, Kaiwen Yuan, Ali Shademani, Yibo Zhang and Zhengmu Wang at the UBC MEMS lab for sharing their insights and helpful ideas.  Words cannot describe my utmost gratitude to my parents for all their sacrifices, support and ever-lasting love, and to my lovely sister Mojdeh for her continuous encouragement. This thesis would have not been possible without them.  xv  Dedication         To My Parents  ϡέΩΎϣ ϭ έΪ̡ ϪΑ ϢϳΪϘΗ1  Chapter 1: Introduction  1.1 Prostate Cancer and Current Treatments Prostate cancer is the most common type of cancer in men. The Canadian Cancer Society estimates that in 2014, 24% of all male cancer incidence in Canada will be prostate cancer. It is the second leading cause of cancer related deaths in men, surpassed only by lung cancer  [1]. The method of treating prostate cancer highly depends on the stage of the cancer and the conditions of the patient. For early-stage prostate cancer, radical prostatectomy (RP) and radiation therapy (RT) are the primary choices for treatment  [2]. New treatments for localized cancer such as HIFU (High-Intensity Focused Ultrasound) and Cryosurgery (freezing cancer cells to induce cell death) are also emerging  [3]. When cancer becomes metastasized, meaning that the tumor is no longer confined to the prostate and it has spread to other parts of the body, androgen deprivation therapy (ADT, also known as hormone therapy) or Chemotherapy is used to treat the patient  [3].  Most prostate cancer treatments have side effects; some of which significantly deteriorates the quality of life of the patients. Both radical prostatectomy and radiation therapy can cause incontinence, bowel dysfunction and impotence [4, 5]. There have also been reports on post-operative strictures caused by RP  [4]. New localized cancer treatments (HIFU and Cryosurgery) are still in clinical trial stage in some countries and doubts exist regarding their long-term effectiveness  [6]. Most prostate cancer cells become resistant to hormone therapy after few years (castration-resistant prostate cancer) and continue to grow despite the treatment  [7].  Chemotherapy using docetaxel (Taxotere) is a standard treatment for patients with metastatic cancer but it is also used in adjuvant and neoadjuvant settings in organ-confined cancers  [8]. For localized treatment, the efficacy of chemotherapy depends on the amount of drug available at the 2  tumor site. In order to have therapeutic effects, the amount of drug administered through systemic circulation should be high due to the rapid clearance of these potent taxane-based drugs from the blood  [9], which increases the risk of systemic toxicity. The adverse effects of systemic toxicity become more significant when drug interactions are considered. Pharmacodynamics and pharmacokinetics of drug which is available in circulatory system with other foods and drugs can lead to more toxicity or decreased effectiveness of the drug at the cancer site  [10]. This indicates that an effective solution for localized tumor treatment is a system capable of providing localized and controlled drug delivery, preventing the drug from spreading to other organs and hence, reducing the chances of toxicity and potential side effects due to less systemic exposure.  1.2 Controlled Drug Delivery In order to have therapeutic effects, drug concentration should remain in the therapeutic window as it is shown in Figure 1.1 (figure used with permission). If drug concentration exceeds the upper limit of the window, the toxicity associated with the drug may cause undesired outcomes and if the concentration is below the lower limit of minimum drug efficacy, it may not be therapeutically effective and it may increase the chances of drug resistance  [11]. Conventional routes of drug administration such as oral or systemic administration could result in a high initial concentration of drugs in the blood (above toxic level) followed by a concentration decline below the minimum effective level. For those drugs with a narrow therapeutic window such as docetaxel, the chances of serious toxicity are high due to these fluctuations  [12]. In such cases, a controlled drug delivery implant capable of maintaining the drug concentration in the therapeutic window would be advantageous. 3   Figure ‎1.1 Controlled drug delivery compared to conventional drug delivery. Image created based on the image from ‎[13].  1.2.1 Passive Drug Delivery Implants Localized drug delivery implants can provide drug release directly to the tumor at the disease site. To date, various localized drug delivery devices have been proposed. Passive drug delivery implants were the first devices to emerge. In these devices, material properties determine the release rate of drugs from the device and usually the rate cannot be controlled after implantation  [14]. The simplest form of such devices could control the release rate by a membrane, either porous or non-porous  [15]. Drug diffusion through the membrane depends on the molecular size of the drug and the pore size. Biocompatible polymeric devices that encapsulate drugs and rely on polymer degradation for releasing have also been shown  [16]. 4  Other passive systems rely on the change in their surroundings such as pH or temperature change  [17] or osmotic pressure  [14] to slowly release the drug over time with very limited or no control over the rate and time of release. This method of drug delivery can maintain the drug concentration in the therapeutic window for longer periods, but it does not offer dosing flexibility or timing of drug release. In these devices the release rate is predetermined regardless of patient needs and changing conditions  [18].  1.2.2 Active Drug Delivery Implants The shortcomings of passive drug delivery systems resulted in the advent of active drug delivery devices which offered on-demand drug release and dosing flexibility after implantation. Contrary to passive systems, active drug delivery devices that require some kind of actuation in order to release their drug content provide a temporal control over drug release. One of the advantages of the on-demand drug delivery is that drug release can be switched on and off based on the conditions of the patient or the duration of the treatment. Timing drug release can be of utmost importance when there is an unexpected change in the conditions of the patient which requires a relevant change in the dosing. Many devices that can be triggered with external stimuli such as near-infrared light, visible light or ultrasound have been proposed in the literature  [18]. In such devices, the time of the release can be actively controlled but controlling the rate and achieving reproducible release profiles are still challenging. An ideal drug delivery device should also be able to adjust the drug release rate after implantation. Such devices can increase the efficacy of drug therapy since they can provide a specific release profile tailored to each patient’s unique physiology. MEMS serves as a platform for the development of devices that can precisely control drug release kinetics (i.e. time and 5  amount), improving the performance of conventional drug therapies in terms of pain, convenience, efficacy and safety  [19]. MEMS-based devices can eliminate the need for drug injection with remote and repeatable switching of drug release rate  [11]. These devices can also enable complex dosing schedules and deliver a cocktail of anti-cancer drugs  [19]. In many MEMS-based devices, drug is contained in a sealed reservoir which can stabilize active ingredients of the therapeutic agent and keep the drug intact inside the reservoir for long periods  [19].  Active MEMS-based drug delivery devices can be categorized into two main groups: reservoir-based and micropumps.   1.2.2.1 Reservoir-based Implants 2D fabrication processes used for computer chip fabrication were utilized to fabricate conventional silicon-based MEMS devices. They incorporated reservoirs with channels and valves into a single implantable chip. In some of the designs, drug was released from the reservoir by electronically opening the seals or valves of the reservoir [20-25]. Electrochemical  [26] and electrothermal  [27] opening of thin film reservoir seals were shown by Santini et al. The device could achieve controlled release of drugs in dogs for 6 months  [28]. In another design, the reservoir served as a solid drug-containing matrix created by polymeric thin films  [29]. The film was dissolved with the application of electric current and drug molecules that were selectively embedded inside the film were released. Microresistors inside the reservoir were also used to release the drug from the sealed reservoirs. A very good example of such device is an Implantable Rapid Drug Delivery Device (IRD3)  [30] from Massachusetts Institute of Technology, designed to release drugs in emergency situations. 6  As microresistors heat up with the passing current, bubbles are formed and increase the internal pressure of the device until they rupture the sealing membrane and a burst release of drugs takes place in a short time.  1.2.2.2 Micropumps Micropumps with diaphragms have also been proposed for drug delivery. A manually actuated drug delivery pump for treatment of ocular diseases has been developed by Lo et al.  [31]. The device is placed on the eye and it releases its drug contents by manually pressing on. Manual actuation of such devices reduces their applications to the surface of the body and they usually cannot be implanted inside the body due to inaccessibility. Li et al.  [32] proposed an electrochemically actuated device with a refillable drug reservoir and a one-way Parylene check-valve. The passage of current through the electrodes generates gas which increases the pressure on the membrane and causes the drug to release from the Parylene cannula. This device can achieve variable release rates by adjusting the passing current through the electrodes. Other micropumps with shape memory alloy or piezoelectric actuation have also been developed  [33].  1.2.3 Challenges of the Active Implants  Although most of the aforementioned active drug delivery devices can precisely control the time and the dose, they do not fully satisfy all the requirements for an implantable drug delivery device. All of these devices need a power source in order to operate. Current technology has been unable to make batteries small enough for these microdevices. Therefore, the overall size of 7  these devices is determined by their battery size. Not only these batteries are large compared to the size of the devices, but they also need to be replaced when they run out of power.  One way to overcome this challenge is to eliminate the onboard power source from the device. Many material-based systems in which the release mechanism is triggered by a laser, near IR light, visual light or ultrasound have been proposed  [18]. On DemandTM Therapeutics has developed a laser-activated device for the treatment of ocular diseases  [14]. The device has multiple hermetically sealed reservoirs loaded with drugs and it is implanted inside the eye with an intravitreal injection. When on-demand dosing is required, an opening is created in one of the reservoirs using a laser beam. Additional dosing can take place by puncturing the seals of other reservoirs. This battery-less device can control release kinetics but it should only be implanted in the eye where it can be accessed by a laser beam.  