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Doctoral Thesis
Fabrication and functionalization of magnetic helical microrobotsfor potential biomedical applications
Author(s): Qiu, Famin
Publication Date: 2015
Permanent Link: https://doi.org/10.3929/ethz-a-010532942
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ETH Library
DISS. ETH NO. 22671
FABRICATION AND FUNCTIONALIZATION OF MAGNETIC HELICAL MICROROBOTS FOR POTENTIAL BIOMEDICAL APPLICATIONS
A thesis submitted to attain the degree of
DOCTOR OF SCIENCES of ETH ZURICH
(Dr. sc. ETH Zurich)
presented by
FAMIN QIU
M. Eng. in Zhejiang University, China
born on 17.01.1985
citizen of China
accepted on the recommendation of
Prof. Dr. Bradley J. Nelson, examiner Prof. Dr. Li Zhang, co-examiner
2015
i
“Man's dearest possession is life.
It is given to him but once,
and he must live it so as to feel no torturing regrets for wasted years,
never know the burning shame of a mean and petty past.”
“人最宝贵的是生命。
生命每个人只有一次。
人的一生应当这样度过：回忆往事，他不会因为虚度年华而悔恨，
也不会因为生活庸俗而羞愧。”
- Nikolai Alexeevich Ostrovsky
from “How the Steel Was Tempered”
ii
Acknowledgements
This dissertation presents the cumulative knowledge and results of my doctoral study in ETH
Zurich. During the past three and half years in the Institute of Robotics and Intelligent Systems
(IRIS) lab, I got a lot of support and help from both people who belong to and people who are
outside of the lab. Without their help, I would not have managed to write this dissertation.
First of all, I would like to give special thanks to my supervisor, Professor Bradley J. Nelson,
who offered the great opportunity for me to work in this group and in this exciting area,
microrobotics. He always encouraged and supported me when I proposed ideas to him. His
useful suggestions, such as, “to look at the big picture of your research”, “be nice to people and
work hard”, have led me through problems both in my research and in my personal life.
I would like to thank my co-examiner, Professor Li Zhang, who has supported me continuously,
especially in the first one and a half years of my studies. In the research, Li guided me into this
interesting topic and taught me gradually about how to do research. His enthusiasm for science
and hard-working attitude inspire and encourage me. Li has played multiple roles in my life,
as an advisor, a friend and a big brother.
This dissertation would not have been possible without the collaboration and support of
colleagues in my lab, university students who worked in the project and world-wide
collaborators. I would like to thank my fellow PhD students, Kathrin E. Peyer, Soichiro Tottori,
Tianyun Huang, Simone Schurle and a pervious master student, Erdem Siringil, for their help
on the fabrication and swimming characterization of microrobots. I would like to thank Klaus
Marquardt (University of Zurich), Krzysztof K. Krawczyk, Marco Casarosa (ETH Zurich),
Prof. Alfredo Franco-Obregon (National University of Singapore, Singapore) and Prof.
Hongsoo Choi (DGIST, South Korea) for their help on the cytotoxicity of microrobots. I am
also thankful to Rami Mhanna, Benjamin Simona, Christopher Millan, Prof. Marcy Zenobi-
Wong, Prof. Janos Vörös from ETH Zurich, Dr. Satoshi Fujita (from AIST, Japan), Stefano
Fusco and Dr. Salvador Pané i Vidal from IRIS for their help on the functionalization of
microrobots. I would like to thank Franziska Ullrich, Dr. Roel Pieters (IRIS, ETH Zurich), Dr.
Ania Servant and Prof. Kostas Kostarelos (University of Manchester, UK) for their support and
collaboration on the in vivo tracking of microrobots. I also thank the students I supervised for
their outstanding work, namely Yun Ding, Kourosh Schneeberger, Meijun Liu, Jan-Philipp
iii
Peyer, Geevan Punnackalkilukken and Baris Atakan.
Many members in the lab have not only helped my research, but also my personal life during
my studies. Besides the people mentioned above, I would like to thank: Brigitte Geissmann
and Kerstin Degen for their support on paper work, Dimitris Felekis, Taylor Newton and Dr.
Roel Pieters for their IT support, and Dr. Salvador Pané i Vidal, Dr. Mahmut Selman Sakar,
Dr. Olgac Ergeneman, Dr. Chengzhi Hu and Dr. Xiangzhong Chen for their helpful scientific
discussions. I would like to thank the PhD students (Franziska Ullrich, Stefano Fusco, Taylor
Newton, George Chatzipirpiridis, Bumjin Jang, Hen-Wei Huang and Naveen Shamsudhin)
who have shared office space with me and made it an enjoyable environment. Thank Juho
Pokki, André Lindo, Hsi-Wen Tung and Berna Özkale for the help in the chemical lab B11. I
would like to thank Jo Walker for improving the English in this dissertation and thank
Franziska Ullrich for her translation of the English abstract into German. I would like to thank
the FIRST lab team of ETH Zurich for technical support on the micro-fabrication, and also the
financial funding of my project, including the Swiss National Science Foundation (SNSF)
Project NO. 200021-130069, the European Research Council Advanced Grant “BOTMED”
and the Sino-Swiss Science and Technology Cooperation (SSSTC, Grant No.
IZLCZ2_138898) Grant.
Last but not least, I would like to deeply thank my parents and my wife. My parents always
give me support selflessly and encourage me to go through problems in my life. They are
greatest parents in the world. Special thanks to my wife, Mengfei, for her company. She takes
care of our little daughter and most of the housework without any complaint so that I am able
to concentrate on my research. There are, of course, many other people who I have neglected
to mention individually here and I extend my thanks to all those people without whom I would
not be who I am today.
iv
Abstract
Mobile microrobots show potential applications in various fields, especially in the biological
and medical fields. Due to their small size and mobility, they are promising tools for minimally
invasive surgery, cell manipulation and analysis, and targeted therapy. Magnetic helical
microrobots, referred to as artificial bacterial flagella (ABFs), are one kind of mobile
microrobot. Inspired by the flagellar propulsive motion of the bacteria E. Coli, these ABFs can
perform 3D navigation in a controllable fashion with micrometer precision under low-strength
rotating magnetic fields (< 10 mT). They are promising tools for biomedical applications, such
as targeted drug delivery in vitro and in vivo. This dissertation focuses on using these ABFs for
potential targeted therapies, and the work includes the fabrication, wireless actuation,
biocompatibility, biomedical functionalization and in vivo localization of these ABFs.
A single ABF consisted of a helical-shaped polymeric body and metallic layers of a magnetic
material and titanium. The helical body with the length ranging from 4 m to 300 m was
fabricated by 3D laser direct writing (a kind of 3D lithography) and the magnetic layer, such
as iron and nickel, and titanium were coated onto the helical body by physical vapor deposition.
The ABFs showed 3D navigation in liquid using a corkscrew motion under rotating magnetic
fields. The layer of titanium improves the biocompatibility of ABFs and ABFs were nontoxic
to mouse myoblast C2C12 cells over three days.
The functionalization of ABFs with drugs is essential to enhance their biomedical performance
for targeted drug therapy. Lipid-based nanoscale drug carriers were successfully functionalized
on the ABFs. The functionalized ABFs (f-ABFs) showed the abilities to be wirelessly steered
to specific sites and release the carried drug models (calcein, a green fluorescent probe, and
DNA) into targeted cells in vitro. For in vivo applications ABFs were functionalized with a
near-infrared dye NIR-797 for tracking. The simultaneous injection of over 80,000 f-ABFs into
a mouse peritoneal cavity, in vivo tracking using near infrared fluorescence, and integrated
wireless control of ABFs using rotating magnetic fields within the mouse were demonstrated.
As only weak magnetic fields are required for actuation with feedback provided by in vivo
tracking, the approach can be used deep within tissue relatively far from an organism’s surface.
v
Zusammenfassung
Mobile Mikroroboter weisen viele potentielle Anwendungsgebiete auf, ins besondere in
biologischen und medizinischen Bereichen. Aufgrund ihrer kleinen Grösse und Wendigkeit
sind sie vielversprechende Werkzeuge für minimal invasive Operationen, Manipulation und
Analyse einzelner Zellen und zielgerichtete Therapien. Eine Art mobiler Mikroroboter sind
magnetische spiralförmige Mikrostrukturen, die Artificial Bacterial Flagella (ABFs). Inspiriert
durch den Flagellenantrieb des E. Coli Bakteriums, können ABFs, angetrieben durch rotierende
schwache Magnetfelder (< 10 mT), kontrollierte dreidimensionale Bewegungen mit einer
Präzision von wenigen Mikrometern ausführen. Diese Strukturen erweisen sich als
aussichtsvolles Werkzeug für biomedizinische Anwendungen, wie zum Beispiel den gezielten
Wirkstofftransport in in vitro und in vivo Versuchen. Der Schwerpunkt dieser Dissertation liegt
auf der Anwendung von ABFs in gezielten Therapien. Diese Arbeit beschreibt die Fabrikation,
die drahtlose Aktivierung, Biokompatibilität der Strukturen, die biomedizinische
Oberflächenfunktionalisierung und in vivo Lokalisierung von ABFs.
Ein einziges ABF besteht aus einem spiralförmigen Polymer Körper und wird mit je einer
Schicht eines magnetischen Metalls und einer Schicht Titanium beschichtet. Der spiralförmige
Körper, mit einer Länge zwischen 4 µm und 300 µm, wird mittels 3D Laserdirektschreiben
(einer Art der 3D Lithografie) hergestellt. Die magnetische Schicht, bestehend aus Eisen oder
Nickel, und die Titanium Schicht werden mittels Abscheidung aus der Dampfphase hergestellt.
In Flüssigkeit untergeht das ABFs eine Korkenzieher-ähnliche Bewegung, angetrieben durch
rotierende Magnetfelder. Die dünne Titanium Beschichtung verbessert die Biokompatibilität
des ABFs und die Struktur zeigt sich nichttoxisch gegenüber Maus Myoblasten C2C12 in
einem Zeitraum von drei Tagen.
Die Oberflächenfunktionalisierung des ABFs mit Medikamenten ist eine wesentliche
Voraussetzung zur gezielten Wirkstoffabgabe und steigert den biomedizinischen Nutzen.
ABFs konnten erfolgreich mit Liposomen, Systeme zur Wirkstoffabgabe im Nanobereich,
funktionalisiert werden. Es wurde gezeigt, dass die funktionalisierten ABF (f-ABFs) mittels
Magnetfeldern gezielt zu Zellen navigiert und einen Modell-Wirkstoff (Calcein, grüne
fluoreszierende Farbe und DNA) in vitro abgeben konnten. Um ABFs in vivo anzuwenden,
müssen diese visuell verfolgt und lokalisiert werden während sie sich im lebenden Körper
befinden. Die ABFs wurden mittels einer nah-infraroten Farbe NIR-797 für in vivo
vi
Lokalisierung funktionalisiert. Mehr als 80.000 f-ABFs wurden unter das Bauchfell einer Maus
injiziert und konnten unter nah-infrarot Fluoreszenz in vivo lokalisiert werden. Die Bewegung
der f-ABFs wurde mittels rotierender Magnetfelder gesteuert. Da die benötigten Feldstärken
zur Aktivierung der ABFs sehr gering sind, kann diese Methodik auch innerhalb des Körpers
und nicht nur an der Oberfläche der Maus verwendet werden.
vii
Contents
Acknowledgements .................................................................................................................. ii
Abstract .................................................................................................................................... iv
Zusammenfassung.................................................................................................................... v
Contents .................................................................................................................................. vii
List of Tables ........................................................................................................................... ix
List of Figures ........................................................................................................................... x
1 Introduction ...................................................................................................................... 1
1.1 Background ................................................................................................................. 1
1.2 Bio-inspired approaches and contributions of the dissertation ................................... 2
2 Propulsion mechanisms of microrobots ......................................................................... 7
2.1 Swimming at low Reynolds number ........................................................................... 7
2.2 Propulsion mechanisms ............................................................................................... 8
2.3 Magnetic actuation methods ...................................................................................... 14
2.4 Magnetic helical micro/nanorobots ........................................................................... 21
2.5 Conclusion ................................................................................................................. 25
3 Fabrication of ABFs ....................................................................................................... 26
3.1 Introduction ............................................................................................................... 26
3.2 Fabrication of microstructures using DLW ............................................................... 29
3.3 Conclusion ................................................................................................................. 34
4 Motion control of ABFs.................................................................................................. 35
4.1 The torque on an ABF ............................................................................................... 35
4.2 Magnetic setup .......................................................................................................... 36
4.3 Swimming behavior of ABFs .................................................................................... 38
4.4 Conclusion ................................................................................................................. 43
5 Cytotoxicity of ABFs ...................................................................................................... 45
viii
5.1 Introduction ............................................................................................................... 45
5.2 ABFs made from IP-L and SU-8 with Ni/Ti coating ................................................ 46
5.3 ABFs made from ORMOCOMP and Fe/Ti coating .................................................. 49
5.4 Conclusion ................................................................................................................. 56
6 Functionalization of ABFs for potential biomedical applications .............................. 58
6.1 Introduction ............................................................................................................... 58
6.2 ABFs functionalized with liposomes for drug delivery in vitro ................................ 58
6.3 ABFs functionalized with lipoplexes for gene delivery in vitro ............................... 80
6.4 ABFs functionalized with near-infrared dyes for in vivo tracking and actuation ..... 88
6.5 Conclusion ............................................................................................................... 107
7 Summary and future work .......................................................................................... 109
7.1 Summary ................................................................................................................. 109
7.2 Future work ............................................................................................................. 111
References ............................................................................................................................. 112
Appendix ............................................................................................................................... 130
Appendix A ........................................................................................................................ 130
Appendix B ........................................................................................................................ 134
CV .......................................................................................................................................... 135
Publications .......................................................................................................................... 137
ix
List of Tables
Table 1. Fluorescent intensity of f-ABFs measured at varying swimming times. ................... 79
Table 2. The components of the final lipoplex (50 µl in total).. .............................................. 82
Table 3. The coordinates of the center of the clouds. ............................................................ 105
x
List of Figures
Figure 1.1. The field of mobile microrobots is a multi-interdisciplinary research area. ........... 2
Figure 1.2. E. coli bacteria and how they swim. ........................................................................ 3
Figure 1.3. The ABF made by self-scrolling technology........................................................... 4
Figure 1.4. A roadmap for magnetic helical microrobots ......................................................... 5
Figure 2.1. The two-hinged swimmer presented by Purcell. ..................................................... 8
Figure 2.2. The two proposed swimming strategies used by microorganisms .......................... 8
Figure 2.3. Mobile microrobots or micro motors propelled by different mechanisms. ............. 9
Figure 2.4. Three propulsion mechanisms of microorganisms. ............................................... 10
Figure 2.5. Miniature semiconductor diodes powered by an alternating electric field. ........... 12
Figure 2.6. Optical tweezers trap and manipulate small objects. ............................................. 12
Figure 2.7. Nanorobots powered by ultrasound. ...................................................................... 13
Figure 2.8. The typical hysteresis loops of hard and soft magnetic materials. ........................ 16
Figure 2.9. Magnetic fields for actuation of micro/nanorobots. .............................................. 17
Figure 2.10. Micro/nanodevices actuated by rotating magnetic fields. ................................... 18
Figure 2.11. Microrobots actuated by oscillating and on-off magnetic fields. ........................ 19
Figure 2.12. Micro/nanostructures actuated by magnetic fields with gradients. ..................... 20
Figure 2.13. An overview of magnetic helical micro/nanorobots. .......................................... 22
Figure 2.14. Schematic of an ABF. .......................................................................................... 24
Figure 3.1. An overview of fabrication methods for magnetic helical micro/nanorobots. ...... 27
Figure 3.2. Schematic of a direct laser writing tool. ................................................................ 30
Figure 3.3. The principle of DLW based on TPP. ................................................................... 30
Figure 3.4. General fabrication flow using DLW. ................................................................... 31
Figure 3.5. Fabrication of ABFs. ............................................................................................. 32
Figure 3.6. Three-dimensional non-helical microstructures fabricated by DLW .................... 33
Figure 4.1. The principle of motion control of an ABF in 3D ................................................. 36
xi
Figure 4.2. The magnetic setup. ............................................................................................... 37
Figure 4.3. Swimming behavior of ABFs. ............................................................................... 38
Figure 4.4. The wobbling and corkscrew regions of ABF swimming behavior. ..................... 39
Figure 4.5. Schematic of the misalignment angle .................................................................... 40
Figure 4.6. The step-out frequencies of an ABF in different field strengths ........................... 41
Figure 4.7. The swarm control of ABFs. ................................................................................. 43
Figure 5.1. The optical images of cells on different samples after one-day incubation .......... 47
Figure 5.2. Comparison of cell concentrations on different samples ...................................... 48
Figure 5.3. SEM micrographs of cells ..................................................................................... 49
Figure 5.4. Helical microstructures made from ORMOCOMP using DLW ........................... 51
Figure 5.5. The chemical basis of MTT assay ......................................................................... 52
Figure 5.6. Cell viability of Fe-ABFs and FeTi-ABFs. ........................................................... 53
Figure 5.7. Comparison of MTT values with and without cells. ............................................. 54
Figure 5.8. Cell morphology on Fe-ABFs and FeTi-ABFs after one-day incubation. ............ 55
Figure 5.9. Swimming behaviors of Fe-ABFs and FeTi-ABFs. .............................................. 56
Figure 6.1. The structures of lipid vesicles. ............................................................................. 59
Figure 6.2. The trigger-release mechanisms of different liposomes for drug delivery. ........... 60
Figure 6.3. Schematic of functionalization of ABFs with liposomes. ..................................... 60
Figure 6.4. Release of ABFs by sonication. ............................................................................. 61
Figure 6.5. Images of ABF arrays before and after sonication. ............................................... 62
Figure 6.6. A close look at ABFs after sonication. .................................................................. 63
Figure 6.7. Comparison of step-out frequencies of the untethered ABFs ............................... 64
Figure 6.8. SEM image of high and low magnifications of ABFs .......................................... 66
Figure 6.9. Preparation flow for coating ABFs with unilamellar DPPC liposomes. ............... 67
Figure 6.10. The extruder used to produce unilamellar liposomes. ......................................... 68
Figure 6.11. QCM-D signals of DPPC liposome adsorption on a TiO2 crystal....................... 71
Figure 6.12. Fluorescent images of DPPC-coated ABFs. ........................................................ 71
xii
Figure 6.13. Calcein release from DPPC/MSPC functionalized ABFs ................................... 72
Figure 6.14. Quantitative analysis of the calcein signals released from ABFs ........................ 73
Figure 6.15. Calcein release from DPPC/MSPC functionalized TiO2-coated surfaces ........... 74
Figure 6.16. Liposomes adsorb in an intact state to the surface of ABFs. ............................... 77
Figure 6.17. F-ABF swimming and calcein delivery to single cells ........................................ 78
Figure 6.18. Calcein delivery to single cells in vitro. .............................................................. 79
Figure 6.19. QCM-D measurement of the adsorption of the lipoplex ..................................... 84
Figure 6.20. Functionalization of ABFs with the lipoplex. ..................................................... 84
Figure 6.21. Swimming performance of f-ABFs in cell medium. ........................................... 85
Figure 6.22. The transfection efficiency tests of the lipoplex .................................................. 86
Figure 6.23. pDNA transfection and protein expression ......................................................... 87
Figure 6.24. Gene transfection and protein expression in the cells contacting f-ABFs. .......... 87
Figure 6.25. Actuation of a swarm of ABFs in vitro tracked by an optical microscope. ........ 89
Figure 6.26. Actuation and in vivo tracking of a swarm of f-ABFs. ........................................ 92
Figure 6.27. Fabrication and functionalization of ABFs with NIR-797 dyes .......................... 93
Figure 6.28. The f-ABFs characterization. .............................................................................. 95
Figure 6.29. Derivatization of the hydroxyl groups into primary amine groups ..................... 96
Figure 6.30. Fluorescence signals of f-ABFs in different solvents .......................................... 98
Figure 6.31. Controlled swimming of a f-ABF swarm in vitro at 9 mT and 90 Hz. ............. 100
Figure 6.32. In vitro forward velocity of ABFs ..................................................................... 101
Figure 6.33. Data analysis of the swarm swimming in vitro. ................................................ 102
Figure 6.34. A swarm of f-ABFs tracked and actuated in the intra peritoneal cavity ........... 104
Figure 6.35. In vivo movement of a swarm of f-ABFs. ......................................................... 105
Figure 7.1. Summary of the key findings .............................................................................. 110
1
1 Introduction
1.1 Background
1.1.1 Mobile microrobots and their applications
Robots are becoming more and more common and increasingly benefit the quality of our lives.
They have been widely used in manufacturing, space exploration, laboratory research, surgical
procedures and more. Mobile microrobots, a branch of robotics, are miniature robots with a
size ranging from the micrometer (m) to centimeter (cm) scale. They have abilities to access
small spaces down to the microscale, such as inside the human body and micro-channels, and
to interact with and manipulate micro/nanoscale objects. Moreover, due to their small size, they
can be manufactured with a lower cost than big robots and used in a large number to work as a
swarm or a team to produce multi-functionalities.
Mobile microrobots show great potential to be used in various fields due to their small size and
mobility. In biological and medical fields, they are promising tools for minimally invasive
surgery, cell manipulation and analysis, and targeted therapy [1, 2]. In the environmental field,
they show the ability to be used for decontamination and toxicity screening under conditions
too dangerous or too small for humans to access [3, 4]. In microfluidics, they can be used for
manipulation and transportation of micro-objects and chemicals in lab-on-a-chip devices [5,
6].
1.1.2 The challenges of mobile microrobots in biomedical applications
In biomedical applications such as targeted drug delivery, mobile microrobots must: (1) be
small enough to be put in a human body, (2) be able to be actuated in a controlled fashion in a
fluid, (3) be biocompatible with cells and tissues, (4) be localized and tracked, (5) be able to
carry drugs to specific areas and (6) be able to be removed from the human body or
biodegradable after tasks are accomplished. Hence, mobile microrobots in biomedical
applications is a multi-disciplinary task (Figure 1.1), which requires knowledge from many
different fields, such as mechanical engineering for the knowledge of robotic design and
powering, computer science for the knowledge of control and localization, materials science
for the knowledge of microfabrication and characterization, life science for the knowledge of
biocompatibility/safety of devices and functionalization related to bio-applications.
2
There are several challenges to address in order to make mobile microrobots useful for real
biomedical application such as targeted therapies. First, the small size of microrobots requires
one to find suitable technologies in micro/nanofabrication to build the robots since traditional
machining methods are not sufficient to build microstructures and it is difficult to assemble
tiny parts to form a microrobot. Second, the question of how to power microrobots is an issue
since on-board micro-sized batteries do not exist. Additionally, when the size of the robot
decreases to microscale, the inertial force becomes negligible while the drag force from the
liquid dominates. Different actuation methods for making microrobots movable have to be used
compared to the propulsive methods of larger robots. Third, when microrobots are used in
biomedical applications, the biocompatibility of the materials which are used to build the robots
should be taken into account. Further, the functionalization of the tiny robots with therapeutic
drugs and the release of the drugs should be investigated. Finally, the tracking or localization
of the microrobots when they are performing in the human body is another challenge.
Figure 1.1. The field of mobile microrobots is a multi-interdisciplinary research area.
1.2 Bio-inspired approaches and contributions of the dissertation
1.2.1 Bio-inspired approaches
Nature has inspired scientists and engineers to build many useful machines. For example, the
invention of the airplane was inspired by birds, and the creation of the sonar and radar, used to
localize objects in real-time, were inspired by bats. Nature not only inspires us to make
3
machines in the macroscopic scale which can be visualized by the human eye, but also in the
microscopic scale, researchers gained ideas from motile microorganisms to create mobile
micromachines due to the development of microscopes. Towards the end of the 19th century,
researchers established that all motile microorganisms use either flagella or cilia as means for
motion generation. In 1973, the biologist, H. Berg, proved the same that microorganisms, such
as Escherichia coli (E. coli) bacteria, swim in various liquids by rotating their helical flagella
in a helical wave using molecular motors (Figure 1.2) [7].
Figure 1.2. E. coli bacteria and how they swim. (a) Optical image of an E. coli bacteria (K-12 type). Reused from
[8]. (b) E. coli bacteria swim by rotating their flagella in a helical wave. Adapted from [9].
Inspired by this propulsion motion of the flagella , the Nelson group in Switzerland fabricated
the first helical-shaped microrobot mimicking this propulsion method, named artificial
bacterial flagellum (ABF) in 2007 [10] and further studied by Zhang et al. from the same group
in 2009 [11, 12]. The ABF consisted of a helical ‘tail’ made from semiconductors and a square
‘head’ made from magnetic materials nickel (Ni) for magnetic actuation (Figure 1.3). The total
length of these artificial bacterial flagella (ABFs) ranged from 30 m to 100 m. Zhang et al.
showed that the ABFs could be wirelessly powered and steered in liquid using low-strength
rotating magnetic fields (< 10 mT and < 100 Hz). By changing the conditions of the magnetic
fields, ABFs were able to be wirelessly controlled and swim in three dimensions (3D) with
micro-scale precision. The propulsion modelling of ABFs was explained using resistive force
theory by Abbott and Peyer et al. [13, 14] (See Section 2.4.2). In 2010, a review paper
“Artificial bacterial flagella for micromanipulation” by Zhang et al. summarized the
4
fabrication, magnetic actuation and control of the ABFs and showed a clear roadmap of the
current progress and future directions of ABFs for biomedical applications, such as
sensing/marking and targeted therapies (Figure 1.4).
Figure 1.3. The ABF made by self-scrolling technology. (a-f) The fabrication flow of the ABF. (g) A scanning
electron microscope (SEM) image of an untethered ABF. The scale bar is 4 m. Reused from [11].
ABFs have the ability to navigate in liquid using low-strength rotating magnetic fields. Since
low-frequency and low-strength magnetic fields are harmless to living cells and tissues in the
human body and can penetrate through a human body allowing remote control of microrobots,
these magnetic powered flagella-like devices are proposed to be one of the promising tools for
in vitro and in vivo biomedical applications [6, 14]. Initially the ABFs were made from
semiconductor materials (InGaAs) which are expensive and toxic to cells and tissues, hence,
hindering the biomedical application of the devices. It, therefore, became necessary to devise
a new method to fabricate lower cost ABFs with better biocompatibility.
1.2.2 Contributions of the dissertation
In this dissertation, a straightforward method to make ABFs from polymers is introduced. In
addition to fabrication, motion control by magnetic fields, cytotoxicity and the
functionalization of ABFs for potential biomedical applications are studied. The structure of
the dissertation is as follows.
5
Figure 1.4. A roadmap for magnetic helical microrobots, also called ABFs, in biomedical applications. Reused
from [6].
Chapter 2 introduces the actuation methods of micro and nano robots presented in the literature
with a special focus on magnetic actuation. Magnetic actuation is wireless and shows no harm
to cells and tissues, hence, it is a reasonable choice for biomedical applications. Following this,
one type of microrobot, magnetic helical microrobots which mimic the propulsion method of
bacteria E. Coli, is discussed in detail, and the modelling of the helical propulsion method is
explained.
In chapter 3, the fabrication of magnetic helical microrobots is presented in detail. Four main
methods for making magnetic helical microrobots are introduced and one of the methods, laser
direct writing (DLW, a kind of 3D laser lithography) is focused on in more detail. The
fabrication process is explained, followed by the motion control and the swimming behavior
of ABFs in liquid is presented in chapter 4. ABFs show the ability of 3D navigation in liquid
and the frequency-dependent swimming behaviors are explained.
For biomedical applications, the biocompatibility and the functionalization of ABFs should be
6
pursued. The biocompatibility of ABFs, especially the in vitro cytotoxicity, is presented in
chapter 5. In chapter 6, the functionalization of ABFs with lipid-based drug carriers is studied
and the localized drug delivery in vitro using these functionalized ABFs (f-ABFs) with model
drugs and DNA is demonstrated. For in vivo applications, the ABFs have to be tracked and
steered. The in vivo tracking and actuation of ABFs in a mouse peritoneal cavity by
combination of a near-infrared tracking system and a magnetic actuation system is shown.
Chapter 7 contains a short summary of the findings in this work and some opinions about
further directions are discussed.
7
2 Propulsion mechanisms of microrobots
This chapter provides a review of the actuation methods commonly used to power mobile
microrobots. First, the definition of low Reynolds number is introduced, since the most of
microrobots are moving in a fluid at a low Reynolds number. Second, actuation methods are
reviewed using two main categories: chemical means and physical means. Third, the magnetic
actuation, a subcategory of physical means, is reviewed in more detail. Finally, magnetic
helical microrobots, the microrobots used in this dissertation, are discussed.
2.1 Swimming at low Reynolds number
As previously mentioned, most potential applications for mobile microrobots are in biological
and medicinal fields, and in most situations the microrobots have to move in liquid. In order to
understand how they move, it is important to study the environment, which is the liquid they
are swimming in. In fluid mechanics, the term Reynolds (Re) number, is commonly used to
characterize the conditions of flow in a fluid. Re is a dimensionless quantity defining the ratio
of inertial forces to viscous forces when an object moves in a fluid (Equation (2-1)).
𝑅𝑒 =𝑢𝐿𝜌
𝜂~
𝐹𝑖𝑛𝑒𝑟𝑡𝑖𝑎𝑙
𝐹𝑣𝑖𝑠𝑐𝑜𝑢𝑠 (2-1)
where 𝑢 and 𝐿 are the speed of motion and the characteristic length of the object, respectively,
while ρ and η are the density and the viscosity of the flow, respectively. Generally, flows can
be divided into three types by the Re number; laminar, transitional and turbulent flows.
Laminar flow has a Re number less than 2000, turbulent flow has a Re number higher than
4000, and transitional flow has a Re number between 2000 and 4000.
At low Re, we are in a world that is very viscous, very slow, or very small [14]. Mobile
microrobots, like most microorganisms, swim in a low Re regime on the order of 10-4. The
flow around a body at low Re is laminar. At low Re, the flow is effectively reversible and
consequently, reciprocal motion, i.e., body motion that simply goes back and forth between
two configurations, results in negligible net movement. In 1977, the paper “Life at low
Reynolds number” from Purcell pointed out that a non-reciprocal motion is required for a net
displacement in low Re number environments and proposed his “scallop theorem” [15]. This
theorem can be understood with a theoretical 3-link swimmer (Figure 2.1). The two hinges on
the structure offer two degrees of freedom (DOF) and the structure, therefore, can move in a
8
series of angle configurations. In Figure 2.1a a net displacement can be generated when the
swimmer moves in the series of configurations ABCDA after one cycle as this is a non-
reciprocal motion. In Figure 2.1b, however, the series of configurations ABCBA is reciprocal
and there is no net displacement after one cycle [6]. According to Purcell’s paper, two
techniques that microorganisms use to swim in low Re number generate nonreciprocal motion:
the “flexible oar” waving an elastic arm and the “corkscrew” rotating a chiral arm (Figure 2.2).
Purcell’s “scallop theorem” tells us the basic requirements for designing micro/nanoscale
swimmers, i.e., swimmers must move in a non-reciprocal motion to achieve a net displacement.
Figure 2.1. The two-hinged swimmer presented by Purcell. (a) A net displacement can be generated when the
swimmer moves in the series of configurations ABCDA after one cycle. (b) The series of configurations ABCBA
is reciprocal and there is no net displacement after one cycle. Reused from [6].
Figure 2.2. The two proposed swimming strategies used by microorganisms swimming in low Re number
environments. Reused from [15].
2.2 Propulsion mechanisms
During the last century, various tiny swimming devices in micro and nanoscale were invented
9
due to the improvement of micro/nanofabrication technologies. The propulsion mechanisms of
those swimmers can generally be divided into two categories, propulsion by chemical means
and propulsion by physical means. A review paper from Wang et al. summarized some of the
actuation methods (Figure 2.3) [16].
Figure 2.3. Mobile microrobots or micro motors propelled by different mechanisms. Adapted from [16].
2.2.1 Chemical means
Chemically powered mobile micro/nanorobots are based on the power generated by chemical
reactions between the surface materials of robots and a solution, usually hydrogen peroxide
(H2O2). During the chemical reaction, oxygen is produced at the surface of the materials
according to the reaction, 2𝐻2𝑂2 (𝑙) = 𝑂2(𝑔) + 2𝐻2𝑂 (𝑙), and provides the driving force for
propulsion. Based on this reaction, two different propulsion mechanisms have been proposed.
The first one is self-electrophoretic propulsion by which Ni/Pt bi-segment metallic nanorods
were able to move in liquid due to the electrons produced by the chemical reaction, flowing
from one side of the rods to the other (Figure 2.3a) [17-19]. The second one is bubble
propulsion, by which Janus microspheres [20] and multilayer microtubes were propelled by the
oxygen generated during the reaction (Figure 2.3b) [21, 22]. The drawback of these propulsive
methods is the need of the toxic fuel solution H2O2, which is infeasible for biomedical
10
applications.
In nature, motile microorganisms swim in liquid by moving their flagella or cilia in a non-
reciprocal way (Figure 2.4) [14]. In this way, microorganisms are able to translate chemical
power to mechanical power. Researchers have harvested the chemical power from organisms
to actuate micro-devices by attaching microorganisms with devices, called bio-hybrid
microrobots. M. S. Sakar et al. integrated swarmer cells of bacteria Serratica marcescens with
SU-8 U-shape microstructures [23]. V. Magdanz et al. presented a micro-bio-robot which
consists of a living sperm cell and a magnetic microtube. The sperm cell moved based on the
flagellar propulsion of the cell [5]. B. J. Williams et al. recently built a microswimmer which
consisted of a polymer filament and cardiomyocytes. The cardiomyocyte cells contracted and
deformed the filament to generate propulsive force [24].
Figure 2.4. Three propulsion mechanisms of microorganisms. (a) Cilia hold perpendicular to the flow during the
power stroke and fold near the cell body during the recovery stroke. (b) Beating flexible flagella in a planar wave.
(c) Rotating the flagella in a helical pattern. Reused from [14].
2.2.2 Physical means
The second main category of propelling micro/nanorobots is physical means. Different external
power sources have been used to actuate tiny devices in liquid, including electric, optic,
acoustic and magnetic fields.