Magnetic actuation of drug delivery devices can be a practical solution for battery-less devices. Magnetic actuation provides a remote and non-invasive access to the implanted battery-less device. Such a device can be designed to precisely control the release rate with the magnetic field strength or number of actuations, eliminating the need for an onboard battery and creating a simple, inexpensive and yet precise drug delivery implant. Use of magnetic field to remotely activate an implant on-demand has been introduced by our group  [34] and the developed ocular implant has been shown to achieve controlled and on-demand release of therapeutics for treatment of diabetic retinopathy  [35].  1.2.3.1 A Magnetically-actuated Battery-less MEMS Drug Delivery Device Pirmoradi et al. [34, 35] have developed a battery-less magnetically actuated MEMS drug delivery device for the treatment of diabetic retinopathy. Unlike other devices, this device is 8  capable of controlling release kinetics without the presence of any onboard batteries. This simple device is remotely controlled by external magnetic fields. A schematic view of the device is shown in Figure 1.2 (figure used with permission).  Figure ‎1.2 Schematic design and principle of operation of the device from ‎[34] (a) before actuation, (b) drug release after the application of the magnetic field. The device consists of a PDMS microreservoir (6mm diameter, 550 μm depth) which is permanently bonded to a thin magnetic PDMS membrane  [36]. A 100×100 μm2 aperture is drilled on the magnetic membrane using laser ablation. Docetaxel (DTX), an anti-proliferative drug widely used in chemotherapy and ophthalmic disease treatments such as diabetic retinopathy, is deposited into the reservoir in solid form. The device is filled with a solution of bovine serum albumin in phosphate buffered saline (BSA in PBS) to enhance hydrophilicity of the membrane and to simulate the interstitial fluids medium inside the body. When this solution fills the reservoir, it mixes with the deposited solid drug and forms a saturated drug solution inside the reservoir. 9  When an external magnetic field is applied, the magnetic PDMS membrane deflects and increases the internal pressure inside the reservoir. Consequently, drug is expelled from the reservoir and through the laser-drilled aperture. The rate of the release can be controlled either by the number of actuations or by the external magnetic field strength. The device is capable of releasing 171 ng of docetaxel per actuation in a 255 mT magnetic field. When magnetic field is removed, the membrane rebounds and fresh BSA/PBS solution refills the reservoir and forms a new saturated drug solution. Interstitial fluids inside the body would play the role of the BSA in PBS solution once the device is implanted inside the body. Since docetaxel has a very low aqueous solubility in water, it only dissolves a very small amount following each actuation and permits thousands of repeat doses before the drug is gone. If enough mixing time is given to the solution, the concentration of DTX in the solution returns to the maximum level (i.e. saturation limit) after each actuation as long as solid drug exists inside the reservoir. Therefore, the concentration of DTX in the released solution does not decrease following consecutive actuations, given that enough mixing time has been considered in between the actuations. The device achieved reproducible release rates in vitro over 13 days.  This battery-less device can control the time and dose of release, it is simple and inexpensive made from a biocompatible polymer (i.e. PDMS). However it has shortcomings that limit its applications: 1. Although no onboard batteries are present, the size of the device is still large for implantation. This comes from the fact that the diameter of the membrane should be large enough to provide ample displacements with a relatively small magnetic force. 2. The device requires invasive surgery for implantation. 10  3. The magnetic field gradient should be relatively large in order to actuate the device. Strong magnetic field gradients exist at a close proximity to the permanent magnet and the gradient decreases at further distances. Consequently, the external magnet should be placed very close to the device in order to actuate it. In other words, the device is only suitable for implantation near the surface of the body with a close proximity to an external magnet in order to operate. Therefore, although this device could be a good candidate for the treatment of diabetic retinopathy, in which the device is implanted at the back of the eyeball and close to the surface of the body, it might not be a very good candidate for prostate cancer treatment where the gland is located in the middle of the body. In order to be a suitable treatment option for localized prostate cancer, the implantable device should overcome the aforementioned hurdles.  1.2.3.2 Minimally Invasive MEMS Drug Delivery Implant for the Treatment of Prostate Cancer The objective of this research is to design, fabricate and test a battery-less and magnetically controlled drug delivery device for the treatment of prostate cancer. The device should be implanted through a needle with minimally invasive procedures. The geometry and the actuation mechanism of the device should be redesigned, to make it suitable for deeper implantation. This method of implantation is similar to a widely-used technique for localized prostate cancer treatment known as brachytherapy. Brachytherapy is the implantation of radioactive seeds inside the prostate. The seeds with the size of a grain of rice are guided through a long needle and placed inside the prostate. An ultrasound device maps the prostate for accurate seed positioning. Brachytherapy is an effective 11  and convenient treatment for low-risk tumors since there is no need for invasive surgery and patients who receive this treatment are usually able to get back to their daily activities after a few days, but there are complications due to the continuous radiation inside the prostate.  Docetaxel is a potent drug with substantial adverse effects when administered intravenously in chemotherapy for the treatment of localized cancer. It is also used as a secondary treatment when the major therapy is either radiation or radical prostatectomy, to lower the chances of cancer recurring. Efforts have been geared towards the development of less toxic solutions for delivery of taxanes  [9], However, delivering precise doses of DTX directly to the cancer site   Figure ‎1.3 (a) Schematic design and the principle of operation of the device (b) device implantation into the prostate through a needle. 12  can add to its clinical value and opens up new applications by minimizing its adverse effects.  The size and the shape of the new device allow for implantation through a needle with minimally invasive procedures allowing for reduced side effects associated with full prostate removal. This device provides localized drug delivery directly to the prostate, minimizing drug interactions with other tissues and potentially fewer side effects.  Moreover, DTX has unpredictable interpatient variability in efficacy and toxicity and the same recommended doses could lead to overexposure in some patients while providing a suboptimal and ineffective treatment for the others  [37]. An active control over drug release could be used to manipulate drug release rate after implantation based on the patients’ physiological response to the drug and optimize their treatment  [19], a capability lacked by passive drug delivery systems. The drug dosing flexibility provided by this device could facilitate achieving accurate and complex dosing schedules required for cancer treatments (i.e. timing of release)  [19], controlling the duration of the treatment  [18] while enabling precise adjustments of the release rate within DTX’s therapeutic range with respect to patient needs and their changing physiological state. This device can be added to the active surveillance strategy in prostate cancer management, to prevent or slow down cancer progression in patients who have low-risk localized prostate cancer. Figure 1.3 shows the schematic design of the device and its implantation through a needle.  1.3 Thesis Overview This thesis is prepared by an introductory chapter and followed by two main chapters and a concluding chapter that summarizes the work and gives future directions. Chapter 1 briefly reviews prostate cancer and its current treatments. The shortcomings of the current approaches are highlighted and targeted drug delivery using MEMS platform is proposed 13  as an alternative approach with potentially fewer side-effects and increased efficacy. MEMS drug delivery devices are grouped into active and passive systems and active systems are briefly reviewed due to their superior control over drug release kinetics. These active systems are further categorized in two main groups, reservoir-based and micropumps. The current challenges of the existing devices are mentioned and magnetic actuation is proposed to trigger the release mechanism. Chapter 2 provides an overview into the design, fabrication and characterization of the two types of devices that were presented in this work. Key design parameters including the enhanced actuation mechanism and needle implantation technique are discussed and the step-by-step fabrication process for each type of the devices is given. This is followed by describing the actuation setup used for three different set of experiments: Methylene Blue controlled release, Docetaxel controlled release and tissue implantation experiments. The device is magnetically characterized and the theoretical amount of magnetic force is calculated. Then membrane deflection is simulated using COMSOL Multiphysics Software. The data is used for a set of calculations such as drug release time. Two time constants are calculated for continuous actuation of the device to assure constant concentration of drug in the discharged solution. Chapter 3 includes the data and discussion from all the experiments. Cumulative release profile and the average release rates of a model drug and DTX are presented, followed by a qualitative study of device operation implanted in a fresh swine bladder tissue through a needle. Chapter 4 summarizes the work presented in this thesis and proposes future directions for further development of the device.  14  Chapter 2: Design, Fabrication and Characterization  2.1 Design The device consists of a microreservoir, an elastic PDMS membrane and a magnetic block bonded to the membrane. An aperture (100×100 μm2) is drilled on the PDMS membrane to provide an opening for the reservoir contents to be released upon actuation. Additional housing is attached to the reservoir and the membrane to prevent the tissues from coming into contact with the thin and sensitive membrane. When magnetic field is applied, it exerts force on the magnetic block and attracts it. This causes the magnetic block to move towards the positive field gradient and consequently, the membrane that is permanently bonded to the magnetic block deflects. The deflection of the membrane increases the pressure inside the reservoir and pumps out the drug through the aperture. The drug enters the housing which serves as a medium in which the drug can diffuse to the adjacent tissues over time.  