Electric field: Under an electric field E, a particle with a charge q experiences an electric force
FE = q E and moves in parallel with the direction of electric field E. The positively charged
11
particle moves along the direction of E, while the negatively charged particle moves in the
opposite direction. This phenomenon is called electrophoresis. Electrophoresis has been used
to move charged micro/nano-objects in liquid or soft materials. For example, gel
electrophoresis is a common technology in life science to separate DNA sequences or proteins
inside a gel solution. Another important phenomenon when an electric field is applied in liquid
is electroosmosis. In electroosmosis, a bulk of liquid moves in response to the electric fields
[25] and propels objects. Chang et al. published a paper in Nature Materials in 2007 showing
that microdiodes and microparticles were actuated by external electric fields based on
electroosmotic effects (Figure 2.5a) [26]. The diode motility results from a local electroosmotic
flux, powered by the external field (less than 120 V·cm-1). The microdevices moved on the
surface of water and the speed of millimeters per second was achieved (Figure 2.5b). The
direction of the movement depended on the orientation of the anode and cathode. Two diodes
with different orientations of anodes and cathodes moved in opposite directions under the same
electric field (Figure 2.5c) [26]. Hwang et al. used the same mechanism to power mm-sized
helical swimmers in a viscous solution [27, 28].
Optical tweezers: The principle of optical tweezers is to use forces generated by a highly
focused beam of light, such as a laser beam, to trap and manipulate small objects [29]. The
optical gradient forces in the axis of the light can trap small objects in 3D near the focal point
of the light (Figure 2.6a). They can independently trap and move objects ranging from tens of
micrometers to tens of nanometers in size with computer-programed trapping patterns. Figure
2.6b shows 36 polystyrene spherical beads of 800 nm in diameter trapped in a plane with two
different computer-programmed patterns. Optical tweezers have been used for in vitro
biomedical applications such as probing the viscoelastic properties of DNA, cell membranes
and protein fibers and manipulating living cells. The optical traps in optical tweezers are
typically very close to the microscope objective and require a strong laser beam, which hinders
in vivo applications of optical tweezers due to the distance between the optical trap and targeted
samples in vivo and the harm of a strong laser beam to cells and tissues.
12
Figure 2.5. Miniature semiconductor diodes powered by an alternating electric field. (a) The actuation principle
of the floating diode and the schematic of the experimental setup for measurement of the moving speed. (b) The
overlay of images showing the movement of a microdiode on the surface of water. (c) Two diodes moved towards
the top or bottom depending on the orientation of their anodes under the same electric field. Adapted from [26].
Figure 2.6. Optical tweezers trap and manipulate small objects. (a) The principle of optical tweezers. The gradient
force generated by a laser beam traps a colloidal particle in 3D near the focal point. (b) Polystyrene spherical
beads with 800 nm in diameter were trapped in a plane two different computer-programmed trapping patterns.
Adapted from [29].
13
Ultrasound power: Ultrasound is oscillating sound pressure waves with frequencies above the
threshold value of the human hearing range. The ultrasound waves deliver energy to the objects
in the range of the waves. Ultrasound has been used for many applications such as detecting
objects, measuring distances, cleaning and mixing, and manipulating micro/nanoobjects in
liquid. Ultrasound propulsion has been used to power micro/nanorobots (Figure 2.3c). Mallouk
is one of the pioneers who used continuous or pulsed ultrasound to power micro and nanoscale
metallic rods [30, 31]. The ultrasound was produced by a generator mounted on the bottom of
the experiment cell (Figure 2.7a). Ultrasonic waves in the MHz frequency have been shown to
rotate, align, assemble and propel metallic microrods (330 nm in diameter and 2 μm in length)
in water and in solutions of high ionic strength. The fast axial motion of the microrods at ∼200
μm/s was achieved using continuous or pulsed ultrasound [31]. The shape asymmetry of the
microrods as a result of template electrodeposition was proposed to explain the axial propulsion
of the rods. The shape asymmetry of the rods induced non-equivalent distributions of the
ultrasound pressure (One side of the rod has a high acoustic pressure and the other side has a
low acoustic pressure (Figure 2.7b).) Therefore, a pressure gradient was generated to propel
the rods directionally. Recently, the authors showed navigation of these microrods inside living
Hella cells by acoustic propulsion [32]. Since ultrasound is a sound wave that can penetrate
living bodies and is biocompatible to cells and tissues, ultrasound has potential to power
micro/nanorobots for in vivo applications.
Figure 2.7. Nanorobots powered by ultrasound. (a) A schematic of the experimental setup for studying the
nanorobots powered by ultrasound. Adapted from [31]. (b) The RuAu hybrid nanorods were propelled by
ultrasound propulsion. Adapted from [33].
Magnetic fields: Magnetic fields are generated from electric currents and magnetic materials.
A magnetic field at any position in space is defined as a vector field including a direction and
14
a magnitude/strength. Powering micro/nanorobots using magnetic fields has gained lots of
interests due to its biocompatibility with tissue and the good penetration through human tissue,
which enable wirelessly powering micro/nanorobots at a long distance (Figure 2.3d). The
magnetic actuation of micro/nanorobots will be introduced in detail in the following section.
2.3 Magnetic actuation methods
2.3.1 The principle of magnetic actuation
The basic principle of magnetic actuation is to move a magnetic object by applying a magnetic
force and/or magnetic torque onto it. When an external magnetic field is applied to a magnetic
object, the magnetic force FM [N] and the magnetic torque TM [N·m] acting on the body is
given by the following two equations:
𝐹𝑀 = 𝜇0𝜗(𝑀 ∙ ∇)𝐻 (2-2)
𝑇𝑀 = 𝜇0𝜗𝑀 × 𝐻 (2-3)
where 𝜇0 is the permeability of free space (𝜇0 = 4π × 10-7 T·m·A-1, T represents the unit Tesla),
ϑ [m3] is the magnetic volume of the object, M [A·m-1] is the magnetization of the object, H
[A·m-1] is the external magnetic field and the magnetic field also can be expressed as the
magnetic flux density/magnetic induction B [T] ( 𝐵 = 𝜇0𝐻), and 𝛻 [A·m-2 or T·m-1] is the
gradient of the magnetic field.
From Equation (2-2), we can see that the magnetic force is governed by the parameters ϑ, M,
𝛻, H and the angle between H and M. From Equation (2-3), the magnetic torque is governed
by the parameters ϑ, M, H and the angle between H and M. So there are five parameters (ϑ, M,
𝛻, H and the angle between H and M) in total which govern the magnetic force FM and torque
TM of a magnetic body. Among these five parameters, only ϑ depends on the volume of
magnetic materials or the size of robots. H and 𝛻 depend only on the external magnetic fields.
The angle between M and H depends on the shape of robots. M is related to the magnetic
properties of magnetic materials which is a function of H.
2.3.2 Magnetic materials
Magnetic materials can be classified into three main types according to the response of
materials in an external magnetic field: ferromagnetic materials, paramagnetic materials and
15
diamagnetic materials. Simply speaking, ferromagnetic materials are strongly attracted to a
magnet, and examples of these materials are iron (Fe), cobalt (Co), nickel (Ni), some rare earth
metals and their alloys. Paramagnetic materials are slightly attracted to a magnet and examples
of these materials are platinum (Pt), aluminum (Al) and manganese (Mn). Diamagnetic
materials are weakly repelled by a magnet and examples of these materials are silver (Ag),
copper (Cu), gold (Au), beryllium (Be) and bismuth (Bi) [34].
Ferromagnetic materials are mostly used for building magnetic micro/nanorobots since they
have much higher M than the other two types of material in the same external magnetic field
H. Therefore, they experience more magnetic forces and torque under the same conditions (the
same ϑ, H, 𝛻 and the angle between H and M) than the other two material categories experience
(Equation (2-2) and (2-3)).
Ferromagnetic materials have been used in many different applications such as in permanent
magnets, magnetic recording, electrical motors and power generators. They are generally
divided into two broad classifications based on their coercivity: hard magnetic materials and
soft magnetic materials. Coercivity Hc is the intensity of the magnetic field applied which is
needed to drive the magnetization M of a material to zero after it has previously reached
magnetic saturation point (which is shown in the hysteresis loops of hard and soft magnetic
materials in Figure 2.8). Generally, materials with Hc above 10,000 A·m-1 are hard magnetic
materials while soft magnetic materials have Hc below 1,000 A·m-1 [34]. Once hard magnetic
materials are magnetized and saturated by a strong external field H, they will keep M when
external magnetic fields are moved away. That is why hard magnetic materials are also called
permanent magnets. For hard magnetic materials M shows independently on the applied field
H below a certain value (Figure 2.8a), while for soft magnetic materials M is not constant, but
closely depends on the field H before saturation (Figure 2.8b).
Hard magnetic materials have been used to build robots at the millimeter-scale and larger. For
example, a spiral-type cm-sized robot used a samarium-cobalt (SmCo) magnet
(1mm×1mm×1mm) for magnetic actuation [35], and a cylindrical neodymium-iron-boron
(NdFeB) magnet (500 m in diameter, 600 m in length) was used to power magnetic robots
with the total length around 1mm [36]. It is a challenge to make microscale-sized hard magnetic
materials with arbitrary shapes. On the other hand, soft magnetic materials (such as Fe, Ni, Co)
are commonly used to build smaller magnetic robots in micro and nanoscale since they are
flexible in size and shape.
16
Figure 2.8. The typical hysteresis loops of hard and soft magnetic materials.
For soft magnetic materials, the magnetization M of the body is a nonlinear function of the
magnetic field H. Abbott et al. developed a model for magnetic torque and force on soft
magnetic objects with axial symmetry, for instance, an ellipsoid [37]. The hysteresis loop of
soft magnetic materials can be characterized by three regions: linear-magnetization, nonlinear
and saturation regions (Figure 2.8b). In the linear-magnetization region when the H is relatively
low, the M is related to the applied field 𝐻 by an apparent susceptibility tensor 𝜒𝑎 (Equation
(2-4)). The susceptibility tensor 𝜒𝑎 is a function of the susceptibility of the material 𝜒 and a
demagnetization factor 𝑁 [37].
𝑀 = 𝜒[1 + 𝜒𝑁]−1 = 𝜒𝑎𝐻 (2-4)
The demagnetizing factor 𝑁 depends on the shape of the magnetic bodies and varies in the
different directions within the magnetic body, which is known as shape anisotropy. The
demagnetizing factors are largest along the short axis of a body. So an easy magnetization axis
refers to the long axis of a body, since it is a relatively easy direction to magnetize the materials
[14]. For instance, a magnetic body with a large aspect ratio, such as a compass needle, has an
easy axis in the direction of the length of the needle.
In addition to ferromagnetic materials, paramagnetic materials such as iron oxide micro and
nanoparticles have been used for building magnetic micro/nanorobots. A microrobot consisting
of a red blood cell as a “head” and a chain of iron oxide microparticles linked by DNA as a
“flexible tail” was self-propelled by planar beating of the flexible tail [38]. Superparamagnetic
iron oxide nanoparticles were embedded into a helical-shaped polymer to create magnetic
helical microrobots [39, 40].
17
2.3.3 Magnetic fields
A magnetic field can be generated by an electric current or a permanent magnet, hence,
electromagnets or permanent magnets have been used to generate different types of controlled
magnetic fields, such as rotating magnetic fields, oscillating magnetic fields and magnetic
fields with gradients (Figure 2.9), for manipulating magnetic micro/nanorobots. A summary of
these different magnetic fields was presented in a review paper from Peyer et al. [41].
Figure 2.9. Magnetic fields for actuation of micro/nanorobots. (a) Rotating fields with the field vector B rotated
in a plane, called normal fields. (b) Rotating fields with B rotated along the mantel of a cone, called precession
fields. (c) Oscillating field. (d) On-off field (e) Field gradients with the gradient along the direction of B. (f) Field
gradients with the gradient perpendicular to the direction of B. Reused from [41].
Rotating magnetic fields: In rotating magnetic fields, the field vector B rotates around an axis
continuously. The normal rotating field has the field vector B rotating in a plane which is
perpendicular to the rotational axis (Figure 2.9a). Another type of rotating fields is a precession
field, in which, the field vector B rotates along the mantle of a cone (Figure 2.9b).
Helical microrobots are actuated well in rotating magnetic fields. The microrobots rotate
around their helical axis by following the rotating fields and move in a direction parallel to the
helical axis (Figure 2.10a) [11]. Hybrid nanorods with Ni and Au segments connected by
flexible Ag were actuated by rotating magnetic fields. Under a normal rotating field, the Ni
segment started to rotate and induced the rotation of the Au segment, which broke the symmetry
of the system and induced a forward movement (Figure 2.10b) [42]. By changing the direction
of the rotational axis, these two types of swimmers can move in 3D in liquid without a surface
support. Magnetic micro/nanostructures with different shapes can be actuated and move on a
surface under rotating fields. Ni nanowires move parallel to a surface plane by rotating or
18
tumbling on a surface (Figure 2.10c) [43], as well as the self-assembled chain of
superparamagnetic beads (Figure 2.10d) [44] and permanently magnetized sphere (Figure
2.10e) [45]. These structures are also called “surface walkers” due to the necessity of a surface
to support their movement. A precession rotating magnetic field has been used recently to
actuate a microrobot with a rigid helical “tail” [46]. The results showed that the helical
microrobot has better stability and swimming performance in a low rotating frequency range
under a precession field than a normal field (Figure 2.10f).
Figure 2.10. Micro/nanodevices actuated by rotating magnetic fields. (a-e) Micro/nanostructures actuated in
normal rotating fields. (a) Helical-shaped microrobot. Adapted from [47]. (b) A flexible nanorod with Ni/Ag/Au
segments. Adapted from [42]. (c) Ni nanowire. Adapted from [43]. (d) The self-assembled chain of
superparamagnetic beads. Adapted from [44]. (e) Permanently magnetized sphere. Adapted from [45]. (f)
Microrobot with a helical “tail” in a procession magnetic field. Adapted from [46].
Oscillating and on-off magnetic fields: In an oscillating field, the field vector B moves up
and down in a plane (Figure 2.9c). An artificial microswimmer with a flexible tail was actuated
under this field [38]. The microrobot had a red blood cell which served as the “head”, and a
linear chain of magnetic particles linked by DNA serving as the flexible “tail” (Figure 2.11a).
The flexible tail aligned with the field and moved in a beating pattern that propelled the
19
structure under an oscillating field (Figure 2.11b). Using oscillating fields, microrobots can
“walk” on a surface by a stick-slip motion (Figure 2.11c). The V-shaped microrobot was made
from NdFeB fabricated by laser micromachining and achieved the speed of 2.8 mm/s, 11 body
lengths per second (Figure 2.11d) [48].
Figure 2.11. Microrobots actuated by oscillating and on-off magnetic fields. (a) A microrobot with a red blood
“head” and a flexible magnetic “tail”. Adapted from [38]. (b) Beating pattern of the flexible-tail microrobot
(shown in (a)) under oscillating fields. The arrows on swimmers (a-t) represent the directions of oscillating fields.
Adapted from [38]. (c) Stick-slip motion of a microrobot under oscillation fields. The upper images are real-time
images observed by a high speed camera, and the lower images are analogous steps from simulated results.
Adapted from [48]. (d) An SEM image of V-shaped microrobot. Adapted from [48]. (e) A SEM image of
MagMites. Adapted from [49]. (f) The schematic of PolyMites. Adapted from [50]. (g) Mechanical model of
PolyMites and MagMites. Adapted from [50]. (h) The real-time impact-motion of MagMites captured by a high
speed camera under on-off magnetic fields. Adapted from [49].
In an on-off magnetic field, the magnetic field is applied in an on-off cycle (Figure 2.9d).
Microrobots, such as MagMites (Figure 2.11e) and PolyMites (Figure 2.11f) developed in the
Nelson group [49, 50], were actuated under this field. Generally, they consist of two soft-
magnetic parts (called the “body” and “hammer”) connected by a “spring” (Figure 2.11g). The
20
two parts separate and the spring is in a released state when the magnetic field is off. When the
magnetic field is on, the two soft-magnetic parts are magnetized and attract each other, resulting
in a deformation of the spring. When the magnetic field turns off, the magnetic force between
two parts dissipates and the energy stored in the spring pushes the two bodies apart. The on-
off cycle of the magnetic fields was in the range of kHz near the Eigen frequency of the devices.
In this manner, they transformed the magnetic energy into inertia and impact-driven
mechanical force (Figure 2.11h). They can be actuated on a surface both in wet and dry
environments. The robustness of the MagMites successfully competed in the RoboCup 2007
and 2009 Nanogram competitions [49].
Figure 2.12. Micro/nanostructures actuated by magnetic fields with gradients. (a-b) An elliptical magnetic device
moved in 3D in a field with gradients generated by the OctoMag. Adapted from [51]. (c-f) A single Ni nanowire
moved in 3D under a highly controlled field with gradients generated by the NanoMag. The direction of B and
gradient 𝜵 is parallel in (c-d) while perpendicular to each other in (e-f). Adapted from [52].
Magnetic fields with gradients: In a gradient field, magnetic micro/nanorobots experience
both a magnetic torque and a magnetic force. The torque given by Equation (2-3) aligns the
robots until the M of the robot is parallel to B, and the force given by Equation (2-2) moves the
robots towards the direction where magnetic fields increase. The direction of B can be either
parallel to the direction of gradient 𝜵 (Figure 2.9e) or perpendicular to the direction of 𝜵
(Figure 2.9f). In a gradient field, magnetic micro/nanorobots of any shape can be pulled with a
magnetic force as long as the force is big enough. A mm-sized elliptical robot was actuated by
an electromagnetic actuation (EMA) system, called OctoMag built by the Nelson group at ETH
Zurich, in 3D with high controllability (Figure 2.12a-b) [51]. And a nanoscale Ni nanowire was
21
actuated by another EMA system, called NanoMag built by the same group, in 2D and 3D
space (Figure 2.12c-f) [52].