Similar to the diabetic retinopathy device  [35], docetaxel is deposited in the device as a solid drug prior to filling the reservoir with BSA in PBS solution. After filling, a saturated drug solution is formed in the reservoir and the concentration will remain almost constant after consecutive releases due to the reasons mentioned in section 1.2.3.1. When magnetic field is removed, the membrane rebounds and the pressure inside the reservoir drops. As a result, the fluids surrounding the device refill the reservoir and form a new saturated solution ready for the next release. This device was originally designed to treat localized prostate cancer or prevent cancer progression into higher stages. However, it can be used in other applications (e.g. breast cancer) 15  and will be equally as feasible. Other potent drugs with low solubility in water can also be delivered using this device (e.g. paclitaxel). Two devices are designed and fabricated. One is completely made from PDMS (reservoir, membrane and housing) while the other one has a PDMS membrane but the reservoir and the housing are made from a UV curable photopolymer known as PlasClear used in 3D printing machines (The Freeform Pico, Asiga, CA, USA). PDMS is biocompatible, flexible and transparent which makes it a very good candidate for the flexible membrane fabrication. But due to some limitations for the fabrication of smaller devices with PDMS and also due to the presence of unknown external forces inside the body which could deform the reservoir and result in unwanted drug release, the second smaller and structurally rigid device was designed and fabricated. In both designs, the shape and the actuation mechanism are exactly the same; however, their size, material and some fabrication steps are different. The pioneering device that was designed by our group before  [34] used a flexible and magnetic PDMS membrane by uniformly dispersing coated iron-oxide nanoparticles in a PDMS matrix. In that design, the magnetic membrane was responsible for producing magnetic force and ample deflection simultaneously. However, this approach has a trade-off that ultimately limits the use of the device. The more the amount of particles in the PDMS membrane, the stronger the magnetic force exerted onto the membrane in an external magnetic field. But increasing the weight concentration of particles in the PDMS matrix could lead to particle agglomeration and a rough membrane surface incapable of forming a leakage-free seal with the reservoir. Increasing the thickness of the membrane could accommodate more magnetic particles in the PDMS matrix, but from a mechanical point of view, a thicker membrane results in an increased bending modulus, which requires even more force for deflection. A maximum weight concentration of 16  32% magnetic content in the PDMS membrane (40% weight concentration of particles in PDMS) was achieved for a uniform magnetic membrane in  [34]; only enough to deflect the membrane in strong magnetic field gradients that that exist very close to the external magnet’s surface. In order to increase the actuating distance, the actuation mechanism had to be changed. In the new design, the magnetic force generating component is separated from the deflection generating component, so that each part can be improved separately. Therefore, instead of dispersing the particles inside the PDMS membrane, a rectangular magnetic block was attached to the center of a pure PDMS membrane. All of the magnetic particles are collected in the magnetic block, taking up to 25% of the total area of the membrane. The rest of the membrane is smooth and can be permanently bonded to the reservoir to make a seal. As a result of separating the force generating component from the deflection generating component, the followings can be achieved:  The amount of magnetic particles in the magnetic block can be increased as particle agglomeration will not be problematic.  Thinner membranes can be made which require less force for deflection The magnetic block is made from iron-oxide nanoparticles (Nanostructured and Amorphous materials, Los Alamos, NM, USA) sandwiched between PDMS layers and it is permanently bonded to the center of the membrane. Therefore, when external magnetic field is applied, the translational motion of the block causes the membrane to deflect. The thickness of the magnetic block is independent from the thickness of the PDMS membrane and it can be much thicker than the membrane to contain more magnetic particles. 17  The PDMS device is 12 mm long with 1 mm thick walls. The outside diameter of the reservoir is 5mm and the inner diameter is 3mm. The PDMS membrane is ~55 μm thick and the aperture size is 100×100 μm2. The magnetic block is 1.5×5 mm2 with ~218 μm thickness. Since both the magnetic block and the aperture are located in the center of the membrane, a small 1 mm hole is punched in the middle of the magnetic block to allow the fluids to pass through the aperture when the device is actuated. The housing has the same dimensions as the reservoir, but with a 10 mm long opening to allow the released drug to diffuse to the adjacent tissue. The main role of the housing is to prevent the tissues from touching the sensitive membrane and clogging the aperture. Fabrication of smaller devices made from PDMS is very challenging, but high precision 3D-Figure ‎2.1 Optical images of the devices showing (a) latitudinal cross-section of the PDMS device without the housing (b) size comparison between the 3D printed device (without housing) and a Canadian 25 cent coin and (c) size comparison between the 3D printed device (without housing) and the tip of a pen. 18  printing enables the fabrication of much smaller devices. The 3D printed device is almost 6 times smaller than the PDMS device. This device is 12 mm long but with an outside diameter of 2 mm and 0.3 mm thick walls. The magnetic block is 5×1 mm2 with a small hole punched in the middle. All other dimensions are identical to the PDMS device. The optical images of the devices are illustrated in Figure  2.1.  2.2  Fabrication Traditional 2D fabrication techniques used for computer microchip fabrication such as lithography may not be suitable for the fabrication of cylindrical devices. Therefore, we had to use a combination of molding and 3D printing techniques in order to make our devices. Here, the fabrication steps for both devices are explained in detail.  2.2.1 PDMS Device As shown in Figure  2.2, the fabrication of the PDMS device can be summarized in the following steps:  Step 1: Two 3D printed molds were made from Acrylonitrile Butadiene Styrene (ABS); a common thermoplastic material. One of the molds has a 12 mm long and 5 mm diameter half-cylindrical groove and the other one has a 10 mm long and 3 mm diameter half-cylinder bump. The molds have features that align both half-cylinders when placed together. PDMS (Sylgard 184 Silicone Elastomer, Dow corning Corporation) was prepared with a mixing ration of 10:1 pre-polymer to cross-linker. It was then casted in the mold with a groove and degassed for 30 minutes. The mold with the bump was placed on top of the PDMS filled mold and pressed on it. 19  The molds were placed in a convection oven at 70º C for 4 hours. Then the molds were detached from each other and the PDMS reservoir was peeled off from the molds. At this point, the reservoir has a rough surface due to the surface roughness of the molds and was not suitable for bonding to a smooth membrane.  Step 2: Poly(acrylic acid) (PAA) was used as a sacrificial layer. PAA powder (Mw=1800, Sigma Aldrich) was mixed with distilled water with a 20% w/v concentration. It was mixed and sonicated for 30 minutes and then filtered (4.5 μm pore size, Millipore Corporation, Ma, USA). Three pre-cleaned glass slides were washed with Isopropyl alcohol (IPA) and air dried. The glass slides were treated with air plasma for 75 seconds to enhance their hydrophilicity. Then, PAA was spin-coated on the glass slides in two steps: 500 rpm for 10s and 1000 rpm for 30s. The glass slides were placed on a hot plate at 150º C for 5 minutes to let the water content evaporate and form the sacrificial layer. This process was done on three glass slides and each one was used in a different step during the fabrication process.   Step 3: The prepared PDMS in the first step was spin-coated (500 rpm for 10s and 1500 rpm for 40s) on the glass slides with the sacrificial PAA layer. Before PDMS was cured, the reservoir was placed on the glass and cured at 70ºC for 4 hours. Then the glass slide was immersed in water until the sacrificial layer was dissolved and the reservoir was released from the glass slide with a smooth surface suitable for bonding to the PDMS membrane. The reservoir was then loaded with drug. In this study, we used Methylene Blue (MB) as a model drug due to its high sensitivity of detection using UV-Vis absorbance spectrometry  [34]. Although it was possible to deposit MB as a powder, this method of deposition increased the chances of contamination of the 20  regions that were supposed to be bonded to the membrane afterwards. A better technique was to dissolve MB in water, deposit the solution in the reservoir with a pipette and let the water evaporate. Multiple depositions and evaporations can take place to increase the amount of solid MB in the reservoir.  Step 4: The magnetic block was made by sandwiching 2 layers of magnetic particles between PDMS layers. Similar to step 3, PDMS was spin-coated on the second glass slide with the sacrificial PAA layer. Before PDMS was cured, iron-oxide particles were sprinkled on the PDMS layer until everywhere on the glass slide was covered and no more particles could stick to it. When the first layer of PDMS was cured, another layer of PDMS was spin-coated on top of this layer and a second layer of magnetic particles was added to increase the weight concentration of magnetic particles in the magnetic film. This method can be repeated several times if a very large magnetic force is required. Each step adds about 100 μm thickness to the magnetic film if the same PDMS pre-polymer to cross-linker ratio (i.e. 10:1) and spinning parameters (i.e. 500 rpm for 10s and 1500 rpm for 40s) are used. The final layer of the magnetic film is PDMS to prevent particle leaching from the film in strong magnetic fields. The first layer is also a very smooth PDMS layer, allowing the block to bond to the PDMS membrane using air plasma. After all layers were formed, the glass slide was immersed in water and the magnetic film was released from the glass slide. A 1.5×5 mm2 block was cut from the film and a 1 mm diameter hole was punched in the center of the block. In this work, the terms magnetic block and magnetic film are the same magnetic composite and might be used interchangeably.  21  Step 5: PDMS was spin-coated on the 3rd glass slide with the sacrificial PAA layer (500 rpm for 10s and 1500 rpm for 40s) to make the membrane with ~55 μm thickness. After curing in a convection oven at 70° C for 4 hours, the PDMS-coated glass slides, drug-loaded reservoir and magnetic blocks were all treated with air plasma for 75 seconds. The air pressure inside the chamber was maintained at ~700 mTorr during this period. After treatment, magnetic blocks were bonded to the membrane and then, the membrane was bonded to the drug-loaded reservoir in a way that magnetic block was positioned in the center of the device. The glass slide was immersed in water and the sealed reservoir was released with a flat membrane holding the magnetic block at its center.  It should be noted that no plasma induced changes (i.e. change in the UV-Vis absorbance spectrum) were observed in MB following air plasma exposure  [35]. However, in the case of radioactive DTX deposition, DTX solution can be loaded inside the reservoir after plasma treatment in order to avoid plasma exposure and also to avoid chamber contamination with radioactive particles. When the solvents evaporate, the reservoir and the membrane can be bonded together. Instead of water, DTX was dissolved in a 50/50 solution of ethanol and dichloromethane prior to deposition inside the reservoir.  Step 6: The aperture was created by laser ablation using a Nd:YAG laser (Quicklaze, New Wave Research, Sunnyvale, CA). Green laser (532 nm wavelength) with the properties 0.6 mJ (100% high), laser pulses at 35 Hz and a scanning speed of 10 μm/s was used.  After laser ablation, the bottom of the reservoir was inspected and no change was observed, so it's expected that the laser beam does not reach the bottom of the reservoir and decompose the deposited drug content. However, even if the small focused laser beam could reach the bottom of 22  the reservoir, it would still interact with a tiny fraction of the deposited drug and it would not alter the drug composition noticeably.  Step 7: The housing was created with the same procedure as the reservoir (steps 1 and 2). A 4 mm diameter hole was punched on the housing to allow the released drug to diffuse to the surrounding environment. The reservoir and the housing were treated with air plasma for 75 seconds at 700 mTorr air pressure and then bonded together. The hydrophobic nature of PDMS prevents the initial filling of the device with fluids. In order to fill the reservoirs with fluids, the surface of the reservoir and the membrane had to be treated to enhance their hydrophilicity. The devices were put into a 4% solution of BSA in PBS and then inside a vacuum chamber. The air was pumped out of the vacuum chamber. As a result of this pressure drop, the air that was trapped inside the reservoir expanded and most of it left the reservoir through the aperture. The pressure inside the vacuum chamber was then returned to the ambient pressure. The suction caused by the low air pressure inside the reservoir drove the fluids inside the reservoir and partially filled it. Devices were then placed in a 37° C oven over night to incubate. The BSA/PBS solution and the incubation process help the remaining parts of the reservoir to get filled. No bubbles were observed inside the reservoir after this step.  2.2.2 3D Printed Device As mentioned before, due to some constraints with the size and flexibility of the PDMS device, a second smaller and more rigid device was fabricated using high precision 3D printing technology. Because PDMS device is bigger, it is challenging to glue a flat membrane to the reservoir and hence, PDMS device requires more fabrication steps (i.e. surface smoothing, 23  membrane bonding, etc.). But the 3D printed device is smaller and a flat PDMS membrane can be directly glued to the reservoir with fewer fabrication steps. The step by step fabrication process of the 3D printed device is presented in Figure  2.2.  Step 1: Drug reservoir was made with a high precision Freeform Pico 3D printer (Asiga, CA, USA) using PlasClear, a UV curable polymer (Asiga, CA, USA). During printing, the reservoir was partially cured with UV. After printing, the reservoir was thoroughly washed with IPA and put in room temperate to dry. Then, it was placed in a UV chamber (Asiga, CA, USA) for 20 minutes for additional curing which made it more rigid. This was an important step in the fabrication of such a small device. The reservoir volume should be as big as possible to contain more drugs and provide a larger deflection area for the membrane. This means that the walls should get thinner if the outer diameter of the device is constant. The more rigid the thin walls, the easier it is to handle the device and assemble other parts, and the less the chances of device deformation under physical loads.  Step 2: Similar to the 3rd step in the fabrication of the PDMS device, MB was dissolved in water and deposited into the reservoir with a syringe and water was allowed to evaporate. Additional MB can be deposited by repeating this step to increase the amount of solid MB in the reservoir.  Step 3: Similar to the 2nd step in the fabrication of the PDMS device, a glass slide was coated by PAA followed by spin-coating of PDMS with the same recipe. Before PDMS was cured, magnetic block was placed on the glass slide and the reservoir was placed afterwards in a way that the magnetic block was located in the center of the reservoir. The PDMS layer was cured at 24  70° C for 4 hours. The glass slide was then immersed in water and the reservoir was released with a flat membrane glued to the walls by the adhesive nature of PDMS and a magnetic block in the center of the membrane.  Step 4: An aperture was created in the center of the membrane similar to the 6th step in the fabrication of the PDMS device. All laser properties and aperture dimensions are the same as the PDMS device.  Step 5: The housing was made by 3D printing, washed with IPA and cured with UV similar to the first step. In this design, the housing has a 10 mm long and 1 mm wide opening. The housing is placed on the drug loaded reservoir using a universal flip chip die-bonder (JFP Michrothechnic, Marcoussis, France). Using this device, the reservoir and the housing were aligned and their vertical positions were controlled. When the housing touched the membrane, it was held in place until the two ends of the device were glued with two small droplets of PlasClear. The polymer droplets were partially cured with a UV spot cure system (Electro-Lite Corporation, Bethel, CT, USA) for 1 minute and then placed in the UV chamber for 20 minutes to fully cure and glue the ends of the device. 25   Figure ‎2.2 Fabrication process of (a) PDMS device and (b) 3D printed device.  26   Figure ‎2.3 SEM images of the PDMS device showing (a) longitudinal cross-section of the device without the housing and (b) closer view of the membrane, aperture and the magnetic block and (c) the laser-drilled aperture on the membrane.  2.3 Actuation Setup Controlled release studies were conducted in two different setups: PDMS devices were used for in-vitro release experiments due to their bigger sizes and convenience in their handling. 3D-27  printed devices were used for the ex-vivo experiments that qualitatively show the feasibility of implantation of the device into animal tissue through a needle, and also device operation inside the tissue. The 3D printed devices are the ones with the future prospect of implantation in the prostate. Although the amount of drug release in PDMS devices and 3D printed devices could vary, but the release data obtained from the PDMS devices serve as proof-of-concept release studies and demonstrate the consistency of release rates of the device.  2.3.1 In-vitro Testing Setup In this work, PDMS devices were used for controlled in-vitro release studies using the model drug MB and also radioactive DTX. The magnetic field of the permanent magnet was measured as a function of distance from the surface of the magnet with a Bell Gaussmeter (Sypris Test & Measurement, FL, USA) and it is depicted in Figure  2.4. The permanent magnet is a cylindrical NdFeB magnet (𝐷𝑖𝑎𝑚𝑒𝑡𝑒𝑟 =12", 𝐿𝑒𝑛𝑔𝑡ℎ =34")  positioned on a motorized stage (an Atmega328 microcontroller on the Arduino Duemilanove board, by Arduino©, Italy) and controlled by a computer. The device was placed inside a petri dish and then on a height-adjustable stage next to the magnet. The motorized magnet could move back and forth underneath the device (Appendix A). All distances were measured from the surface of the magnet to the magnetic block on the membrane. The device was glued (LePage, Henkel Canada Co, Mississauga, Ontario, Canada) to the bottom of a petri dish so that it would not float in the BSA/PBS solution and that the magnetic field would always remain perpendicular to the membrane. Once glued, the device was first rinsed with 1 mL of BSA/PBS solution to get rid of any dust particles. Then, the petri dish was filled with BSA/PBS solution until the entire device was submerged and the volume was recorded for 28  calculations. The petri dish was located on two microscopic slides that were attached to the stage. The height of the stage was adjusted so that the magnetic field at the location of the magnetic block was in the desired range.  Figure ‎2.4 Magnetic flux density and magnetic field strength of the permanent magnet as a function of distance from the surface of the magnet.  2.3.2 Ex-vivo Testing Setup 3D printed devices were implanted inside swine bladder tissue and activated for two hours. MB was used as a model drug in this study but the amount of the released MB was not measured. The 29  aim of this study was to show drug release from the smaller 3D printed devices qualitatively and to demonstrate the feasibility of implanting the devices through needles in a fresh animal tissue.  Two inch long reusable blunt-tip dispensing needles (McMaster-Carr, Aurora, OH, USA) were purchased and used to implant the devices into the tissue. The needles used were gauge 12 (ID =  2.16 mm, OD = 2.77 mm) and gauge 9 (ID = 3 mm , OD = 3.76 mm). The tissue was first pierced with a sharp-tip hypodermic needle followed by the insertion of blunt-tip needles for device implantation. The bladder tissue was cut into small pieces with a thickness of roughly 3 mm each. The tissues were fresh and kept inside Tyrode’s solution (pH = 7.4) and refrigerated to stay fresh. After implantation, the tissues were washed with the buffer solution and placed in a petri dish. The tissue did not float in water and therefore did not require gluing to the bottom of the petri dish. The petri dish was filled with Tyrode’s solution until the tissue was submerged, then placed on the glass slides attached to the height-adjustable stage. The device was actuated 360 times in two hours. Each actuation cycle consisted of applying magnetic field for 10s (i.e. discharging MB) followed by a 10s relaxation time (i.e. refilling reservoirs). After this time, the tissue was taken out, and cut open. The devices were removed and the released MB was compared to the control tissue with a device implanted, but with no magnetic actuations.  