2.4 Magnetic helical micro/nanorobots
As mentioned before, magnetic helical micro/nanorobots mimic the flagellar propulsion
method of bacteria such as E. coli. They have the ability to perform 3D navigation in low-Re-
number environments with micrometer precision under low-strength rotating magnetic fields
(less than 10 mT). In contrast to surface walkers, magnetic helical micro/nanorobots do not
need any support surface and do not need a field gradient, which is advantageous as generating
enough gradient to pull micro/nanoscale magnetic objects over a long distance is a limitation
of the usage of gradient fields. The magnetic helical micro/nanorobots combine the advantages
of both magnetic actuation and helical propulsion. It has been proposed as one of the most
promising propulsive methods for biomedical applications, especially for in vivo applications
[14, 53]. In the remaining section of this chapter, an overview of magnetic helical
micro/nanorobots and the modelling of the helical propulsion are presented.
2.4.1 An overview of magnetic helical micro/nanorobots
In 1996, Honda and his co-workers proposed the helical-type swimming mechanism for
microrobots in low Re number and fabricated the first prototype of magnetic helical
microrobots in cm-size. The swimmer consisted of a cubic SmCo magnet (1mm×1mm×1mm)
attached to a Cu helical wire (Figure 2.13a). The cm-sized model was wirelessly actuated by
an external rotating magnetic field and proved to be able to swim in low-Re-number fluid by
swimming in a highly viscous silicon oil [35]. In 2005, the same group developed a similar
helical robot with a smaller size. The total length of the robot was reduced to 5.55 mm (Figure
2.13b). The authors showed that the mm-sized helical robot was able to trail a wire and change
the motion direction of the wire in a narrow fluidic channel. They proposed that the helical
microrobots have a great potential for navigating medical catheters in blood vessels [54].
22
Figure 2.13. An overview of magnetic helical micro/nanorobots. (a) The first prototype of magnetic helical robot
in cm-size in 1996. Adapted from [35]. (b) Guidance of wires in a narrow channel by a magnetic helical microrobot
in 2005. Adapted from [54]. (c) The first microscale prototype of helical microrobots made by self-scrolling
technology in 2007. Adapted from [10]. (d) The first nanoscale helical propeller made by glancing angle
deposition in 2009. Adapted from [55]. (e) Open-loop velocity control with gravity compensation. Adapted from
[56]. (f) Magnetic helical microrobots made by 3D laser lithography. Adapted from [57]. (g) Helical
microstructures derived from spiral vessels of different plants. Adapted from [58]. (g) Helical nanoswimmer made
by a template electro-synthesis method. Adapted from [59].
In 2007, the first microscale prototype of magnetic helical microrobots, artificial bacterial
flagella (ABFs), were invented by a self-scrolling technique in the Nelson group at ETH
Zurich. The microswimmer has a soft magnetic ‘head’ and a helical ‘tail’ with the diameter of
3 m and the length of 30-40 m (Figure 2.13c) [10]. The detailed fabrication process is shown
in Section 3.1.1. The magnetic actuation and swimming behaviors of ABFs were then
23
characterized by Zhang and Peyer et al. from the same group [6, 11, 12, 60]. In 2012, a new
fabrication process, 3D laser lithography, was used to fabricate ABFs (Figure 2.13f). Details
of this fabrication method are in Section 3.1.3. With this new fabrication method,
microstructures with almost arbitrary shapes can be fabricated, for example, a helical swimmer
with a claw for 3D transportation of micro-beads [57] and helical microswimmers with
mastigoneme-inspired appendages [61]. The swimming performance of ABFs in
heterogeneous viscous environments was studied by Peyer and her co-workers [62]. The further
functionalization and biomedical applications of ABFs were demonstrated in these publications
[63-68] and also included in Chapter 5 and 6.
In 2009, the first nanoscale magnetic helical robot was fabricated using glancing angle
deposition by Ghosh and Fischer (Figure 2.13d). The detailed fabrication is shown in Section
3.1.2. The nano-propellers could navigate in water under rotating magnetic fields [55].
Recently, these nanopropellers showed the possibility of actuation in viscoelastic media [69]
and in human blood [70].
In 2010, Thomas et al. showed the control of scaled-up magnetic helical microrobots using a
non-uniform magnetic field generated by a rotating permanent magnetic manipulator [71].
Further, Mahoney et al. from the same group presented an algorithm enabling velocity control
with gravity compensation of scaled-up magnetic helical microrobots (Figure 2.13e) [56].
In 2013, by harnessing the existing biological structures of nature, Gao et al. fabricated
magnetic helical microstructures derived from spiral vessels of different plants (Figure 2.13g).
The detailed fabrication is shown in Section 3.1.4. This fabrication process was simple and
cost-effective for the mass-production of helical microswimmers [58]. In 2014, a large-scale
fabrication of magnetic helical nanoswimmers by a template electro-synthesis method was
presented by the same group (Figure 2.13h) [59].
2.4.2 Modelling of helical propulsion
Modelling the helical propulsion of ABFs has already been explained by previous researchers
using resistive force theory (RFT) [12, 14, 60]. The motion of helical propulsion can be
approximated by a 1D propulsion matrix [15]. At low Re number (Re << 1), the inertial force
is negligible and a linear relationship between the force, the velocity and rotational speed is
expected, and the propulsion matrix can be captured in the following matrix,
24
(𝐹
𝑇) = (
𝑎 𝑏𝑏 𝑐
) (𝑢
𝑓) (2-5)
where F and T are the magnetic force and the torque acting on the ABF, respectively. The
vector of u and f represents the velocity and the rotational frequency of the ABF, respectively
(Figure 2.14). The coefficients a, b and c are scalars and a function of geometric parameters of
the ABF and the viscosity of the fluid.
Figure 2.14. Schematic of an ABF. A magnetic field Bm rotates around the helical axis under a frequency f. The
torque T rotates the robot and the helical shape of the ABF transforms the rotational motion into a linear velocity
u. Reused from [40].
For a helix with a helicity angle θ, these coefficients a, b and c are given as,
𝑎 = 2𝜋𝑛𝐷 (
𝜉ǁ cos2 θ + 𝜉⊥ sin2 θ
sin θ) (2-6)
𝑏 = 2𝜋𝑛𝐷2(𝜉ǁ − 𝜉⊥) cos θ (2-7)
𝑐 = 2𝜋𝑛𝐷3 (
𝜉⊥ cos2 θ + 𝜉ǁ sin2 θ
sin θ) (2-8)
where 𝑛 is the number of the helix turns, D is the helix diameter, r is the filament radius, and
𝜉⊥ and 𝜉ǁ are the viscous drag coefficients found by RFT parallel and perpendicular to the
filament, respectively, given by [14, 72].
𝜉⊥ =4𝜋𝜂
ln(0.18𝜋𝐷𝑟 sin θ
) + 0.5 (2-9)
𝜉ǁ =2𝜋𝜂
ln(0.18𝜋𝐷𝑟 sin θ
) (2-10)
Assuming that there is no external force in the direction of the helical axis, which is reasonable
when the microrobot swims horizontally, the first line of Equation (2-5) can be written as,
25
𝑢 = −
𝑏
𝑎 𝑓 (2-11)
The equation shows the linear relationship between 𝑓 and 𝑢, indicating a frequency-dependent
behaviour of the ABF. The further discussion of ABF swimming behaviors is in Section 4.3.
2.5 Conclusion
Magnetic helical microrobots, mimicking the flagellar motion of bacteria E. Coli, can be
powered and steered in 3D in liquid under low-strength rotating magnetic fields. The flagellar
propulsion method has been proposed as one of the most promising means for biomedical
applications, especially for in vivo application. The fabrication of these helical robots is
fundamentally important for their further applications. The next chapter introduces the detailed
fabrication of magnetic helical microrobots.
26
3 Fabrication of ABFs
This chapter focuses on the fabrication of magnetic helical microstructures. First, it starts with
an introduction of various fabrication methods for magnetic helical micro/nano structures.
Next, one of the methods, direct laser writing (DLW), a type of 3D laser lithography, is
explained in detail including its basic principle, fabrication process and typical examples. With
DLW, almost any arbitrary 3D microstructure can be fabricated. The magnetic helical
microrobot studied in this dissertation, the artificial bacteria flagellum (ABF), was fabricated
using DLW. Some results in this chapter have been published in two journal papers [53, 57].
3.1 Introduction
A magnetic helical microstructure consists of at least two components, a helical body and a
magnetic material. The helical body mimics the helical propulsion motion of bacterial flagella
and provides the structure the ability to perform translational movement when it rotates along
the helical axis. The magnetic material enables the structure to rotate by following external
rotating magnetic fields. The helical body is a 3D microscopic structure which limits the choice
of fabrication methods. In this introduction, we summarize the fabrication methods of magnetic
helical microstructures into four categories, which are rolled-up (Figure 3.1a), glancing angle
deposition (GLAD) (Figure 3.1b), direct laser writing (DLW) (Figure 3.1c) and template-
assisted methods (Figure 3.1d) [53, 73].
3.1.1 Rolled-up method
In 2007, the first truly microscale magnetic helical microrobot, the artificial bacterial flagellum
(ABF), was fabricated by the rolled-up method, also known as self-scrolling technology [10].
The ABF had the total length of 30 m to 40 m and consisted of a helical ‘tail’ made from
semiconductor materials (InGaAs) and a magnetic Ni ‘head’. The fabrication was based on
traditional thin film deposition methods and mono-crystalline thin film growth. By depositing
or growing bi- or tri-layers of material, internal stresses in the structure result in a bending of
the thin films [74, 75]. In this manner straight ribbon patterns can form helices in a controllable
fashion (Figure 3.1a) [76]. By controlling the deposition parameters, such as the film thickness,
the ribbon width, or the orientation of the ribbon with respect to the crystalline structure of the
metal, the curvature can be finely tuned. By connecting these self-scrolled helical structures
27
with a Ni plate, ABFs with diameters of around 3 µm and variable lengths of 10 to 100 µm
were achieved. Like the SEM image shown in Figure 3.1e, the ABFs have a non-magnetic
helical tail with the magnetic material localized only in the head design. The swimming and
actuation have been explored in detail in further publications in 2009 by Zhang et al [11, 12].
Soft materials, such as lipid bilayers, also rolled up to helical or tubular microstructures.
Magnetic helical lipid microstructures were fabricated by electroless plating of helical
microstructures with a magnetic material CoNiReP [77].
Figure 3.1. An overview of fabrication methods for magnetic helical micro/nanorobots. (a-d) Schematic
representations of fabrication methods for magnetic helices; (a) Rolled-up method. Reused from [53]. (b) Glancing
angle deposition method. Reused from [53]. (c) Direct laser writing method. Reused from [53]. (d) template-
assisted method. Adapted from [58]. (e-i) Fabricated structures; (e) an ABF made from rolled-up method. Adapted
from [12]. (f) The nano helical structures fabricated by GLAD. Adapted from [55]. (g-h) The ABFs fabricated by
DLW. Adapted from [57, 78]. The scale bar in (g) is 2 m. (i) the helical structures fabricated by template-assisted
method. Adapted from [58].
3.1.2 Glancing angle deposition method
In 2009 a publication from Ghosh and Fisher demonstrated the batch fabrication of helical
nanoswimmers by means of glancing angle deposition (GLAD) [55]. Spherical seeds were
densely packed on a substrate, and nano-pillars were deposited at an oblique angle in GLAD.
By continuous rotation of the substrate, the pillars grew into a helical shape (Figure 3.1b).
These helices had a diameter of 200 to 300 nm and a length of 1 to 2 µm (Figure 3.1f). Magnetic
material Co was then deposited on one side of the nanohelices along the whole body length
28
and was permanently magnetized in the radial direction of the helical structures.
3.1.3 Direct laser writing method
In 2012 a new method was developed to make ABFs from polymers by direct laser writing
(DLW) and e-beam deposition [57]. DLW is a kind of 3D laser lithography method and it
allows the creation of arbitrary 3D microstructures. A photosensitive resist was deposited on a
glass substrate which can be moved in 3D with a piezoelectric stage. Laser beams were focused
into the resist and two-photon polymerization (TPP) occurs at the focal point of the laser. By
moving the focal point in a helical path, a helical microstructure remained after removal of the
undeveloped negative photoresist (Figure 3.1c). Next, magnetic materials were deposited by
electron beam evaporation on the entire helical structure. With this method, 3D microstructures
with arbitrary shapes can be fabricated. The ABFs used in this dissertation were fabricated by
this method. The detailed principle and fabrication process will be introduced in the following
Section 3.2. Other researchers have fabricated magnetic helical microstructures by mixing
magnetic nanoparticles inside the polymer instead of coating the surface of a polymer [39].
Recently, magnetic helical microstructures with hybrid materials were presented by Zeeshan
et al. in 2014 [78]. Hollow microstructures with a rectangular head contacting with a helical
body were fabricated by DLW in a positive-tone photoresist, and the hybrid microstructures
were obtained by subsequent deposition of CoNi magnetic materials and polymer poly(pyrrole)
(PPy) by electro-deposition (Figure 3.1h).
3.1.4 Template-assisted method
In 2013 Gao et al. showed a method to fabricate helical microswimmers using vascular plants
as helical templates [58]. Various vascular plants, from fruit-bearing trees to decorative shrubs,
have spiral xylem vessels. The diameters of these helical structures vary from 10 m to over
60 m. First, the authors isolated the helical microstructures from vascular plants by washing
the plant leaves in a low concentration KOH solution. After that, the helical microstructures
were coated with Ni and Ti layers by electron beam evaporation. Finally, the long
microstructures were cut into short helices of a few turns (Figure 3.1d-i). The geometries of
the magnetic helical swimmers can be controlled by the intrinsic structures of different plants.
It is a cost-effective method which allows for mass-production [58]. In 2014 the same group
presented another template-assisted method to mass-produce nano-sized magnetic helices [59].
They used porous AAO membranes with pore sizes of 100 to 400 nm as templates. Pd/Cu
29
nanorods were deposited into the nano-pores of the templates by electro-deposition. The Pd
helical nanoswimmers were fabricated by removal of Cu and the addition of an electron-beam
coating of a magnetic Ni layer.
3.2 Fabrication of microstructures using DLW
3.2.1 The principle of DLW
The principle of DLW is two-photon polymerization (TPP) [79]. The TPP technique is to use
a femtosecond laser to induce a highly localized chemical reaction, which leads to a local
polymerization of the photosensitive material with a resolution of approximately 100 nm [80].
TPP can fabricate 3D microstructures with arbitrary geometries at sub-micro resolution. Since
1998 when Kawata et al. established TPP as a technique to fabricate 3D microstructures [81],
researchers in various fields such as nanophotonics [82], microfluidics [83], and bioengineering
[84, 85], were able to create 3D microdevices to fit their purposes. The advantages of TPP
compared to conventional photolithography are: (1) Microstructures with arbitrary 3D
geometries can be fabricated. (2) A high resolution down to 100 nm can be achieved [86] [87].
The DWL tool used in this dissertation was from Nanoscribe GmbH in Germany. Figure 3.2
shows a schematic of the DWL tool. A femtosecond laser beam is generated and transported
through a microscope objective. The laser beam is focused onto a photoresist by the objective.
The photoresist at the laser focus point changes its properties by chemical reaction induced by
TPP. Related structures are created by moving the photoresist or the laser beam in a predefined
path. The schematic of writing a 3D helix using DLW is shown in Figure 3.3a. A helix is
fabricated by moving the laser focus point in a designed helical path. The different optical
properties between polymerized and unpolymerized resist enables the real-time visualization
of the writing process (Figure 3.3b). The focus laser point (the bright point in Figure 3.3b)
wrote helices in a commercial gel-like negative photoresist ORMOCOMP. Helixes with a
length of 16 m and a diameter of 5 m was written in 2.3 s [65].
30
Figure 3.2. Schematic of a direct laser writing tool. The original photo is from http://www.nanoscribe.de.
Figure 3.3. The principle of DLW based on TPP. (a) Schematic of writing a helix in a photoresist. The photoresist
present at the laser focal point polymerizes based on TPP. The 3D helical structure is written by moving the laser
focus point in a helical path. (b) The real-time writing process of helical structures. The bright point is the laser
focus point. From (1) to (3), a helix was finished, then the laser focus point moves to the next defined position to
write another helix (4). Reused from [65].
3.2.2 Fabrication process
This section introduces the fabrication of 3D microstructures using DLW. The general
fabrication flow is: (1) Design of 3D structures using CAD software or Matlab software (2)
31
Slicing the CAD 3D structures into 3D coordinates which the DLW tool can recognize, such
as a ‘.gwl’ file. (3) Prewriting to find the proper writing parameters. The two crucial writing
parameters are laser power and scan speed. The first parameter defines the power of the laser,
and the second defines the moving speed of the laser. (4) Writing microstructures using the
proper parameters.
Here, we take the fabrication of one helix to give a step-by-step demonstration of the process.
The geometries of a helix (the total length, the diameter and the pitch) are defined (Figure 3.4a)
in a code, normally in a “.gwl” file type, including 3D coordinates (x, y, z) of a helix generated
using a custom-made Matlab program (see Appendix A). Writing the helix using different
writing parameters, laser power and scan speed, is carried out, and the writing results are
characterized using scanning electronic microscope (SEM) (Figure 3.4b). The results show that
the structures written using laser power of less than 14% (100% laser power corresponds to 20
mW) with a scan speed of 25 m/s were mechanically too weak to stand up and collapsed onto
the glass substrate while the structures written using laser power less than 16% collapsed with
a scan speed of 50 m/s. The proper writing parameters can be 14% or 16% as laser power and
25 m/s as scan speed, or 16% with a scan speed of 50 m/s. The final step is to write arrays
of helices using the correct writing parameters.