2.4 Characterization Iron-oxide nanoparticles Fe3O4 (Nanostructured and Amorphous materials Inc., Los Alamos, NM, USA) were purchased as a powder and used to make the magnetic block. Although other magnetic particles with higher magnetization (e.g. Iron nanoparticles) could be used to provide a larger force, Iron-oxide was used to enable comparison of the current design with the previously 30  fabricated device  [34] that used a magnetic membrane made from the same particles. Moreover, Iron-oxide particles are widely used in biomedical and drug delivery applications. The powder is 98% pure with an average particle size of 20-30 nm. In order to maximize the magnetic force, the amount of particles should be increased. This is in contrast with maximizing deflection, as a thin membrane is required to achieve large deflections. Therefore, by separating the magnetic part from the membrane, each component can be improved individually. The thickness of the magnetic block was increased to accommodate more magnetic particles without interfering with the membrane deflection.  In this study, we sandwiched two layers of iron-oxide nanoparticles between 3 layers of PDMS. The average thickness of the fabricated magnetic film was measured from the SEM images of the sample and it is 218±20 μm. The density of the fabricated magnetic film was measured by dividing the measured weight of three different samples and dividing them by their volume obtained from image processing. The average density of the magnetic film is 1.960±0.066 g/cm3.  We used a superconductive quantum interface device (SQUID) to measure the magnetization curve of the magnetic film versus the applied field. Assuming that the magnetic film was in the x-y plane, the magnetic field was applied in z direction (perpendicular to the magnetic film) and the magnetism was also measured in that direction (i.e. Mz). Magnetization of the sample was measured at two different temperatures; room temperature at 25° C and body temperature 37° C, to account for the differences between the laboratory conditions and when the device is implanted in the prostate. The measured values of Mz at these two temperatures were very close and therefore, we used the magnetization values at 37° C for all of the calculations. The data is shown in Figure  2.5. 31   Figure ‎2.5 Magnetization curve of the magnetic film versus the applied magnetic field.  2.4.1 Magnetic Force When a magnetic dipole is placed inside a magnetic field, there are two types of loads that are exerted on the dipole. One is the magnetic torque ?⃗? = 𝜇0𝑣?⃗⃗? × ?⃗?  that rotates the magnetization axis ?⃗⃗?  towards the direction of the applied magnetic field. μ0 is the permeability of free space which is equal to 4𝜋 × 10−7  N/A2 and 𝑣 is the volume of the dipole. This torque depends on the magnetic field strength ?⃗? . Another load is the magnetic translational force 𝐹 = 𝜇0𝑣(?⃗⃗? . ∇)?⃗?  which pulls the dipole towards the positive field gradient and depends on the magnetic field gradient. In our case, magnetic translational force is responsible for the movement of the magnetic block and the deflection of the membrane. This indicates that a larger field gradient at 32  the location of the magnetic block, results in a larger magnetic force and hence, larger deflections. Equation (2.1) indicates that the magnetic field strength H (A/m) can also be expressed by the magnetic flux density B (T) as follows:  𝐵 = 𝜇0𝐻 (2.1) Since there is no electric current flowing in the dipole, Maxwell’s equations provide the constraint ∇ × 𝐻 = 0. Therefore, magnetic force can be expressed as follows  [38]:  𝐹 = 𝜇0𝑣[      𝜕𝜕𝑥𝐻𝑇𝜕𝜕𝑦𝐻𝑇𝜕𝜕𝑧𝐻𝑇]      𝑀 (2.2) We assume that the magnetic field variation in the x-y plane is negligible. Therefore, the translational force in the z direction is simplified as:  𝐹𝑧 = 𝜇0𝑣𝛥𝐻𝑧𝑡𝑀𝑧 (2.3) where 𝑡 is the thickness of the magnetic block and 𝛥𝐻𝑧 is the difference in the magnetic field strength at the top and bottom surfaces of the magnetic block. Therefore, 𝛥𝐻𝑧𝑡 corresponds to the magnetic field gradient at the location of the magnetic block and it is calculated at different distances between the magnet and the magnetic block based on the data shown in Figure  2.4. The value of 𝑀𝑧 is obtained from the data shown in Figure  2.5 after conversion into SI units by using the density of the magnetic film.  33  2.4.2 Membrane Deflection In order to estimate the amount of deflection in our device, COMSOL multiphysics software was used. These simulations were verified by experimental deflection measurements. To further verify the results of the COMSOL model with analytical solutions, the deflection of a circular membrane with a uniform load was simulated and compared with the analytical solution (Appendix B). The results of the circular membrane deflection were in excellent agreement with the analytical solution which indicates that this method could be used to estimate the deflection of a more complex rectangular membrane with a uniform load applied to a portion of the total area of the membrane (i.e. magnetic block area). The analytical magnetic force values obtained from the previous section were used as uniform loads exerted on the magnetic block. The elastic modulus for Sylgard 184 PDMS with a mixing ratio of 10:1 pre-polymer to cross-linker was reported at 1.8 MPa in the literature [39, 40]. PDMS Poisson’s ratio was reported at 0.5 in the literature [41, 42]. For our simulations we used the value 0.45 for Poisson’s ratio which is acceptable in literature  [43] since higher values led to solution divergence. The typical thickness of the membrane is ~55 μm. The dimensions of the magnetic block were accurately measured by image processing techniques and were equal to 1.8 × 5.4 × 0.218 mm3. The membrane deflection was measured under different magnetic fields in order to verify the results obtained from the COMSOL model. The PDMS device with the aforementioned dimensions was placed on a glass slide and taped to a height-adjustable stage. The permanent magnet was placed on a magnet holder and screwed to a vertical microstage (Melles Griot, NY, USA). A stereo microscope (Olympus, MA, USA) was aligned perpendicular to the axis of the magnet and parallel to the membrane. The distance between the magnetic block and the magnet and the corresponding deflection was measured by image processing using ImageJ software. 34  Figure  2.6 shows the simulated membrane deflection and the measured deflection of the membrane and the results are compared in Figure  2.7. The results are summarized in Table  2.1.  Figure ‎2.6 Deflection of the membrane; (a) simulated deflection in a 92.9 mT field, (b) actual deflection of the membrane in 3 different magnetic fields.  Distance from Magnet (mm) Magnetic Field (mT) Magnetic Force (mN) Simulated Deflection (μm) Measured Deflection (μm) Displaced Volume (μl) % of Reservoir Volume 1 424 12 244 244±44 3.3 9.4 3 259 6.3 181 203±12 2.5 7 5 166.4 3.4 129 148±14 1.8 5.1 7 111.8 1.8 85 101±19 1.2 3.4 9 77.9 1 52 58±29 0.7 2.1 Table ‎2.1 Theoretical forces on the magnetic block and the corresponding deflections of the membrane. 35   Figure ‎2.7 Comparison between deflections from simulation and experimental measurements. The error bars represent one standard deviation from the measured values.  2.4.3 Release Time The device should be actuated several times so that the released drug concentration reaches the therapeutic range. In order to release repeated doses of drug in consecutive actuation cycles, the membrane should be given time to achieve maximum deflection. This way, we ensure that the volume of the released drug solution in each actuation cycle is constant and equal to the maximum displaced volume of the device. If the release time is too short, then the amount of released drug would be less than the maximum displaced volume. This could increase the required number of actuations in order to get to the therapeutic concentrations and increases the overall time of therapy. Therefore, drug release time is the minimum time required for the release of the maximum achievable volume of the discharged solution (equal to the maximum 36  displaced volume, see Figure  2.8) under a specific magnetic field in a single actuation. In order to estimate drug release time, the velocity of the discharging solution from the aperture should be measured. An MB-loaded PDMS device was placed in a petri dish. The petri dish was then filled with 1% BSA/PBS solution and placed on a stage. A stereo microscope (Olympus, MA, USA) was aligned horizontally and focused on the membrane. Drug release was recorded at 20 frames per second with a camera attached to the end of the microscope (Basler vision technologies, Ahrensburg, Germany). Therefore each frame is 50 ms. Figure  2.9 shows the maximum travelled distance of the discharged MB in three different magnetic fields. As the magnetic field gradient becomes stronger, it exerts more force on the magnetic block, leading to an increase in the acceleration of the block. This causes the membrane to deflect faster and discharge the MB solution with a higher velocity.   Figure ‎2.8 Displaced volume under actuation and pumped-in solution after removing the magnetic field. 37   Figure ‎2.9 Maximum travelled distance of the discharged solution in (a) 135.7 mT, (b) 65.8 mT and (c) 32.9 mT magnetic fields. It is reported that the viscosity of the BSA solution does not change by the addition of Trypan Blue (TB) at a concentration of 1 mg/mL and that it is expected that no change in viscosity of the BSA solution would occur if nanomolar concentrations of DTX are used  [35]. MB has very similar physical characteristics to TB and it is also expected that changes in the viscosity of the BSA solution by addition of MB in low concentrations are negligible.  Therefore, the measured velocity of the released MB is assumed to be equal with the velocity of the discharged DTX solution. It is assumed that the velocity of the discharged solution is constant during the release time and equal to the initial velocity of the discharging solution. However, this is a simplifying assumption as the velocity of the released solution would probably decrease over time. Still, this simplified discharge velocity could be used for estimation purposes. Therefore, drug release time can be calculated from the following equation:  𝑡𝑑  =𝛥𝑉𝐴. 𝜈 (2.4) In which td is the time of release in seconds, ΔV is the stroke volume calculated in the previous section, A is the aperture area and ν is the measured velocity of the released solution obtained 38  from the first four frames of the recorded MB release videos. Drug release times in three different magnetic fields are demonstrated in Table  2.2. Distance from Magnet (mm) Magnetic Field (mT) Maximum Travelled Distance (μm) Initial Velocity (mm/s) Displaced Volume (μl) Release Time (s) 6 135.7 1269±74 4.4±0.5 1.5 34.1 10 65.8 644±25 1.4±0.2 0.6 42.9 15 32.9 440±28 0.6±0.1 0.2 33.3 Table ‎2.2 Estimated release time in different magnetic fields.  2.4.4 Mixing Time Another important time constant required for continuous actuation of the device is the mixing time. Mixing time is the minimum time required for the concentration of the pumped-in solution to return to its saturation limit after the magnetic field is removed (see Figure  2.8). To obtain this time constant, Pirmoradi et al. solved the mass balance equation for transport using COMSOL Multiphysics software  [35]. This equation is defined as follows:  𝜕𝑐𝜕𝑡+ 𝑢. 𝛻𝑐 = 𝛻. (𝐷𝛻𝑐) (2.5) Where c is the concentration, D is the diffusion coefficient and u is the velocity vector. The term 𝑢. 