Figure 3.4. General fabrication flow using DLW. (a) Designation of the structure using 3D CAD design software
or Matlab software. (b) Parameters test. Laser power and scan speed are two crucial parameters.
3.2.3 Fabrication of ABFs
As mentioned before, one single ABF has a helical body and contains a magnetic material.
After an array of polymeric helical bodies was successfully produced on the substrate by DLW,
32
the entire substrate was coated with magnetic materials, such as Ni or Fe, using electron beam
deposition with a tilting angle of 15 degrees to reduce shadowing effect (Figure 3.5a) [57].
Figure 3.5. Fabrication of ABFs. (a) The fabrication flow of ABFs. Adapted from [57]. (b) A horizontal array of
ABFs fabricated from IP-L photoresist and coated with Ni/Ti layers, the scale bars are 2 m and 10 m in the
inset and overview, respectively. Reused from [57].
The SEM image in Figure 3.5b shows the top-view array of ABFs, made from IP-L photoresist
(a commercial negative photoresist from Nanoscribe GmbH) and Ni/Ti coating, on the glass
substrate. The insert shows the side-view of four ABFs. The total length of a single ABF is 8
m and the diameter is 2 m. All ABFs in Figure 3.5b look identical, which indicates the
reproducibility of the fabrication process, therefore a batch of ABFs can be produced in one
substrate by this method. Helices of different geometries with a diameter ranging from 1 m
to 76 m and a total length ranging from 4 m to 280 m were fabricated.
Arbitrarily-shaped 3D microstructures can be created by DLW. By changing the CAD design
33
and choosing different photoresists, other non-helical microstructures were fabricated for
different research purposes. For example, 3D porous scaffold-like magnetic microrobots could
be used for 3D transportation of cells (Figure 3.6a-b) [88]. 3D microstructures with four bonds
arranged in different angles, hollow tubes, woodpile-like structures and microscale helicopters
were fabricated (Figure 3.6c-f).
Figure 3.6. Three-dimensional non-helical microstructures fabricated by DLW. (a-b) 3D porous scaffold-like
magnetic microrobots [88]. (b) 3D microstructures with four bonds arranged in different angles. (c) Hollow
microtube. (e) Woodpile-like microstructure. (f) Microscale helicopter.
34
3.3 Conclusion
In this chapter, the straightforward fabrication process of ABFs using DLW and electron beam
evaporation was introduced in detail. ABFs fabricated with the same process parameters look
identical, which indicates the reproducibility of the fabrication process. A large number of
ABFs can be produced in one substrate. This method can be used to create helical
microstructures as well as other complex 3D magnetic microstructures. The next chapter
considers the processes following fabrication, i.e. the actuation and swimming behavior of
ABFs under rotating magnetic fields.
35
4 Motion control of ABFs
4.1 The torque on an ABF
In Section 2.3.1, it was mentioned that the principle of magnetic actuation is to apply a
magnetic force or magnetic torque to a magnetic object. From Equation (2-2) and (2-3), we
know when a magnetic field is apply to a magnetic object, the magnetic force 𝐹𝑀 and torque
𝑇𝑀 acting on the body can be written as,
𝐹𝑀 = 𝜗(𝑀 ∙ ∇)𝐵 (4-1)
𝑇𝑀 = 𝜗𝑀 × 𝐵 (4-2)
where ϑ is the magnetic volume of the object, and B is the magnetic flux density. M is the
magnetization of the microrobot under the magnetic field B. In a hard magnetic material, the
magnetization of the object is independent of the applied field B. While in a soft magnetic
material, the magnitude of the vector M is dependent on the B field and the direction of M
depends on the geometry of the object.
ABFs were coated with soft-magnetic materials, Fe or Ni. When an ABF is exposed to a
uniform B field, where the gradient 𝛻𝐵 is zero, the soft-magnetic ABF is magnetized by the B
field. The torque 𝑇𝑀 drives the ABF and aligns it immediately to the B field. Once the direction
of M and B is the same, the torque 𝑇𝑀 becomes zero and the ABF maintains the alignment as
long as B is unchanged. An ABF is placed in an XYZ 3D coordinate frame (Figure 4.1a), and
we suppose the M of the ABF is perpendicular to the long axis of the ABF once it is under a B
field. When a uniform B field is applied, in which the direction of B is minus X (-X in Figure
4.1a), only a torque 𝑇𝑀 (Equation (4-2)) is generated. The torque 𝑇𝑀 brings M to align with B
and vanishes. The schematic in Figure 4.1b-c shows the ABF alignment on the XY plane and
XZ plane, respectively. When the direction of B rotates a number of degrees to B1 on the XZ
plane (Figure 4.1c), a new torque is generated, and the ABF again aligns with the B1 field,
which makes the ABF rotate a number of degrees along its long axis (Y direction in Figure
4.1a). When the B field is continuously rotated in a circle on the XZ plane, the ABF rotates
around its helical axis continuously. As mentioned before, a net-displacement is generated
when a helix rotates, which generates a translational movement to make the ABF move
forward. When the rotating axis of the B changes its direction on the XY plane (Figure 4.1b),
36
the direction of ABF movement changes accordingly on the horizontal plane. When the rotating
axis of B field changes its direction on the YZ plane, the ABF swims out of the horizontal plane
which enables 3D movement of the ABF. In summary, an ABF moves forward in a rotating
magnetic field B. By simply changing the rotational axis of the B field, we can steer the ABF
in 3D wirelessly. The following section discusses the magnetic setup which can generate the
required rotating B field.
Figure 4.1. The principle of motion control of an ABF in 3D under a uniform rotating magnetic field. (a) An ABF
in an XYZ coordinate system. The alignment of an ABF exposed to a B field on the XY plane (b) and the XZ
plane (c).
4.2 Magnetic setup
Normally three pairs of Helmholtz coils are used for the actuation of magnetic helical robots,
such as the ones in Figure 4.2a. The three pairs of Helmholtz coils are placed orthogonally to
generate a uniform rotating magnetic field around the center of the coils. Figure 4.2b shows
the principle of one pair of Helmholtz coils. It consists of a pair of conducting circular coils
with a specific configuration. In each coil there are N turns of wires. The radius of the two
circular coils is r. These coils are placed parallel to each other at a distance of r. When a current
I is applied to the coils in the same direction, a homogeneous magnetic field B is generated in
the X direction on the mid-plane (X=0 in Figure 4.2b) of the two coils, and the strength of B is
given as,
𝐵 =32 𝜋 𝑁 𝐼
5 √5 𝑟 × 10−7 Tesla (4-3)
An ABF aligns to B in the middle of one pair of these coils (Figure 4.2b). The direction of B
can be changed from X to –X (Figure 4.2b) by changing the direction of current, and the
strength of B can be easily tuned by changing the value of the current. When three pairs of
Helmholtz coils are placed in an orthogonal way (Figure 4.2a), the total B field is a sum of the
37
field generated by each pair and a B field in any 3D direction can be generated by independently
changing the current inputs in each of three pairs. By programing the input current in each pair,
different magnetic fields can be generated such as uniform rotating magnetic fields (Figure
2.9a), precession fields (Figure 2.9b), oscillating fields and On-off fields (Figure 2.9c-d).
Figure 4.2. The magnetic setup. (a) The three-pairs Helmholtz coils setup. Adapted from [89]. (b) Schematic of a
pair of Helmholtz coils. (c) Control flow of the magnetic experimental setup.
Figure 4.2c shows the control flow of the experimental setup. An ABF is placed in a tank
(Figure 4.2a-3) in the central space of the three coil pairs (Figure 4.2a-2). An objective lens is
placed above the tank to magnify the ABF (Figure 4.2a-1). Objective lenses of different
magnifications of 2×, 10×, 20× and 50× can be chosen depending on the size of the structures.
The reflected light is collected by a CCD camera above the lens. The real-time signals and
images are collected by the control software (Figure 4.2c). The software controls the visual
conditions of the camera and the input currents in the three pairs of coils via the current
controller (Figure 4.2c). In this system, the real-time swimming of ABFs can be visualized and
steered manually. Automatic control of ABFs is possible. Most researchers use a Helmholtz
38
coil setup to generate rotating magnetic fields, however actuation of helical swimmers using a
non-uniform rotating magnetic field generated by rotating permanent magnets is also possible
[71].
4.3 Swimming behavior of ABFs
Figure 4.3 shows the swimming velocities of an ABF as a function of rotating frequency (f) of
the external magnetic fields at a magnetic strength of 3 mT. The ABFs of 16 m in length and
5 m in diameter were made from IP-L photoresist by DLW with Ni/Ti (50 nm/5 nm) coating.
The ABFs on the substrate were untethered by micromanipulation, gently pushing an ABF
using a sharp tungsten probe (Type T-4-44 from GGB Industries, Inc. in US) mounted on a
micromanipulation stage. The experiments were conducted in deionized water (DI-water) and
the ABF swam on a clean and polished silicon (Si) wafer.
Figure 4.3. Swimming behavior of ABFs. The ABF was made from IP-L with Ni/Ti coating. The swimming
experiments were conducted at 3 mT. The swimming behavior can be divided into three regions, wobbling,
corkscrew and step-out regions. The insert shows the forward velocity (Vforward), drift velocity (Vdrift) and the total
velocity (Vtotal).
Overall, the ABFs show frequency-dependent swimming behavior and the swimming plot can
39
be divided into three regions: wobbling, corkscrew and step-out regions. In the wobbling
region, the ABF is actuated at a low frequency (f < 10 Hz) and wobbles around the helical axis
while moving forward. When the frequency increases, the ABF swims in a stable corkscrew
motion without wobbling, the swimming velocity increasing linearly until it reaches the max.
The input frequency, at which the ABF reaches its maximum speed, is called the step-out
frequency (fstep-out). When the frequency increases further, the ABF motion is in the step-out
region. In this region the velocity decreases dramatically while the frequency increases. Since
the ABFs rotate near a surface, the ABFs also have a drift speed perpendicularly to the helical
axis due to the surface effect. The total velocity of the ABF is the sum of the forward velocity
and drift velocity (the insert in Figure 4.3). The drift velocity is always present if the ABFs
move near a surface.
4.3.1 Wobbling region
Figure 4.4a-b shows the wobbling behavior of an ABF at a low frequency of 6 Hz. The ABF
moves forward while wobbling around the moving direction with a precession angle . The
precession angle decreases as the frequency increases [89].
Figure 4.4. The wobbling and corkscrew regions of ABF swimming behavior. (a) Schematic of the wobbling of
an ABF. (b) The time-lapse images of an ABF wobbling at a low frequency of 6 Hz. (c) Schematic of stable
corkscrew motion of an ABF. (d) The time-lapse images of an ABF swimming in a high frequency at 58 Hz, The
time step is 1/6 s per picture in (b) and (d).
40
Wobbling occurs when the magnetization M of the ABF is not perpendicular to the helical axis.
In Section 4.1, we assumed that the M of an ABF is perpendicular to its helical axis, however,
M is not always perpendicular to the ABF helical axis. An ABF has a nano-sized film of soft-
magnetic magnetic materials covering a 3D helical shape. Due to the shape anisotropy of the
magnetic materials and the helical shape of the ABF, the M of an ABF is rarely perpendicular
to its helical axis. There is an angle between the M of the ABF and the perpendicular direction
of the helical axis (y in Figure 4.5), which is called the misalignment angle.
Figure 4.5. Schematic of the misalignment angle X is the helical axis of the ABF and y is the perpendicular
direction of x.
4.3.2 Corkscrew region
As the rotating frequency of the B field increases, the precession angle gradually decreases,
and, finally, a corkscrew motion occurs (Figure 4.3). This is due to the total drag from the
surrounding fluid acting on the ABF being minimized if the ABF rotates around the helical
axis. Figure 4.4c-d shows the stable corkscrew motion of an ABF at a high frequency of 58 Hz.
In this region, the velocity of an ABF increases linearly as the frequency increases, which is
indicated by Equation (2-11). When the frequency reaches a critical value, called the step-out
frequency (fstep-out), the maximum velocity of the swimmer is reached.
4.3.3 The step-out frequency
The step-out frequency (fstep-out) is the maximum frequency at which the ABF can rotate
synchronously with the rotation speed of the B field. The step-out frequency depends on the
geometry of an ABF, the viscosity of the liquid, and the input magnitude of the B field [62]. It
is given by
41
𝑓𝑠𝑡𝑒𝑝−𝑜𝑢𝑡 =𝑎
𝑎𝑐−𝑏2 𝑇𝑀𝑚𝑎𝑥. (4-4)
where the parameters a, b, c are related to the geometry of the ABFs and the viscosity of the
liquid, while the maximum torque 𝑇𝑀𝑚𝑎𝑥 relates to the magnetic volume ϑ, the B field strength,
and the magnetization M of an ABF (Equation (4-2)). ABFs with different geometries, such as
diameter, total length, filament radius and helicity angle of helices, have a different step-out
frequency. For the same ABF, the step-out frequency increases when the applied B field
increases, and, thus, the ABF reaches a higher maximum speed. Figure 4.6 shows forward
velocities of an ABF as a function of frequency at strengths of 1 mT and 3 mT. The ABF was
made from ORMOCOMP and coated with 25 nm Fe. At 1 mT strength, the ABF had the
maximum speed of 5.3 m/s at the step-out frequency of 8 Hz, whereas the same ABF had the
maximum speed of 18.0 m/s at a step-out frequency of 28 Hz at 3 mT strength. It is reasonable
to predict that the step-out frequency and the maximum speed continuously increase by further
increasing the B field up to the limitation of the magnetic setup (< 10 mT). The viscosity of the
liquid also influences the step-out frequency in that the step-out frequency decreases when the
viscosity of the liquid increases [62].
Figure 4.6. The step-out frequencies of an ABF in different field strengths, 1 mT and 3 mT. The ABF was made
from ORMOCOMP and coated with 25 nm Fe. Adapted from [65].
42
4.3.4 Step-out region
When the frequency of the B field is higher than the step-out frequency of the ABF, the rotating
speed of the ABF can no longer follow the rotating speed of the B field. The velocity of the
ABF decreases dramatically as the frequency increases (Figure 4.3 and Figure 4.6). The
velocity of the ABF drops to zero when the frequency of the B field is further increased.
4.3.5 Drifting
Figure 4.3 shows that the ABF has both forward and drifting velocities. The forward velocity
is a component of the total velocity in the direction of the helical axis, while the drifting velocity
is a component in the perpendicular direction of the forward velocity (the insert in Figure 4.3).
The drifting of the ABF is due to the drag imbalance between the part of the ABF which is
closer to and the part which is away from the surface (Si wafer on which the ABF swims), also
called the wall effect [89]. When the ABF moves away from the solid surface, the drifting
velocity falls to zero since the wall effect does not exist. The drifting effect is obvious when
the ABF is moving in the wobbling region since the forward velocity is small in this region. In
the stable corkscrew region and some part of step-out region, the drifting velocity is negligible
compared to the increased forward velocity, and the plot of forward velocity overlays the one
of total velocity in Figure 4.3.
4.3.6 Swarm control
For biomedical applications such as drug delivery, a swarm of ABFs is necessary to increase
the drug loading capability. When a swarm of ABFs is placed in a uniform rotating magnetic
field, the external B field applies to all of the ABFs in its range. If the ABFs are fabricated
under the same parameters, they have the same or similar swimming behaviors. Figure 4.7
shows nine ABFs in the camera screen, which were fabricated in the same batch with the same
fabrication parameters, move synchronously in DI-water at 9 mT and 90 Hz. The swimming
direction of all the ABFs was tuned synchronously by simply changing the B field.
43
Figure 4.7. The swarm control of ABFs. The ABFs were made from IP-L and coated with 50nm Ni/5nm Ti. The
ABFs were actuated at 90 Hz at 9 mT and the average forward velocity of ABFs was 69.12 m/s.
When ABFs with different parameters, which in turn cause the ABFs to have different step-out
frequencies, are placed in a uniform magnetic field, different ABFs will act differently. For
example, some ABFs are actuated in the corkscrew region and others are actuated in the step-
out region in a same B field, which allows separate control of the movement of each ABF. The
individual control of three different ABFs made by a self-scrolling technique was demonstrated
by Peyer et al. [41].
4.4 Conclusion
In the chapter, motion control of ABFs in a uniform rotating magnetic field generated by three-
pairs of Helmholtz coils and swimming behavior of ABFs were explained. The swimming
behavior of ABFs is frequency-dependent. The velocity plot of ABFs as a function of the
frequency of the external magnetic fields can be divided into three regions, wobbling,
corkscrew and step-out regions. ABFs wobble at a low frequency, and move in a stable
corkscrew motion in the corkscrew region. In this region, the swimming speed increases
linearly as the frequency increases. The speed reaches a maximum value when the step-out
frequency is reached. In the last region, the velocity of the ABFs drops dramatically when the
frequency is higher than the step-out frequency of the ABFs. Finally, the swarm control of
44
multiple ABFs with similar swimming behaviors was demonstrated and nine ABFs moved and
responded to the magnetic fields synchronously. The individual control of several ABFs with
different swimming behaviors in the same rotating B field was discussed.