𝛻𝑐 on the left is the convective transport and the term on the right-hand side of the equation is the diffusive transport. In  [35], the velocity field inside the reservoir was neglected since the mixing time required by the diffusion was longer than that of convection because of the low Péclet number (1<Pe<1000). Moreover, it was stated that the velocity field only existed for 0.1% of the total mixing time inside the reservoir which facilitates the mixing. So by ignoring the velocity field, an over-estimated value for the required mixing time inside the reservoir was 39  obtained. Equation 2.5 was reduced to Fick’s law for diffusive transport and solved in a 2D space with the assumption of isotropic diffusion. The mixing time for DTX with the diffusion coefficient of 9 × 10−10  m2/s was calculated at tm = 200 seconds. This is the time required for the pumped-in solution to reach 95% of the initial reservoir concentration.  Since our device has a smaller stroke volume in the same magnetic field compared to the device in  [34] (for instance, 2.3 μl compared to 2.8 μl in a 230 mT magnetic field  [34]) the volume of the pumped-in solution after removing the magnetic field is smaller. Therefore the required mixing time for our device is smaller than 200 seconds. However, we use 200s mixing time for the in-vitro release studies. This over-estimated mixing time ensures consistent release rates of DTX per actuation. For the ex-vivo bladder implantation we used 10s on / 10s off actuation mode and examined the total released drug qualitatively after two hours.   40  Chapter 3: Results and Discussions  3.1 MB Controlled Release PDMS devices were used to study the controlled release of the model drug MB. The actuation setup for MB was previously discussed in chapter 2. MB has a very high solubility in water and concentrations up to 40 g/L were reported by the manufacturing companies. Therefore, it was expected to see a reduction in the concentration of the reservoir solution after few actuations, and hence, a decline in the amount of released MB following consecutive actuations. In order to avoid this and to make the device’s release rate as consistent as possible, MB loaded devices were actuated in low magnetic fields (32.9 mT), to minimize the amount of released MB and to ensure that solid drug lasts in the device for longer periods. It is worth noting that drugs with high solubility in water would not be an ideal candidate for this device, however, MB has favorable characteristics such as ease of measurement and it is non-radioactive which makes its handling very easy. Therefore it might be a good candidate as a model drug to show proof-of-concept of the device. Figure  3.1 shows the cumulative release profile of the MB from the device for the period of three days and the average release rate of MB per actuation interval. Each actuation interval consists of 5 consecutive actuation cycles, and each cycle consists of 50s actuation time and 200s relaxation time. Therefore each actuation interval is 17.5 minutes. For MB, the minimum non-actuation period between each actuation interval is 60 minutes. This time is required so that the concentration of the solution inside the reservoir returns to saturation after each actuation interval. The 50s release time is longer than the estimated release time in a 32.9 mT magnetic field shown in Table  2.2, but it ensures all of the MB solution is released in each actuation cycle. 41  The device has an average release rate of 27.61±0.79 μg per actuation interval in a 32.9 mT magnetic field. The small standard deviation compared to the average release rate highlights the consistency of the released MB rate. The leakage of MB from the aperture of the device (i.e. background diffusion) for the same period of one actuation interval (i.e. 17.5 minutes) is 1.42±0.16 μg which is almost 20 times smaller than the average release rate during the actuation period.  It should be noted that the release rate of drugs at the onset of actuation is very high. Therefore, the devices had to go through an initial actuation step called priming (i.e. 10s on / 10s off actuation for two hours) to become stable. In the case of MB, priming results in rapid depletion of solid MB from the reservoir due to high solubility and therefore all solid MB could be gone during this step. Therefore, instead of priming the devices, which could have led to significant release of deposited MB, the amount of the released MB was measured from the beginning. The device went through 15 actuation intervals for three days (5 intervals each day), but the data from the last 7 intervals were used to measure the average release and background diffusion of the device since the device became more stable after the first 8 actuation intervals (i.e. constant release rate per actuation interval with less than 3% standard deviation from the average release rate). The cumulative release chart however contains all the measurements from the beginning to the end of MB release experiments.  42   Figure ‎3.1 (a) Cumulative MB release profile over 3 days in a 32.9 mT magnetic field and (b) average MB release rate per actuation interval in a 32.9 mT magnetic field. 43  3.2 DTX Controlled Release In contrast with MB, DTX has a very low solubility in water. It’s also an antiproliferative drug widely used in chemotherapy, which makes it an ideal drug candidate for this device. The release kinetics of DTX over a period of 11 days was studied and the results are presented here. DTX was prepared by mixing tritium-labeled DTX with normal DTX in a 50% ethanol and 50% dichloromethane (DCM) solution. 160 μl of tritium-labeled docetaxel in ethanol (activity 1 mCi/mL) was mixed with 12.8 mg of DTX dissolved in 160 μg of DCM, resulting in a drug concentration of ~40 mg/mL. 2 mg of MB powder was also used to add color to the mixture for visual aid. 10 μl of ethanol was added to the solution to compensate for the evaporated solvents when handling the solution. 800 μg of DTX was loaded into a PDMS device. The device was filled with 4% w/v BSA in PBS solution and incubated at 37 °C over night. After incubation, the device was inspected for any remaining bubbles inside the reservoir then glued to the bottom of a petri dish that was later filled with 10 mL of sterile 1% w/v BSA in PBS solution. The petri dish was placed on the height-adjustable stage with a magnet-to-membrane distance of 6 mm resulting in a 135.7 mT magnetic field at the magnetic block location.  The release kinetics of the device was studied over a period of 11 days and the cumulative released DTX profile is demonstrated in Figure  3.2. The device went through five actuation intervals per day, each consisting of five actuation cycles. The actuation time (50s) and the mixing time (200s) were selected slightly more than the estimations to ensure that the DTX release rate remains constant during the actuation period. At the end of each actuation interval, the solution inside the petri dish was stirred to become uniform and then three samples with different volumes were collected from the solution and analyzed by a LS 6500 series, 44  multipurpose scintillation counter (Beckman Coulter Inc., Brea, CA, USA). The total amount of the released DTX was calculated based on the average DPM (disintegration per minute) values of the samples. The smallest DPM value was noticeably larger than the background radiation of the environment. Moreover, DPM values of the samples were linearly proportional to the volume of the collected samples, indicating a good level of reliability for the measurements. The entire amount of released DTX after 11 days was ~17 μg, accounting for 2% of the initial deposited solid DTX in the reservoir. Therefore, a device with 800 μg of deposited DTX should theoretically have enough drugs to release over 1.5 years in a 135.7 mT magnetic field with the same actuation parameters used in this study.  Figure ‎3.2 Cumulative DTX release profile in a 135.7 mT magnetic field over 11 days.  45  3.3 Tissue Implantation In order to see how the device works when it is implanted inside a tissue, and to demonstrate the feasibility of implanting the device through a needle, a qualitative study with the model drug MB was undertaken. Smaller 3D printed devices were used for this experiment as the PDMS devices were too large for implantation through a needle. 3D printed devices are 2mm in diameter and 12 mm long, and can be implanted into the tissue through a gauge 12 needle (inner diameter of the needle is 2.16 mm). Two devices were filled with 4% BSA solution and implanted into fresh bladder tissue slices of a swine (3-4 mm thick). Each tissue was then placed in a petri dish and filled with Tyrone’s solution. The petri dish was placed on the height-adjustable stage and actuated in a 206.3 mT magnetic field. The actuation cycle for this experiment was 10s on / 10s off for 2 hours. Another device was implanted inside the bladder tissue but it was not actuated. This served as a control sample to be compared with the actuated device after 2 hours. After actuation, the tissues were cut open and visually inspected for the released MB. Figure  3.3 shows the implantation of the device with a blunt tip needle and the released MB in the actuated sample and the control sample with no actuation. The released MB from the actuated device is clearly noticeable while the non-actuated device has no signs of released MB.  46   Figure ‎3.3 (a) Device implantation in swine bladder tissue with a needle, (b) control sample with no actuation after two hours and (c) released MB after two hours of actuation in a 206.3 mT magnetic field. 47  Chapter 4: Conclusions and Future Work  The availability of therapeutic compounds and their concentrations at the disease site play a crucial role in the efficacy of drug therapy. Recently, efforts have been geared towards the development of controlled release systems that are capable of controlling drug release rate and the time of release, getting one step closer to the realization of the personalized medicine concept, in which there would be a unique drug release profile specific to each patient based on their conditions. In this thesis we described the design and fabrication of a minimally-invasive MEMS drug delivery implant for the treatment of prostate cancer and presented proof-of-concept release studies. The device releases docetaxel (an anti-proliferative drug widely used in chemotherapy) on demand with the application of an external magnetic field. Device implantation into the prostate is done through a needle. This minimally invasive procedure is similar to the radioactive seed implantation in the prostate in brachytherapy for the treatment of localized prostate cancer. The work in this study included an introductory chapter on the prostate cancer, current treatments and current active drug delivery systems. It was followed by the second chapter describing the design, fabrication and characterization of the device. Drug release studies and ex-vivo implantation results and discussions were presented in chapter three. This final chapter summarizes the work, highlights some concluding remarks and gives future directions.  