45
5 Cytotoxicity of ABFs
As previously mentioned, ABFs have great potential in biomedical applications, such as
targeted drug delivery. For biomedical applications, the biocompatibility or safety of the
devices should be taken into consideration. This chapter focuses on the studies of the
cytotoxicity of the ABFs, which is a preliminary test of biocompatibility. It starts with a general
introduction of biocompatibility, followed by in vitro cytotoxicity tests of the ABFs. The ABFs
were made from different materials including three photoresists for making the polymeric
helical bodies and two different metal coating (Ni/Ti and Fe/Ti). Results shown in this chapter
have been previously published in two journal papers [57, 65].
5.1 Introduction
The definition of biocompatibility, which has been widely used in the community of
biomaterials and medical devices, was described by Williams in 1987 as “the ability of a
material to perform with an appropriate host response in a specific application” [90]. In 2008,
Williams re-defined the definition of biocompatibility as “Biocompatibility refers to the ability
of a biomaterial to perform its desired function with respect to a medical therapy, without
eliciting any undesirable local or systemic effects in the recipient or beneficiary of that therapy,
but generating the most appropriate beneficial cellular or tissue response, and optimizing the
clinically relevant performance of that therapy.” [91]. The biocompatibility of a material/device
depends on several factors: (1) the location of the material/device inside of a human body, (2)
the length of time the material/device is placed inside a body, so called short- and long-term
biocompatibility, and (3) the material properties itself, such as chemical composition, surface
topography and surface energy.
Biocompatibility includes in vitro assessment of cell and tissue compatibility and in vivo
assessment of tissue compatibility [90]. The biocompatibility of a device/material can be tested
or assessed using in vitro and in vivo assays, which are well-defined by national regulatory
agencies and are generally used by the biomaterials community. For example, the IOS 10993
standards “Biological evaluation of medical devices” are regulated by the US Food and Drugs
Administration (FDA). These tests primarily focus on toxicology in vitro and the foreign-body
reaction in vivo [90]. The aim of this chapter is to examine the cytotoxicity of ABFs in vitro, a
preliminary but important test of biocompatibility. Cell proliferation, cell viability and cell
46
morphology after contact with the same materials used to make ABFs or ABFs themselves are
presented.
5.2 ABFs made from IP-L and SU-8 with Ni/Ti coating
5.2.1 Cell proliferation tests
During the fabrication process, the DLW tool allows us to fabricate ABFs with different
photoresists, such as IP-L, SU-8 and ORMOCOMP, and e-beam deposition allows us to deposit
different metals including magnetic and non-magnetic materials. Two different photoresists IP-
L and SU-8 were used for fabrication of the polymeric helical bodies, and Ni was used as the
magnetic material. In order to improve the biocompatibility of ABFs, a thin layer of titanium
(Ti) was deposited and covered the Ni layer on the ABFs. The potential cytotoxicity of these
ABFs was investigated by assessing the cell proliferation and cell morphology after a 3-day
incubation period with cells.
Materials and methods
Cell Culture: C2C12 mouse myoblast cells (mammalian muscle cells) were chosen to conduct
all of in vitro cytotoxicity tests in this chapter, since skeletal muscle, as our largest tissue and
whole body homeostatic maintenance organ, is an extremely important target with which to
test cytotoxicity. The cells were purchased from the American Type Culture Collection
(ATCC) (LOT: 58127344), maintained in T75 tissue culture flasks (SPL Life sciences) at 37°C,
7% CO2 and split every second day (while still below 50% confluence) at 1:10 for propagation.
Proliferating myoblasts were maintained in Growth Media (GM) consisting of Dulbecco’s
modified eagle’s medium (DMEM) (from Invitrogen), 20% fetal bovine serum
(Lot#41F0482K, from Invitrogen) and were supplemented with 2mM L-Glutamine (from
Invitrogen).
Sample preparation: Flat circular glasses (12 mm in diameter) coated with the same materials
as ABFs were used to represent the ABFs for cell proliferation tests. First, the photoresist, IP-
L or SU-8, was spin-coated onto the flat glass substrates. Next, Ni/Ti bilayers were evaporated
onto the photoresist substrates with a thickness of 10 nm and 5 nm, respectively. The sample
coated with IP-L is referred to a “IP-L with Ni/Ti”, and “SU-8 with Ni/Ti” represents the
sample coated with SU-8.
47
Cell proliferation tests: Clean circular glasses without coating served as a control [92]. The
glass samples with and without coating were sterilized in 10% penicillin/streptomycin (from
PAA) in phosphate-buffered saline (PBS, from Invitrogen) solution for 30 minutes followed
by a single rinsing with fresh PBS for 5 minutes. C2C12 cells (1×104 cells/ml) were then seeded
onto the samples in GM at 37°C with 7% CO2 and incubated for 72 hours. Micrographs were
taken with an inverted light microscope every 24 hours (Figure 5.1), and the sampled cell areas
in the images were taken randomly. The cells were counted using ImageJ software and the cell
counter plugin.
Figure 5.1. The optical images of cells on different samples after one-day incubation. (a) Cells on control. (b)
Cells on IP-L with Ni/Ti. (c) Cells on SU-8 with Ni/Ti. Reused from [57].
Results and discussions
Figure 5.2 shows a comparison of the proliferation of C2C12 cells on three different samples,
Control (glass), IP-L with Ni/Ti and SU-8 with Ni/Ti, after 72 hours of incubation. On the “SU-
48
8 with Ni/Ti” sample, the cells grew at approximately the same rate as they did on the control.
The proliferation was slower on the glass coated with IP-L and Ni/Ti bilayer, which may be
due to the observed irregularities of its surface (Figure 5.1b) compared to two other samples
(Figure 5.1a,c) [93]. The results indicate that the C2C12 cells readily tolerated the surface
materials and continued to proliferate on Ni/Ti bilayers in three days.
Figure 5.2. Comparison of cell concentrations on different samples (flat glass, IP-L and SU-8 coated with Ni/Ti
bilayers). The error bars showed the standard deviation from four different cell culture regions. Reused from [57].
5.2.2 Cell morphology tests
Another method to monitor the viability of cells is to examine their interactions with their
surroundings. Healthy cells are more proliferative, which requires focal adhesion contacts with
their substrates for dynamic lamellipodial and filopodial attachment and retractions. The
interaction between the ABFs and C2C12 mouse myoblasts was analyzed. For these studies
the cells were cultured on substrates with both horizontal and vertical arrays of ABFs. After 24
or 72 hours of incubation in GM, the samples were gently rinsed twice with PBS. The cells
were then fixed in 3.7% formaldehyde in PBS for 10 minutes followed by a second fixation of
2.5% glutaraldehyde in PBS overnight. The samples were rinsed the next day with fresh PBS
49
and then immersed in 10%, 25%, 50%, 75%, 90% and 100% ethanol in water for dehydration.
Each step was maintained for 5 minutes. The samples were then dried in a critical point dryer
(CPD 030, BAL-TEC) and sputtered with a thin film of platinum as a conductive layer for SEM
inspection.
Figure 5.3 shows SEM images of the interaction between the cells and the ABFs. These results
verify that for both the horizontal (Figure 5.3a-b) and vertical (Figure 5.3c-d) arrays of ABFs
prepared from IP-L and SU-8, respectively, the cells adhered well and extended apparently
normal lamellipodia and filopodia over the surface of the devices within three days.
Figure 5.3. SEM micrographs of cells resting on the horizontal array of IP-L helices (a-b) and vertical array of
SU-8 helices (c-d) after three days culture, respectively. The insert in (b) shows the SEM image of a cell contacting
part of the ABF by the lamellipodia and filopodia after one-day culture, and the scale bar is 1 m. Reused from
[57].
5.3 ABFs made from ORMOCOMP and Fe/Ti coating
Although the Ti layer improves the biocompatibility of ABFs that have IP-L or SU-8 as the
polymeric bodies and Ni as the magnetic material, the potential cytotoxicity of Ni element [94]
may be an issue, especially when Ni is oxidized into Ni ions (Ni2+). Ni ions showed dose-
dependent cytotoxicity to fibroblast cells [95] and osteoblasts [96]. Pure Fe has better
biocompatibility compared to Ni, and some researchers claim that pure Fe could be
50
biocompatible and even biodegradable materials for large-sized stents [97-99]. In this work,
we chose biocompatible ORMOCOMP to produce the helical bodies of ABFs and replaced Ni
with Fe as the magnetic material for actuation. The in vitro cytotoxicity of Fe-based ABFs was
assessed using 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay
and cell morphology.
5.3.1 Fabrication of ABFs from ORMOCOMP and Fe/Ti coating
ORMOCOMP, a member of the ORMOCER® family, is a commercially available gel resist
composed of inorganic-(Si-O-Si)-organic groups [92, 100]. While the detailed chemical
composition of ORMOCOMP is proprietary, previous results have shown that ORMOCOMP
is biocompatible, and supports cell proliferation, viability and normal morphology of various
cell lines [84, 92, 101]. A growing interest in ORMOCOMP in biomedicine and tissue
engineering exists [101, 102], where hollow microneedles for transdermal drug delivery [84,
103] and elastic three-dimensional scaffolds for single-cell force measurements have been
developed [100, 104].
ORMOCOMP (Micro resist technology GmbH) was coated on the clean glass substrate by a
droplet. The 3D helices were then written in the resist using a DLW tool (from Nanoscribe
GmbH) with the oil-immersion 100× objective (NA = 1.4 from Zeiss, NA denotes numerical
aperture). The two important writing parameters, laser power and scan speed, were 5 mW and
25 m/s, respectively. Two-minute post bake at 80ᵒC followed after writing. The sample was
immersed in Ormodev developer (from micro resist technology GmbH) for 7 minutes to
remove the un-polymerized resist and then rinsed in isopropanol (IPA) immediately followed
by a gentle nitrogen blow. Then, Fe or Fe/Ti thin films were deposited using electron beam
deposition with a rotational speed of 10-15 rpm. The tilt angle for deposition was 15ᵒ in order
to reduce the shadowing effect [57]. The deposition rate was 0.6-2 Å/s.
Figure 5.4 shows the SEM image of arrays of ABFs. The mean length and diameter of one
helix is 15.7 m and 5.3 m, respectively (the inset of Figure 5.4). A single helix can be
produced every 2.3 s at a writing speed of 25 m/s. ORMOCOMP alleviates the need for spin-
coating and pre-backing steps, thus, saving fabrication time compared to SU-8. ORMOCOMP
also has better cell curability in biomedical applications [101, 105], as it does not require further
surface modifications such as O2 plasma treatment and chemical treatment, as does SU-8 due
to its hydrophobicity [105]. After the fabrication of helical arrays from ORMOCOMP, a Fe (25
51
nm) or Fe/Ti (25 nm/ 15 nm) thin film was coated onto the bodies by electron beam deposition.
We refer to “Fe-ABFs” and “FeTi-ABFs” to represent Fe-coated ABFs and Fe/Ti-coated ABFs,
respectively.
Figure 5.4. Helical microstructures made from ORMOCOMP using DLW, and the inset is a magnification of a
single helix. Reused from [65].
5.3.2 Cell viability of ABFs assessed by MTT assay
Materials and methods
Preparation of samples for MTT test: The MTT test (introduced in the next paragraph) was
used to check the cell viability of ORMOCOMP/Fe and ORMOCOMP/Fe/Ti layers. A thin
layer of ORMOCOMP was spin-coated on a clean glass substrate with a diameter of 12 mm,
and a Fe layer or Fe/Ti bilayer was deposited on the ORMOCOMP layer by e-beam deposition
to mimic the material composition of ABFs.
MTT assay: The MTT assay is a simple, rapid and non-radioactive process, which has been
broadly used to assess the cell viability, proliferation or cell cytotoxicity in vitro [94, 106-109].
For most living cells, the mitochondrial activity is constant. Therefore, a change in the number
of living cells is linearly related to the mitochondrial activity. Mitochondrial activity is
reflected by a reduction of the MTT (yellow color) into formazan crystals (purple color) (Figure
5.5). Hence, any decrease or increase in living cell number can be identified by tracking the
52
formazan concentration, which can be simply characterized by measuring the optical density
of the sample using a plate reader [110].
Figure 5.5. The chemical basis of MTT assay: the reduction of MTT to MTT formazan. Adapted from [106].
The prepared samples were sterilized in 10% penicillin/streptomycin (from PAA) in phosphate-
buffered saline (PBS, from Invitrogen) solution for 30 minutes, followed by rinsing twice with
fresh PBS for 5 minutes. C2C12 mouse myoblasts (from the American Type Culture Collection
(ATCC) (LOT: 58127344)) were seeded onto the substrate films in 24-well plates at a density
of 2.5×105 cells·ml-1 and cultured within a standard tissue culture incubator at 37ᵒC with 7%
CO2. The GM for culturing the C2C12 myoblasts consisted of DMEM (Invitrogen)
supplemented with 20% fetal bovine serum (Lot#41F0482K; Invitrogen) and 2mM L-
Glutamine (Invitrogen)). After incubating the cells on the substrates for 24 hours, the GM was
replaced with 500 l of MTT reagent (2.5 mg/ml MTT in PBS; Sigma No. M2128). Following
incubation for 3 hours the MTT reagent was replaced by 200 l dimethyl sulfoxide (DMSO),
shielded from light with aluminum foil and placed on a shaker (54 rpm; Stuart, mini gyro-
rocker, SSM3) for 10 minutes. Subsequently, 40 l of solution from each well was transferred
to the 96-well plate, mixed with 40 l DMSO, and the optical density measured by a microplate
reader at 540 nm wavelength. The tests were done for three time periods, 24 hours, 48 hours
and 72 hours.
SEM of cell samples: In order to examine how cells interact with ORMOCOMP/Fe surface,
C2C12 cells (2.5×105 cells·ml-1) were seeded onto the glass substrate with ORMOCOMP/Fe
layers and incubated in their presence for 24 hours. The cells on the sample were then
chemically fixed, dehydrated and dried in a critical point dryer (CPD 030, BAL-TEC) as
previously described in Section 5.2.2. Next, a thin film (about 10 nm) of Platinum was sputtered
onto the sample for SEM inspection (Zeiss ULTRA 55).
53
Results and discussions
Cell viability was assayed on two-dimensional surfaces coated with ORMOCOMP/Fe or
ORMOCOMP/Fe/Ti layers by MTT, as previously described [111]. Figure 5.6a shows C2C12
mouse myoblasts grown on distinct substrates after 24, 48 and 72 hours. The clean glass
substrate served as “Control” [92]. By comparing the optical intensity of formazan between
cells grown on different substrates (Figure 5.6b), we obtained a measure of cytotoxicity of the
distinct components. Figure 5.6a shows that the cell viability of ORMOCOMP/Fe film is
somewhat lower than that of clean glass (control) after 24 hours of incubation, whereas SEM
Figure 5.6. Cell viability of Fe-ABFs and FeTi-ABFs. (a) MTT tests of cell viability for C2C12 mouse muscle
cells cultured on clean glasses (showed as ‘Control’), ORMOCOMP coated glasses with a Fe metal layer (showed
as ‘Fe’), and ORMOCOMP coated glasses with Fe/Ti layers (showed as ‘FeTi’) after incubation of 24 hours, 48
hours and 72 hours, respectively. The optical intensities were normalized to the intensities of the control. The
error bars show the standard deviation of mean values where n = 8. (b) Optical image of MTT results after two-
day incubation. (c) Cells spread on the entire ORMOCOPM/Fe surface after seeding (2.5×105 cells·ml-1) for 24
hours. Reused from [65].
images (Figure 5.6c) do show that cells spread and make functional membrane contacts on the
ORMOCOMP/Fe surface at this time point. Indeed, ORMOCOMP/Fe/Ti appear to enhance
cell viability after 72 hours of incubation, as previously described for rat myoblasts on titanium
54
containing structures [112].
We also conducted MTT assays on substrates that were never exposed to cells to test for the
possibility that these materials might produce a MTT signal on their own after 72 hours of
incubation. Figure 5.7a shows the optical intensity of three types of samples, Control, Fe, FeTi.
The comparison of MTT viability determination in the presence and absence of cells after 72
hours incubation in Figure 5.7b shows that none of these materials produced a sufficient
amount of MTT formazan to invalidate our initial result.
Figure 5.7. Comparison of MTT values with and without cells. (a) MTT assay run on Control (glass), Fe and FeTi
surfaces without the presence of cells after 72 hours incubation. (b) Comparison of MTT viability determination
in the presence and absence (extracted from panel (a) and compressed to fit in the same plot) of cells after 72
hours incubation. All values are normalized to the control condition in the presence of cells. Reused from [65].
5.3.3 Morphology of cells on ABFs
The MTT assay shows that the ORMOCOMP/Fe and ORMOCOMP/Fe/Ti films are not
cytotoxic to C2C12 cells. The cell morphology was inspected by SEM to examine their
interactions with their surroundings. C2C12 cells (1×105 cells·ml-1) were seeded onto glass
substrates with incorporated arrays of Fe-coated ABFs and incubated in their presence for 24
hours. The cells on the sample were then chemically fixed, dehydrated, dried in a critical point
dryer (CPD 030, BAL-TEC) and inspected. Figure 5.8 shows C2C12 cell contacting and
spreading on Fe-ABFs (Figure 5.8a-b) and FeTi-ABFs (Figure 5.8c-d) with lamellipodial and
filopodial interactions of normal morphology [53, 57, 63].