4.1 Summary Chapter 1 briefly reviews prostate cancer and its current treatments. Controlled drug delivery using MEMS implants was suggested as an alternative treatment for prostate cancer with 48  potentially fewer side-effects than current treatments. The implants were grouped in two categories; active and passive systems. Active systems were briefly reviewed and further categorized into reservoir-based systems and micropumps. Finally, magnetic actuation was proposed to remotely trigger drug release from the implanted devices due to its advantages. In chapter 2, the design of the device was explained first. In order to implant the device in prostate, the device had to be designed in a way that it could be actuated in longer distances from the external magnet (i.e. smaller field gradients). In order to achieve large deflections in smaller field gradients, a magnetic block in which all the magnetic particles were gathered, was attached to a thin PDMS membrane with a laser-drilled aperture. The membrane sealed a drug-loaded reservoir and deflected under the application of an external magnetic field, therefore discharging the drug solution inside the reservoir through the aperture. Additional housing prevented the surrounding tissues from touching the membrane. Two types of devices (PDMS and 3D printed) were fabricated and the step-by-step fabrication process was presented in chapter 2. PDMS devices were easier to handle due to their larger size and they were used for drug release experiments. However they were too large to be implanted through a needle; so smaller devices were fabricated using 3D printing technology. These devices were more rigid and unlike PDMS devices, they did not deform under external loads. They were more suitable for demonstrating the feasibility of the implantation of the device through a needle and the operation of the device when implanted in a tissue. The PDMS device was used to characterize the membrane deflection under various magnetic fields and to calculate two time constants for continuous actuation of the device. One of the most important characters of this device is the novel actuation mechanism used to maximize the magnetic force. In the previous device from our group  [34], the amount of magnetic particles that 49  could be uniformly dispersed into the PDMS matrix was limited. Dispersing more particles into the membrane could result in the agglomeration of particles and hence, a non-uniform membrane with a rough surface that was unable to form a leakage-free seal with the reservoir. Since the amount of available magnetic force is proportional to the amount of magnetic particles inside the magnetic PDMS membrane, magnetic forces were only large enough to deflect the membrane at a close proximity to the permanent magnet, where large field gradients exist. This limited the application of the previous device  [35] to the implantation sites close to the skin and the outside of the body (e.g. back of the eye). Increasing the thickness of the membrane could increase the amount of magnetic particles inside the membrane, however, from a mechanical aspect, a thicker membrane requires a much larger force to deflect. Therefore, there was a trade-off between the thickness of the membrane and the amount of magnetic particles inside the membrane. The new actuation mechanism separates the magnetic force generating part (i.e. magnetic block) from the deflection generating part (i.e. membrane) and these two do not interfere with each other’s tasks. This allows us to embed more magnetic particles into the magnetic block while minimizing the thickness of the membrane; both of which lead to larger deflections in smaller magnetic field gradients. So the new device is a better candidate for deeper implantation and could be used to treat or slow down the growth of localized prostate cancer. Pirmoradi et al.  [36] reported the fabrication of a uniformly dispersed magnetic PDMS membrane with a 40% w/w ratio of coated iron-oxide magnetic particles (EMG 1400) to PDMS. This is among the highest ratios reported for a flexible structure using solvent casting technique [44-47]. In  [36] the nature of the particle coating was not disclosed, but the weight content of iron-oxide in the particles was 80%. A maximum weight loading of 40% particles in PDMS was 50  achieved in  [36], resulting in a 32% weight concentration of magnetic content (i.e. iron-oxide) in the composite. We can calculate the weight percentage of magnetic content in our magnetic film and compare it with the previous results. The saturation magnetization of our film (Ms) is 35.4 emu/g at 40 kOe magnetic field. The typical magnetization of the particles is 63 emu/g (factory data). Assuming that PDMS has negligible contribution to the magnetization of the film, the weight percentage of the magnetic content inside the magnetic film would be 56.2%. This means that for a same volume of magnetic film (current work) and the magnetic PDMS in  [36], the magnetic film used in this design can generate 75% more magnetic force in the same actuation setup. It is worth noting that since the thickness of the magnetic film in our design does not interfere with the deflection of the membrane, multiple layers of magnetic particles can be deposited into the film, increasing the magnetic content of the film. Moreover, unlike previous fabrication techniques that relied on uniform dispersion of the particles in a polymer (e.g. solvent casting) using surfactant, the new fabrication technique embeds dry magnetic particles between PDMS layers and therefore it does not require particle coating. The absence of surfactant means there would be more magnetic content in a given volume of particles (98% iron-oxide content in the particles used in this study compared to 80% iron-oxide content in the particles used in  [36]), resulting in an increase in the amount of magnetic force. Uncoated particles with much higher magnetization than iron-oxide (e.g. pure iron particles with a saturation magnetization of 210 emu/g  [48]) can also be used in order to increase the magnetic force. Chapter 2 was followed by the estimation of the membrane deflection using COMSOL multiphysics software. The estimated values were then used to calculate the required mixing time and release time to ensure continuous actuation of the device. Table  4.1 shows a 3.6 times 51  increase in the magnetic force and a 1.8 times increase in the actuation distance between this device and the device from  [34] with the similar dimensions. In chapter 3, drug release experiments were presented for a model drug MB and the anti-proliferative drug DTX. The cumulative MB release profile over 3 days and cumulative DTX release profile over 11 days were demonstrated. The device has an average release rate of 27.61±0.79 μg per actuation interval for MB in a 32.9 mT magnetic field. When the device is not actuated, MB background leakage from the aperture for the same period of one actuation interval is 1.42±0.16 μg; almost 20 times smaller than the released amount during the actuation period. The average release rate of DTX in a 135.7 mT magnetic field is 353±36 ng per actuation interval. In chapter 3, device implantation and actuation in the swine bladder tissue was demonstrated. The experiment demonstrated the feasibility of implantation and the operation of the device inside the tissue qualitatively. Device Magnetic Field (mT) Magnetic Force (mN) Distance from Magnet (mm) Current Device 56.8 0.5 11 Current Device 111.8 1.8 7 Device in  [34] 111.5 0.5 6 Table ‎4.1 Comparison between the magnetic force and the actuation distance between the current device and the device from ‎[34].  4.2 Future Work One future direction would be to optimize the device in terms of the available magnetic force and deflection. Uncoated iron particles can be used to make the magnetic block. Iron particles possess stronger magnetization values compared to the iron-oxide particles when placed in the 52  same magnetic field. The number of layers in the magnetic block could be increased to accommodate more particles in the PDMS matrix.  Finite element simulations show that the deflection of the membrane depends on the geometry of the magnetic block. Generally, a longer and narrower block leads to more stroke volume (a narrower block increases the distance between the load and the clamped edge of the membrane, causing more deflection); however making the block narrower means the total amount of particles in the block would be smaller (hence the magnetic force). Therefore, the shape of the magnetic block has to be optimized to achieve the largest magnetic forces and deflections in small magnetic field gradients. Instead of using a soft-magnetic material for the magnetic block, a permanent NdFeB magnet could be attached to the membrane which increases the magnetic force substantially in small magnetic field gradients. Another direction is to coat the 3D printed devices with a biocompatible material to allow permanent implantation of the device inside the body. Thin layers of gold or other biocompatible metals could be a potential solution, but such physical surface modifications would not be a very good solution for long term implantations as they could wear off over time. Chemical surface modification of these devices with a biocompatible coating would be a more suitable solution for such applications. One of the focuses in this study was to improve the actuation mechanism within the device itself. However, the actuator can be designed in a way that maximizes the magnetic field gradient at the location of the magnetic block. In this study we used a permanent magnet for actuation. Although these magnets can be used to generate strong magnetic fields, they do not necessarily generate strong field gradients which usually depend on the shape of the magnet. 53  The geometry of the reservoir can be optimized as well. For 3D printed devices, thinner walls lead to a larger reservoir volume which results in larger deflections of the membrane and larger amounts of deposited drugs. However, a thin-wall device would be hard to fabricate and it would also deform during the curing steps. Efforts should be put towards making a 2mm diameter device with the largest achievable volume of the reservoir and the thinnest walls. The next step in the experiments would be to characterize drug release from an optimized and biocompatible 3D printed device. Once characterized, these devices can be implanted in mice with tumors. After actuation over several days, the tumors will be removed from the mice and examined. These would be the first steps in animal studies using this new implant.  54  References  [1] R. Siegel, D. Naishadham and A. Jemal, “Cancer Statistics, 2013,” CA: A Cancer Journal for Clinicians, vol. 63, pp. 11-30, 2013. [2] A. J. Roth, M. I. Weinberger and C. J. Nelson, “Prostate Cancer: Quality of Life, Psychosocial Implications and Treatment Choices,” Future Oncology, vol. 4, pp. 561-568, 2008. [3] N. Mottet, P.J. Bastian, J. Bellmunt, R. C. N. van den Bergh, M. Bolla, N .J. van Casteren, P. Cornford, S. Joniau, M. D. Mason, V. Matveev, T. H. van der Kwast, H. van der Poel, O. Rouvière and T. Wiegel “Guidelines on Prostate Cancer,” European Association of Urology, 2014. [4] F. J. Fowler Jr, M. J. Barry, G. Lu-Yao, J. Wasson, A. Roman and J. Wennberg, “Effect of radical prostatectomy for prostate cancer on patient quality of life: results from a medicare survey,” Urology, vol. 45, pp. 1007-1015, 1995. [5] F. J. Fowler Jr., M. J. Barry, G. Lu-Yao, J. Wasson and L. Bin, “Outcomes of external-beam radiation therapy for prostate cancer: a study of Medicare beneficiaries in three surveillance, epidemiology, and end results areas,” Journal of Clinical Oncology, vol. 14, pp. 2258-2265, 1996. [6] National Cancer Institute, "Treatment choices for men with early-stage prostate cancer," NIH publication, No. 11-4659, 2011. [7] S. S. Sridhar, S. J. Freedland, M. E. Gleave, C. Higano, P. Mulders, C. Parker, O. Sartor and F. Saad, “Castration-resistant prostate cancer: from new pathophysiology to new treatment,” Europeam Urology, vol. 65, pp. 289-299, 2014. 