Iron is oxidized in an aqueous biological environment. Once the Fe coating is degraded the
55
ABFs will lose their swimming functionality which is driven by magnetic actuation. Figure
5.8b shows a portion of an ABF with the surface roughened due to the corrosion of the exposed
iron surface in physiological saline [113]. By contrast, Figure 5.8d shows the comparably
smooth surface of an ABF whose iron surface was coated by a thin (15nm) coating of Ti that
is highly resistant to corrosion [114]. Hence, the functional life-time of ABFs may be extended
for in vivo applications by coating the surface of Fe-ABFs with anticorrosive Ti.
Figure 5.8. Cell morphology on Fe-ABFs and FeTi-ABFs after one-day incubation. (a) Cells on Fe-ABFs, (b)
Magnified image of rectangular area shown in (a), (c) Cells on FeTi-ABFs, (d) Magnified image of rectangular
area shown in (c). The white scale bar in (a) and (c) is 4 m, and the black scale bar in (b) and (d) is 1 m. Reused
from [65].
Our experiments were conducted in vitro. A previous study by Waksman et al. demonstrated
the safety of iron stents (10 mm in length) in porcine coronary arteries after 28 days with no
evidence of embolization, thrombosis, excess inflammation or fibrin deposition [115].
Furthermore, Peuster et al. studied the long-term biocompatibility of iron stents (20 mm in
length) in the porcine descending aorta for 360 days and showed no evidence for local toxicity
due to corrosion products adjacent to the iron stent [99]. They claimed iron is a suitable metal
for large-size degradable stents with no local or systemic toxicity. We can, therefore, assume
that our Fe-ABFs, which are three orders of magnitude smaller than the iron stents, should be
safe in vivo and maintain functionality. Another advantage of iron for in vivo applications is its
micronutrient value for normal metabolism that is required for efficient systemic gas exchange
56
and mitochondrial function [115].
5.3.4 Swimming performance of Fe-based ABFs
The swimming tests were conducted in DI-water under a Helmholtz coil setup shown in Figure
4.2a. The results in Figure 5.9 show the forward velocity of Fe-ABFs and FeTi-ABFs as a
function of input frequency driven by two magnetic fields with strengths of 1 mT and 3 mT.
The swimming tests were conducted immediately after Fe coating and Fe/Ti coating. The
curves show the standard frequency-dependent swimming behaviors of ABFs. The forward
speed of Fe-ABFs and FeTi-ABFs increases with increasing input frequency below the step-
out frequency. When the frequency exceeds the step-out frequency, the forward speed of the
ABFs drops dramatically. The step-out frequency increases with the strength of the magnetic
field. The maximum forward speed of Fe-ABFs was 48.9 m/s (3.1 body lengths per second)at
the field strength of 9 mT at 72 Hz. After Ti coating, the maximum speeds of ABFs at 1 mT
and 3 mT slightly decreases compared to Fe-ABFs which is probably due to the additional
weight of nonmagnetic Ti.
Figure 5.9. Swimming behaviors of Fe-ABFs and FeTi-ABFs. (a) Forward velocities of Fe-ABFs and FeTi-ABFs
as a function of frequency at two different field strengths of 1 mT and 3 mT. The error bars represent the standard
deviations of mean values from three measurements at each frequency. (b) The swimming performance of an Fe-
ABF at 9 mT at 72 Hz. All swimming tests were conducted in DI-water. Adapted from [65].
5.4 Conclusion
In this chapter the in vitro cytotoxicity of ABFs made from different materials was studied
using C2C12 cells. For the ABFs, which have IP-L or SU-8 as the polymeric bodies and Ni/Ti
layers as coating materials, the outside Ti layer covers the Ni layer and improves the
57
compatibility of the ABFs. The cell proliferation and cell morphology tests show that the ABFs
did not significantly influence the cell proliferation and cells continuously proliferated on the
surface in 3-days culture. For the ABFs, which have biocompatible ORMOCOMP and Fe as
the magnetic material, viability assessment and cell morphology tests reveal little cytotoxicity
for up to three days in culture. The ABFs showed non-cytotoxic and were precisely controlled
by low-strength magnetic fields. For future work, besides the in vitro cytotoxicity tests, a
detailed examination of biocompatibility of ABFs (including in vitro and in vivo cell and tissue
compatibility [90]) according to the defined standards, for example the ISO 10993 standards
by the FDA, should be conducted. These tests should be carried out depending on the place
where ABFs are used and the period the ABFs are used for, especially for in vivo applications.
58
6 Functionalization of ABFs for potential
biomedical applications
6.1 Introduction
Since ABFs can be wirelessly steered in liquid, they can be used to manipulate or transport
micro and even smaller objects in close or open fluidic environments. Previous work showed
that ABFs can be used to manipulate cellular and sub-cellular objects by direct pushing [41,
57] and non-contact methods (agitating the peripheral liquid when an ABF is rotating) [6, 13].
However, for biomedical applications such as drug delivery and wireless sensing, further
surface bio-functionalization with specific chemicals, such as drug molecules and chemicals,
is required [63]. In this chapter, we present the functionalization of ABFs with lipid-based drug
carriers for controlled release and single-cell targeted delivery in vitro, and then move further
to in vivo tracking of ABFs by functionalizing them with near-infrared dyes. Some results in
this chapter have been previously published in three journal papers [66-68].
6.2 ABFs functionalized with liposomes for drug delivery in vitro
In biological and medical fields, liposomes (also called lipid vesicles) have been extensively
studied for various applications including drug delivery systems and cell membrane science
[116-118]. Liposomes are formed based on self-assembly of lipid molecules, which have both
hydrophobic and hydrophilic parts (Figure 6.1a), in aqueous solutions. A liposome is a lipid
vesicle consisting of a self-assembled lipid bilayer (Figure 6.1b) and a liquid compartment in
which DNA, drugs and/or chemicals can be encapsulated (Figure 6.1c).
Liposomes range in size from 20 nm to several hundred m. They can be classified based on
size and lamellarity. Small unilamellar vesicles (SUV) have a diameter of less than 100 nm and
a single bilayer. Large unilamellar vesicles (LUV) have a size between 100 to 1000 nm in
diameter. Giant unilamellar vesicles (GUV) have a diameter of larger than 1 m. Multilamellar
vesicles have several bilayers, while a multivesicular body has several smaller vesicles
encapsulated in one big vesicle (Figure 6.1d) [116].
59
Figure 6.1. The structures of lipid vesicles. (a) A typical lipid molecular with a hydrophilic head and two
hydrophobic tails. (b) Lipid molecules are self-assembled to a bilayer of lipids in an aqueous solution. (c) A lipid
vesicle. (d) The classification of lipid vesicles and their general sizes. Adapted from [116].
As drug carriers, liposomes have the ability to carry both hydrophobic and hydrophilic drugs
benefiting from the amphiphilic properties of lipids. The drugs inside liposomes can be released
when the outer lipid bilayers become unstable and begin to leak. The composition of the lipids
which form liposomes significantly influences the stability of the liposomes. Depending on the
lipid composition, the payload inside the liposomes can be locally and remotely trigger-
released by different stimuli, such as enzymes, pH, ultrasound, light and heat (Figure 6.2)
[119].
The aim of this work is to functionalize the surface of ABFs with liposomes for potential
biomedical applications, such as targeted therapy, genetic transfection and sensing/marking
(Figure 6.3). The combination of liposomes with ABFs, which have the ability of controlled
3D navigation in liquid, may create a multifunctional system which can be wirelessly
controlled and allow targeted drug delivery to specific areas in hard-to-reach areas, such as in
human bodies and in lab-on-a-chip environments.
60
Figure 6.2. The trigger-release mechanisms of different liposomes for drug delivery. Reused from [119].
Figure 6.3. Schematic of functionalization of ABFs with liposomes.
6.2.1 Fabrication of ABF suspensions
Arrays of ABFs tethered on a substrate were fabricated using DLW and e-beam evaporation.
61
A batch process is required to release the tethered devices from the substrate in order to obtain
an ABF suspension in a reasonable amount of time. We report on a time-efficient and low-cost
method to release the ABFs from the original substrate by sonication. All the ABFs were
released in a few minutes. The sonication does not influence the helical shape and swimming
properties of ABFs. The batch release protocol ensures that suspensions of ABFs in different
solutions can be obtained with relative ease and paves the way for their further functionalization
and their applications in biology and biomedicine.
Arrays of ABFs were fabricated by DLW from IP-L photoresist and coated with Ni/Ti (50 nm/
5 nm) using electron beam deposition (Figure 6.4a). Each ABF has a total length of 16 m and
a diameter of 5 m. A schematic of the batch release process is shown in Figure 6.4b. After
fabrication, the glass substrate was cut by a diamond cutter into a 1 cm × 1 cm piece with the
ABF arrays around the center (Figure 6.4b-1), then the small piece was immersed in
isopropanol (IPA) in a glass bottle. Subsequently the bottle was placed in a sonication bath for
1 min at a frequency of 132 KHz (Figure 6.4b-2). Finally the liquid in the bottle was collected
by a pipette and transferred into a centrifuge tube. Centrifuging allows one to control the ABF
concentration and the solvent in the suspension (Figure 6.4b-3) by altering the amount and type
of the liquid.
Figure 6.4. Release of ABFs by sonication. (a) As-fabricated ABF arrays with IP-L polymeric helical bodies
coated by Ni/Ti bilayers. (b) Schematic of processing procedures to untether ABFs in solution using sonication.
62
The ABF array was inspected under an optical microscope before and after the sonication.
Figure 6.5a shows an array of 630 helices on the glass substrate before the sonication process.
A magnified image in Figure 6.5b shows ABFs still anchored to the glass. After one minute of
sonication all of the ABFs were released from the substrate (Figure 6.5c-d); the helical shadows
in Figure 6.5d come from the contrasts between the metal and glass after the ABFs were
detached. The insert in Figure 6.5d is a zoomed image, obtained with a SEM, and confirms that
the ABFs were released from the substrate (the nine lines are the remains of the contact points
between ABFs and the glass substrate).
Figure 6.5. Images of ABF arrays before and after sonication. (a) Arrays of ABFs on the glass substrate. (b)
Magnified image of part of (a). (c) Glass substrate after 132 kHz sonication for one minute. (d) Magnified image
of (c). Inset: SEM image of the substrate after batch release. The scale bar is 6 m.
The ABFs in suspension were also inspected after sonication. A droplet from the ABF
suspension (Figure 6.4b-3) was dried on a glass slide and observed under an optical
microscope. Figure 6.6a shows that most ABFs kept their helical shape, and the difference
between “ABF-1” and “ABF-2” is due to the different contact points between ABFs and the
glass slice (“ABF-1” is similar to the helix in Figure 6.6b while “ABF-2” resembles the helix
in Figure 6.6c). Figure 6.6d shows the magnified image of the rectangular part in Figure 6.6c.
This part was not fully coated with metals since it had contact with the substrate during the
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metal coating. Two different layers can be observed; the outside layer is the metal bilayers
while the inside part is the IP-L polymer body. After sonication, the metal layer on the outside
of ABFs was not detached and remained in contact.
Figure 6.6. A close look at ABFs after sonication. (a) Untethered ABFs obtained from the ABF suspension. (b-c)
SEM images of untethered ABFs. (d) Magnified image of rectangular part noted in (c).
Next, we checked the swimming performance of the ABFs released by sonication and
compared them with the ABFs released by “Micromanipulation” (using a tungsten probe
mounted on a micromanipulator to mechanically detach ABFs), which was mentioned in
Section 4.3. The step-out frequency fstep-out is an important parameter for ABFs. It can be
expressed as Equation (4-4) and is repeated here as follows,
𝑓𝑠𝑡𝑒𝑝−𝑜𝑢𝑡 =𝑎
𝑎𝑐 − 𝑏2 𝑇𝑀𝑚𝑎𝑥
The a, b, c parameters are related to the geometry of ABFs and the viscosity of the liquid, while
the maximum torque 𝑇𝑀𝑚𝑎𝑥 relates to the magnetic field strength and the magnetization of
ABFs. For the ABFs detached by two methods (the micromanipulation and the sonication),
they were fabricated by the same parameters. The swimming tests were conducted in the same
environment (DI-water) and at the same magnetic strength (3 mT). Theoretically, their step-
out frequency should remain the same or similar if the two methods do not make too much
change to the ABFs during their detachment. The plot in Figure 6.7 shows that the average
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step-out frequency of ABFs released by micromanipulation and sonication are similar, around
55 Hz. This implies that the batch release method by the sonication does not detach the surface
coating of metal films of ABFs and does not influence the swimming properties of ABFs
compared to the previous release method. The ABFs untethered by sonication exhibit a
maximum swimming velocity of 75 m/s under a 3 mT magnetic rotating field.
Figure 6.7. Comparison of step-out frequencies of the untethered ABFs by micromanipulation and sonication. The
error bars show the standard deviation from four different ABFs. All swimming tests were conducted in DI-water
with an applied magnetic field strength of 3 mT.
In conclusion, all of the ABFs were batch-released by sonication in one minute at 132 KHz.
After sonication most ABFs keep their helical shapes without structurally affecting the outside
metal bilayers. The ABFs retain their swimming properties with a maximum swimming
velocity of 75 m/s using a rotating magnetic field of 3 mT. This batch release protocol ensures
that suspensions of ABFs in different solutions can be obtained with relative ease and paves
the way for their further functionalization and their applications in biology and biomedicine.
6.2.2 Functionalization of ABFs with temperature-sensitive liposomes
Temperature-sensitive liposomes have been proposed for local hyperthermia treatments in
cancer therapy [120, 121]. Dipalmitoylphosphatidylcholine (DPPC) is commonly used as the
key component for temperature-sensitive liposomes. DPPC has a phase transition temperature
of 41ºC, at which the bilayer of liposomes switch from the solid phase to the liquid-gel phase
65
and become leaky releasing their encapsulated cargo [122]. Adding small amounts of lysolipids
(such as 10% 1-myristoyl-2-stearoyl-sn-glycero-3-phosphocholine (MSCP)) to DPPC
liposomes increases the drug release rate of DPPC liposomes at 39ºC-42ºC [123, 124].
Here, we report on the successful functionalization of the ABFs with temperature-sensitive
dipalmitoylphosphatidylcholine (DPPC)-based liposomes. Liposomes coating on the surface
of ABFs was confirmed using quartz crystal microbalance with dissipation monitoring (QCM-
D) and fluorescent probes. The functionalized ABFs (f-ABFs) showed the ability to incorporate
both hydrophilic and hydrophobic drugs. Finally, thermally-triggered release of calcein (a
common drug analog) from f-ABFs was demonstrated. These f-ABFs have the potential to be
used in targeted and triggered drug delivery, microfluidic devices and biosensing.
Materials and methods
Materials: The photoresist IP-L was purchased from Nanoscribe GmbH, Germany.
Dipalmitoylphosphatidylcholine (DPPC), 1-myristoyl-2-stearoyl-sn-glycero-3-
phosphocholine (MSPC) and lissamine rhodamine B lipids were purchased from Avanti Polar
Lipids, Inc.. Sodium chloride (NaCl), calcein disodium salt and 4-(2-hydroxyethyl)piperazine-
1-ethanesulfonic acid (HEPES) were purchased from Sigma Aldrich Chemie GmbH, Buchs,
Switzerland. The HEPES buffer solution was prepared with 10 mM 4-(2-hydroxyethyl)-1-
piperazineethanesulfonic acid and 0.15 M sodium chloride in Milli-Q water (Milli-Q gradient
A10, Millipore, resistivity 18.3 MΩ∙cm). The pH of the buffer was adjusted to pH 7.4 by a 6
M NaOH solution.
Fabrication of ABFs: ABF arrays were fabricated using direct laser writing (DLW) and e-
beam deposition methods. An individual ABF had a length of 16 μm and a diameter of 5 μm.
The filament cross section had an ellipsoidal shape with a thickness of approximately 1.19 μm
along the short axis and 2.23 μm along the long axis (Figure 6.8). The polymeric helices were
coated with Ni and then Ti layers (25 nm Ni and 15 nm Ti) using electron beam deposition.
The Ni thickness was reduced to 25 nm to improve the transparency of the microswimmers and
consequently the quality of images acquired by confocal laser scanning microscope (CLSM).
The Ti layer is naturally oxidized to TiO2 when exposed to oxygen.
66
Figure 6.8. SEM image of high and low magnifications of ABFs with Ni/Ti (25/15 nm) and measurements of the
filament thickness. (a) SEM image of an ABF at high magnification. (b) SEM image of a broken ABF. (c)
Measurement of the ABF filament thickness along its short axis. (d) Measurement of the ABF filament along its
long axis. Scale bar is 2 µm. Reused from [66].
Preparation of liposome-coated ABFs: The three-part preparation flow of liposome-coated
ABFs is shown in Figure 6.9. First, unilamellar DPPC liposomes are prepared. Second, the
ABF suspension is prepared, and third, the mixture of the two suspensions and washing
generates functionalized ABFs (f-ABFs).
The unilamellar DPPC liposomes were prepared by extrusion (Figure 6.10) [125, 126]. DPPC
lipids in chloroform were completely dried in a glass vial under a gentle N2 flow for 30 minutes
and rehydrated with HEPES buffer. In this step, fluorescent molecules can be dissolved in the
HEPES buffer to be incorporated within the liposomes. The glass vial was subsequently
vortexed to create multilamellar vesicles. The multilamellar vesicle suspension was transferred
into a glass syringe and assembled to form the extruder (Figure 6.10a). The lipid solution was
extruded 31 times through two packed polystyrene membranes (Figure 6.10b) to form uniform-
sized (200 nm) unilamellar vesicles. Extra care was taken to keep the entire extruder system
including the lipid solution above the transition temperature (41ºC) during the extrusion by
pre-warming the system at 65ºC in an oven. The DPPC/MSPC (9:1 w/w) was prepared by
67
adding 10% MSPC lipids in DPPC lipids before drying. All lipid mixtures were dissolved in
buffer at 2.5 mg/ml concentration.