55  [8] A. Michael, K. Syrigos and H. Pandha, “Prostate cancer chemotherapy in the era of targeted therapy,” Prostate Cancer and Prostatic Diseases, vol. 12, pp. 13-16, 2008. [9] G. Gaucher, R. H. Marchessault, and J.C. Leroux, "Polyester-based micelles and nanoparticles for the parenteral delivery of taxanes," Journal of Controlled Release, vol. 143, pp. 2-12, 2010. [10] C. D. Scripture and W. D. Figg, “Drug interactions in cancer therapy,” Nature Reviews Cancer, vol. 6, pp. 546-558, 2006. [11] F. Pirmoradi, “A battery-less MEMS device for on-demand and controlled drug release,” PhD. Dissertation, Department of Mechanical Engineering, University of British Columbia, Vancouver, B.C., 2011. [12] F. K. Engels, W. J. Loos, J. M. van der Bol, P. de Brujin, R. H. J. Mathijssen, J. Verweij and R. A. A. Mathot, “Therapeutic Drug Monitoring for the Individualization of Docetaxel Dosing: A Randomized Pharmacokinetic Study,” Clinical Cancer Research, vol. 17, pp. 353-362, 2011. [13] W. B. Liechty, D. R. Kryscio, B. V. Slaughter and N. A. Peppas, “Polymers for drug delivery systems,” Annual Review of Chemical and Biomolecular Engineering, vol. 1, pp. 149-173, 2010. [14] C. L. Stevenson, J. T. Santini Jr. and Robert Langer, “Reservoir-based drug delivery systems utilizing microtechnology,” Advanced Drug Delivery Reviews, vol. 64, pp. 1590-1602, 2012. [15] Thomson Physicians' Desk Reference, 63th Edition Thomson Publishing, Montvale, N.J, 2009. 56  [16] A. Göpferich, "Mechanisms of polymer degradation and erosion," Biomaterials, vol. 17, pp. 103-114, 1996. [17] R. Yoshida, "Design of functional polymer gels and their application to biomimetic materials," Current Organic Chemistry, vol. 9, pp. 1617-1641, 2005. [18] B. P. Timko, T. Dvir, and D. S. Kohane, "Remotely triggerable drug delivery systems," Advanced Materials, vol. 22, pp. 4925-4943, 2010. [19] M. Staples, “Microchips and controlled-release drug reservoirs,” Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology, vol. 2, pp. 400-417, 2010. [20] J. T. Santini Jr., M. J. Cima and R. S. Langer, “Method for making microchip reservoir device,” United States Patent, US20070047926A1, Feb. 2008. [21] J. T. Santini Jr., M. J. Cima and R. S. Langer, “Microchip drug delivery devices,” United States Patent, US007070590B1, Jul. 2006. [22] J. T. Santini Jr., M. J. Cima and R. S. Langer, “Medical device with array of electrode-containing reservoirs,” United States Patent, US007070592B2, Jul. 2006. [23] J. T. Santini Jr., C. E. Hutchinson, S. A. Uhland, M. J. Cima, R. S. Langer and D. Ausiello, “Microfabricated devices for the delivery of molecules into a carrier fluid,” United States Patent, US006537256B2, Mar. 2003. [24] A. C. R. Grayson, R. S. Shawgo, A. M. Johnson, N. T. Flynn, Y. Li, M. J. Cima, and R. Langer, “A BioMEMS Review: MEMS Technology for Physiologically Integrated Devices,” Proceedings of the IEEE, vol. 92, pp. 6-21, 2004. [25] D. A. LaVan, T. McGuire and Robert Langer, “Small-scale systems for in vivo drug delivery,” Nature Biotechnology, vol. 21, pp. 1184-1191, 2003. 57  [26] J. T. Santini, M. J. Cima, and R. Langer, "A controlled-release microchip," Nature, vol. 397, pp. 335-338, 1999. [27] J. M. Maloney, S. A. Uhland, B. F. Polito, J. N. F. Sheppard, C. M. Pelta, and J. J. T. Santini, "Electrothermally activated microchips for implantable drug delivery and biosensing," Journal of Controlled Release, vol. 109, pp. 244-255, 2005. [28] J. H. Prescott, S. Lipka, S. Baldwin, N. F. Sheppard, J. M. Maloney, J. Coppeta, B. Yomtov, M. A. Staples, and J. T. Santini, "Chronic, programmed polypeptide delivery from an implanted, multireservoir microchip device," Nature Biotechnology, vol. 24, pp. 437-438, 2006. [29] D. Ge., X. Tian, R. Qi, S. Huang, J. Mu, S. Hong, S. Ye, X. Zhang, D. Li and W. Shi, “A polypyrrole-based microchip for controlled drug release,” Electrochimica Acta, vol. 55, pp. 271-275, 2009. [30] N. M. Elman, H. L. Ho Duc and M. J. Cima, “An implantable MEMS drug delivery device for rapid delivery in ambulatory emergency care,” Biomedical Microdevices, vol. 11, pp. 625–631, 2009. [31] R. Lo, K. Kuwahar, P. Y. Li, S. Saati, R. N. Agrawal, M. S. Humayun and E. Meng, “A passive MEMS drug delivery pump for treatment of ocular diseases,” Biomedical Microdevices, vol. 11, pp. 959–970, 2009. [32] P. Y. Li, J. Shih, R. Lo, S. Saati, R. Agrawal, M. S. Humayun, Y. C. Tai and E. Meng, “An electrochemical intraocular drug delivery device,” Sensors and Actuators A: Physical, vol. 143, pp. 41–48, 2008. [33] F. Amirouche, Y. Zhou and T. Johnson, “Current micropump technologies and their biomedical applications,” Microsystem Technologies, vol. 15, pp. 647-666, 2009. 58  [34] F. N. Pirmoradi, J. K. Jackson, H. M. Burt and M. Chiao, “A Magnetically Controlled MEMS Device for Drug Delivery: Design, Fabrication, and Testing,” Lab on a Chip, vol 11, pp. 3072-3080, 2011. [35] F. N. Pirmoradi, J. K. Jackson, H. M. Burt, M. Chiao, “On-demand Controlled Release of Docetaxel from a Battery-less MEMS Drug Delivery Device,” Lab on a Chip, vol. 11, pp. 2744-2752, 2011. [36] F. N. Pirmoradi, L. Cheng, and M. Chiao, "A magnetic poly(dimethylesiloxane) composite membrane incorporated with uniformly dispersed, coated iron oxide nanoparticles," Journal of Micromechanics and Microengineering, vol. 20, p. 015032, 2010. [37] F. K. Engels, A. Sparreboom, R. A. A. Mathot and J. Verweij, “Potential for improvement of docetaxel-based chemotherapy: a pharmacological review,” British Journal of Cancer, vol. 93, pp. 173-177, 2005. [38] J. J. Abbott, O. Ergeneman, M. P. Kummer, A. M. Hirt, and B. J. Nelson, "Modeling magnetic torque and force for controlled manipulation of softmagnetic bodies," IEEE Transactions on Robotics, vol. 23, pp. 1247-1252, 2007. [39] F. Schneider, T. Fellner, J. Wilde and U. Wallrabe, “Mechanical properties of silicones for MEMS” Journal of Micromechanics and Microengineering, vol. 18, 065008, 2008.  [40] K. M. Choi and J. A. Rogers, “A photocurable poly(dimethylsiloxane) chemistry designed for soft lithographic molding and printing in the nanometer regime,” Journal of the American Chemical Society, vol. 125, pp. 4060-4061, 2003. 59  [41] Y. S. Yu and Y. P. Zhao, “Deformation of PDMS membrane and microcantilever by a water droplet: Comparison between Mooney-Rivlin and Linear elastic constitutive models,” Journal of Colloid and Interface Science, vol. 332, pp. 467-476, 2009. [42] A. Folch, Introduction to BioMEMS, 1st ed., CRC Press, 2012. [43] V. Studer, G. Hang, A. Pandolfi, M. Ortiz, W. F. Anderson and S. R. Quake, “Scaling properties of a low-actuation pressure microfluidic valve” Journal of Applied Physics, vol. 95, pp. 393-398, 2004. [44] J. le Digabel, N. Biais, J. Fresnais, J. F. Berret, P. Hersen, and B. Ladoux, “Magnetic micropillars as a tool to govern substrate deformations,” Lab on a Chip, vol. 11, pp. 2630–2636, 2011. [45] N. J. Sniadecki, A. Anguelouch, M. T. Yang, C. M. Lamb, Z. Liu, S. B. Kirschner, Y. Liu, D. H. Reich, and C. S. Chen, “Magnetic microposts as an approach to apply forces to living cells,” Proceedings of the National Academy of Sciences, vol. 104, pp. 14553–14558, 2007. [46] N. J. Sniadecki, C. M. Lamb, Y. Liu, C. S. Chen, and D. H. Reich, “Magnetic microposts for mechanical stimulation of biological cells: Fabrication, characterization, and analysis,” Review of Scientific Instruments, vol. 79, pp. 044302-1–044302-8, 2008. [47] F. Fahrni, M. W. J. Prins, and L. J. van Ijzendoorn, “Micro-fluidic actuation using magnetic artificial cilia,” Lab on a Chip, vol. 9, pp. 3413–3421, 2009. [48] F. Khademolhosseini and M. Chiao, “Fabrication and patterning of magnetic polymer micropillar structures using a dry-nanoparticle embedding technique,” Journal of Microelectromechanical Systems, vol. 22, pp. 131-139, 2013. 60  [49] S. P. Timoshenko and S. Woinowsky-Krieger, Theory of Plates and Shells, 2nd ed., McGraw-Hill, 1959. 61  Appendices  Appendix A  : Actuation Setup   Figure A.1 Actuation setup.    Height-adjustable stage Device in solution Arm Motor Magnet Holder Magnet 62  Appendix B  : COMSOL Simulation Verification  In order to demonstrate the agreement between COMSOL’s membrane deflection simulations and the analytical approximate solution, we consider a uniformly loaded circular magnetic PDMS membrane presented in  [34] and compare the results for both COMSOL simulations and the approximate solution. The membrane has a diameter of 6 mm, thickness of 40 μm, elastic modulus of 1 MPa and Poisson’s ratio of 0.25. When the deformation of a thin plate becomes large compared to the thickness of the plate, the strain of the middle plain should be considered in the analytical solution. In such case, the displacement at the center of the membrane is given by the following equation  [49]:  𝜔0 = 0.662𝑎√𝑝𝑎𝐸𝑡3 (B.1) In which 𝜔0 is the deflection at the center of the membrane, 𝑎 is the membrane radius, 𝑝 is the uniformly distributed load, 𝐸 is the elastic modulus and 𝑡 is the membrane thickness. The results of analytical solution and COMSOL simulations are shown in Table B.1. As can be seen, both results are in very good agreement.   Distance from Magnet (mm) Magnetic Field (mT) Magnetic Force (mN) Analytical 𝜔0 (μm) COMSOL 𝜔0 (μm) 6 111.5 0.5 219 213 4.1 172 1 277 271 2.8 230 1.5 315 311 1.8 295 2 346 343 Table B.1 Comparison between the simulated deflection and the approximate analytical solution for the magnetic PDMS membrane in ‎[34] with‎Φ = 6mm‎and‎t‎=‎40‎μm.  

Cite

Citation Scheme:

        

Citations by CSL (citeproc-js)

Usage Statistics

Share

Embed

Customize your widget with the following options, then copy and paste the code below into the HTML of your page to embed this item in your website.
                        
                            <div id="ubcOpenCollectionsWidgetDisplay">
                            <script id="ubcOpenCollectionsWidget"
                            src="{[{embed.src}]}"
                            data-item="{[{embed.item}]}"
                            data-collection="{[{embed.collection}]}"
                            data-metadata="{[{embed.showMetadata}]}"
                            data-width="{[{embed.width}]}"
                            data-media="{[{embed.selectedMedia}]}"
                            async >
                            </script>
                            </div>
                        
                    
IIIF logo Our image viewer uses the IIIF 2.0 standard. To load this item in other compatible viewers, use this url:
https://iiif.library.ubc.ca/presentation/dsp.24.1-0135633/manifest

Comment

Related Items