Figure 6.9. Preparation flow for coating ABFs with unilamellar DPPC liposomes. Reused from [67].
68
Figure 6.10. The extruder used to produce unilamellar liposomes. (a) Assembled extruder used to obtain the
desired size of liposomes in an oven. (b) The magnified image of extruder. Two polystyrene membranes with 200
nm pore size were packed in the middle of the extruder. And the lipid mixture was loaded in the right syringe.
Reused from [67].
The second step was to prepare the ABF suspension. The ABF array was cleaned in a UV/ozone
cleaner for 30 minutes followed by washing with Milli-Q water. The array was then detached
from the original substrate by the sonication in 1 ml HEPES buffer. After that, the ABFs were
pipetted into a new centrifuge tube. The ABFs were collected from the bottom of the tube by
centrifugation (4000 rpm, 3 min) and the volume of the suspension was reduced to 400 l.
The last step was to mix the liposome suspension (100 l) and ABF suspension (400 l)
together and incubate for enough time with gentle rotation to obtain a saturated adsorption of
liposomes on ABF surfaces, so the final concentration of lipids in the mixture was 0.5 mg/ml.
The incubation time was determined by QCM-D data.
QCM-D measurement: Lipids can form a range of different structures on a solid surface
including monolayers, intact vesicles or lipid bilayers [127]. In this study the goal was to coat
ABFs with intact vesicles which allow entrapment of water-soluble drugs inside the vesicles
and/or lipid-soluble drugs within the membrane lipid bilayer. QCM-D is a commonly used tool
to measure adsorption of lipids and their structure on surfaces [128]. Lipids adsorbing to the
surface of a QCM-D crystal cause a drop in the measured resonant frequency and an increase
in the dissipation of the crystal. By monitoring frequency and dissipation changes, the structure,
mass and viscoelastic properties of adsorbed lipids can be determined.
Since the top surface of ABFs is TiO2, a TiO2-coated crystal was used to simulate the
69
adsorption of DPPC liposomes on ABFs. The crystal was treated in the same way as the ABF
array with UV/ozone cleaner for 30 minutes, followed by washing with Milli-Q water. After
the crystal was assembled in the QCM-D (Q-Sense E4, Gothenburg, Sweden) chamber, HEPES
buffer was injected into the cell and left until a stable baseline was observed. The liposome
solution (0.5 mg/ml) was then injected, and the changes of frequency and dissipation were
recorded to monitor the adsorption and stability of lipids on the crystal surface. After the signal
response reached a plateau, the buffer was injected three times to determine the stability of the
adsorbed liposomes. The QCM-D experiment was repeated three times (n=3).
Confocal laser scanning microscope (CLSM): In order to confirm the coating of liposomes
on ABFs, fluorescent probes, calcein or rhodamine B, were incorporated with liposomes.
Calcein (50 mM in HEPES) was entrapped inside the liposomes. Rhodamine B-labeled lipids
were incorporated in the liposome lipid bilayer by adding 2% (w/w) to the DPPC initial lipid
solution. The f-ABFs were centrifuged (4000 rpm, 3 min) and washed at least five times with
HEPES buffer to remove unbound liposomes from the suspension. Images of f-ABFs were
taken by a CLSM (Carl Zeiss AG/LSM 510, equipped with a 40× 0.6 NA objective). The
calcein signal was detected using a 488 nm excitation laser and a 505-550 nm band-pass filter.
For rhodamine B the laser wavelength was 561 nm, and the filter was BP 575-615 IR.
Calcein release measurement: Calcein release measurements were performed using
DPPC/MSPC (9:1 w/w) liposomes, since this lipid combination has been shown to result in a
better triggered-release at 41ºC than pure DPPC liposomes [124]. The calcein release from
DPPC/MSPC functionalized ABFs (f-ABFs) was qualitatively monitored by CLSM. The
fabricated suspensions of f-ABFs were divided equally into three parts and heated at 33ºC,
37ºC and 41ºC for 1 hour. The fluorescence signals from the three samples were recorded by
CLSM using the same parameters. In order to obtain quantitative data, the calcein release was
measured on a Tecan Infinite M200 PRO plate reader by measuring the fluorescence intensity
at the excitation wavelength of 490 nm and emission wavelength of 520 nm. Ti (15 nm) coated
Si wafers 5 mm × 5 mm were used to simulate an ABF surface. The wafers were cleaned in
UV/ozone cleaner for 30 minutes, and then incubated with liposomes for 3 hours, followed by
washing in HEPES buffer 5 times. The wafer was then placed in a 24-well plate containing 1
ml HEPES buffer. The plate was inserted in a plate reader (Tecan Infinite M200 PRO), and the
fluorescent intensities were measured at 33ºC, 35ºC, 37ºC, 39ºC and 41ºC. The temperature
was kept constant under each condition for 1 hour before measuring. The maximum release of
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calcein was determined by adding 2% Triton X-100 in HEPES buffer to dissolve the vesicles
adsorbed on the wafer. The maximum release was used as the positive control, and the wafer
without calcein-loaded liposomes was set as the negative control. The calcein release efficiency
was calculated using Equation (6-1).
Calcein release % = [ ( IDPPC/MSPC ̶ INegative) / (IPositive ̶ INegative ) ] × 100 (6-1)
where IDPPC/MSPC, INegative and IPositive are the fluorescent intensities of the DPPC/MSPC coated
wafer, negative control and positive control, respectively [129].
Results and discussion
The adsorption of liposomes was assessed using QCM-D. Figure 6.11 shows the QCM-D
results of DPPC lipids on TiO2-coated crystals. The data presented were measured at the third
overtone. The frequency decreased 414 12 Hz while the dissipation increased up to 48 3
×10-6 in the first 30 minutes and reached a plateau after 2 hours. This signal is typical of the
adsorption of intact liposomes and is consistent with previously reported results [127, 128,
130]. There was no significant change in the frequency or dissipation after washing the crystal
three times with HEPES buffer, which suggests that DPPC liposomes were stable on the TiO2
surface. To coat the surface of ABFs with DPPC liposomes, we incubated liposomes with
ABFs for three hours to ensure a saturated adsorption.
QCM-D results showed a stable adsorption of DPPC liposomes on the flat surface of the crystal,
these results were confirmed using CLSM. In order to confirm the adsorption of liposomes on
3D-shaped ABFs, rhodamine B labeled lipids (red) were used as a model for a lipid-soluble
drug. The rhodamine B-tagged lipids are embedded within the liposome lipid bilayer. Calcein,
a green dye, was entrapped within liposomes, mimicking a water-soluble drug. Figure 6.12
shows CLSM images of f-ABFs. We used uncoated ABFs as controls to calibrate the intensity
of the laser ensuring that any obtained signal was a result of the fluorescent dye and not caused
by autofluorescence. Strong signals from both rhodamine B (Figure 6.12a) and calcein (Figure
6.12b) show that liposomes were bound to the ABF surface, which confirms the QCM-D data.
This shows that both hydrophobic and hydrophilic drugs can be incorporated in liposomes. The
release of the trapped drugs can be temperature-triggered [131]. Furthermore, the fluorescent
signal may provide a way to track f-ABFs when they are swimming.
Figure 6.13 shows colored-fluorescent images of f-ABFs at 33ºC, 37ºC and 41ºC, respectively.
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The upper three images in Figure 6.13 show the fluorescent images and the lower three images
show the corresponding overlay images of fluorescence and bright fields. From the upper
images, it can be seen that the fluorescent signals from the f-ABFs hardly changed from
Figure 6.11. QCM-D signals of DPPC liposome adsorption on a TiO2 crystal. The blue and red curves are signals
of frequency and dissipation of the crystal, respectively. All data presented were measured at the third overtone.
Reused from [67].
Figure 6.12. Fluorescent images of DPPC-coated ABFs. (a) An f-ABF with rhodamine B labeled liposomes. (b)
An f-ABF with calcein loaded liposomes. Reused from [67].
33ºC to 37ºC. On the other hand, the fluorescent signals on f-ABFs dramatically decreased
from 37ºC to 41ºC and the background signals increased accordingly, which demonstrates that
72
entrapped calcein was released from the DPPC/MSPC liposomes in significant quantities at
41ºC.
Figure 6.13. Calcein release from DPPC/MSPC functionalized ABFs at 33ºC, 37ºC and 41ºC, respectively. The
upper three pictures are fluorescence images, and the lower three pictures are the combined images of fluorescence
and bright fields. Reused from [67].
Figure 6.14a presents the original quantifiable grayscale images of the three fluorescent images
shown in Figure 6.13. The fluorescent intensity of the background on each grayscale image
was measured using ImageJ software. The way to measure the intensity of the background is
described as follows: Ten images without ABFs were created by the segmentation of each
original image and the intensity values were collected from these segmented images. Since we
know that all calcein incorporates in liposomes at 33ºC [124], the intensity of fluorescence on
each image was offset by the one at 33ºC. The results show that the released calcein in the
background gradually increased with a rise in temperature (Figure 6.14b) and the increase from
37ºC to 41ºC was twice more than the one from 33ºC to 37ºC, which indicates that more calcein
molecules were released from liposomes on ABFs when the temperature was increased from
37ºC to 41ºC.
We also measured the fluorescent intensity of the ABFs using ImageJ software and calculated
the ratio of the remaining calcein on ABFs compared to the fluorescence of ABFs at 33ºC
(Figure 6.14c). We noticed that the fluorescent intensity of ABFs at 33ºC was saturated and
therefore, we could not get precise quantitative values. However, the fluorescent intensities at
37ºC and 41ºC are lower than 86% and 32% of the value at 33ºC. The results show that calcein
on ABFs decreased with a rise in temperature, which indicates the release of calcein from
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liposomes on ABFs as the temperature increased. The calcein remaining on f-ABFs, which was
less than 32%, at 41ºC, which means that the calcein release from f-ABFs at 41ºC was more
than 68%.
Figure 6.14. Quantitative analysis of the calcein signals released from ABFs coated with DPPC/MSPC liposomes.
(a) Original fluorescence images of ABFs functionalized with calcein loaded DPPC/MSPC liposomes. (b) The
fluorescent intensity of calcein released from the liposomes. Multiple images without ABFs were created from
each image and analyzed to get the quantitative data (n=10). (c) Ratio of the remaining calcein incorporated within
DPPC/MSPC liposomes on ABFs compared to the total fluorescence at 33ºC. All quantitative analysis was
performed using ImageJ. Reused from [67].
For detecting calcein release from f-ABFs quantitatively and precisely, we need to prepare and
measure large numbers of the f-ABFs capturing calcein. Because it is technically difficult to
obtain large numbers of f-ABFs and because of errors in the number of ABFs in each
experimental set, we used 5mm × 5mm Ti coated Si wafers to simulate the calcein release from
f-ABFs. The surface area of Ti-coated Si (25 mm2) is equal to the total surface areas of more
than 9000 ABFs, which makes the experiment easier and the total release much higher thus
more reliable in quantification. Figure 6.15 shows the calcein release from Ti-coated Si wafers
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measured at 33ºC, 35ºC, 37ºC, 39ºC and 41ºC. The fluorescent intensity increased dramatically
from 37ºC to 39ºC, which shows that calcein started to leak out from liposomes around 39ºC
and continually released till 41ºC. The release efficiency of calcein at 41ºC was 73 ± 15%,
calculated according to Equation (6-1). This quantity was consistent with the amount of calcein
release from f-ABFs at 41ºC (more than 68%).
Figure 6.15. Calcein release from DPPC/MSPC functionalized TiO2-coated surfaces. The “DPPC/MSPC”
represents the sample coated with DPPC/MSPC liposomes. The “Positive control” represents the liposome-coated
sample washed with 2% Triton X-100 in HEPES buffer. The “Negative control” represents the surface without
liposome coating. The error bars represent the standard deviations of the values of four measuring points in the
well. Reused from [67].
6.2.3 Functionalization of ABFs with cationic liposomes
We also functionalized ABFs with another type of liposome, cationic DOPE/DOTAP
liposomes, and demonstrated the single-cell targeted drug delivery to living cells in vitro using
these functionalized ABFs. The 1,2-di-(9Z-octadecenoyl)-sn-glycero-3-phosphoethanolamine
(DOPE) lipids are known to be fusogenic and positively charged 1,2-dioleoyl-3-
trimethylammonium-propane chloride salt (DOTAP) lipids improve liposome uptake by cells
[132-134]. The adsorption of liposomes on the ABFs was confirmed by QCM-D and
fluorescence recovery after photobleaching (FRAP). The functionalized ABFs (f-ABFs) were
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then placed in contact with C2C12 cells in vitro and the uptake of calcein by cells was
monitored using fluorescence microscopy.
Materials and methods
Materials: photoresist IP-L (from Nanoscribe GmbH, Germany), 1,2-dioleoyl-sn-glycero-3-
phosphocholine (DOPC), 1,2-dioleoyl-3-trimethylammonium-propane chloride salt (DOTAP),
1,2-di-(9Z-octadecenoyl)-sn-glycero-3-phosphoethanolamine (DOPE), 1-oleoyl-2-[12-[(7-
nitro-2-1,3-benzoxadiazol-4-yl)amino]lauroyl]-sn-glycero-3-phosphocholine (NBD-PC) were
purchased from Avanti Polar Lipids. Phosphate buffer saline (PBS), fetal bovine serum (FBS),
cell culture media (Dulbecco’s modified eagle’s medium (DMEM)-Glutamax), antibiotic-
antimycotic, trypsin/EDTA were from Invitrogen AG, Basel, Switzerland. Calcein disodium
salt, sodium chloride (NaCl), 4-(2-hydroxyethyl)piperazine-1-ethanesulfonic acid (HEPES),
were purchased from Sigma Aldrich Chemie GmbH, Buchs, Switzerland.
Functionalization of ABFs with liposomes: The process to functionalize ABFs with
DOPE/DOTAP liposomes was similar to the process used for the functionalization of ABFs
with DPPC-based liposomes (Figure 6.9). The unilamellar DOPE/DOTAP (3:1) were prepared
at room temperature by the extrusion method (Figure 6.10). Calcein loaded vesicles were
prepared by dissolving a DOPE/DOTAP (3:1) lipid film in calcein (50 mM in DI-water) before
extrusion. Fluorescein labelled DOPE/DOTAP (3:1) liposomes were prepared by adding NBD-
PC [(2% (w/w) final percentage of the above lipid composition]. Liposomes were prepared at
a final lipid concentration of 2.5 mg/ml and used at 0.5 mg/ml in all experiments. A chip
containing 10000 ABFs (Figure 6.8) was UV/ozone cleaned (30 minutes) to clean the surface
and create free hydroxyl groups on the Ti surface. ABFs were then released by sonication and
incubated for 3 hours under gentle rotation with a liposome solution (0.5 mg/ml) to obtain
functionalized-ABFs (f-ABFs). F-ABFs were washed in HEPES buffer 5 times by centrifuging
the f-ABF containing solution (4000 rpm for 3 minutes each) to remove unbound liposomes.
Equipment: Assessment of liposome adsorption on TiO2-coated crystals was followed using
QCM-D (Q-Sense E4, Gothenburg, Sweden). Imaging of cells and ABFs and FRAP were
acquired using a confocal microscope (Carl Zeiss AG/LSM 510, equipped with a 40× 0.6 NA
objective and a 488 nm argon laser). Time-lapse images of swimming liposome-coated ABFs
were acquired using a Zeiss Axio Observer equipped with a Hamamatsu C9100-13 Super
sensitive fluorescence camera.
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Cell Culture: C2C12 mouse myoblasts were purchased from American Type Culture
Collection (LGC Standards, Molsheim, France) and seeded at 5000 cells/cm2 on glass bottom
24-well-plates in DMEM supplemented with 1% fetal bovine serum and 1% antibiotic-
antimycotic solution. Cultures were continued in a humidified incubator (37 °C, 7% CO2) and
media was changed every 3 days. For calcein delivery experiments, cells were washed with
HEPES buffer followed by 1 hour incubation with the ABF solution then cultured for 24-48
hours with new cell culture media.
Results and discussion
The liposomes were loaded with calcein which is a fluorophore that does not readily penetrate
cell membranes and, therefore, active uptake of calcein-loaded vesicles is required for cells to
show the calcein signal. Calcein is hydrophilic and can thus, be used as a model for a water
soluble drug [135, 136]. In order to confirm that vesicles used in the current study were
adsorbed in an intact state on TiO2 surfaces, we used QCM-D. Figure 6.16a shows the
adsorption of calcein loaded DOPE/DOTAP (3:1) liposomes on TiO2 coated QCM crystals.
The adsorption of liposomes resulted in a final strong decrease in the frequency (Δf = -151.3 ±
19.9 Hz) and an increase in the dissipation (ΔD = 21.9 ± 1.7). The adsorption curve was roughly
saturated after three hours; therefore, the incubation time of the ABFs with liposomes was set
to three hours in all the experiments of this study. TiO2 is known to support the adsorption of
intact vesicles, as shown in previous literature [137]. The typically observed response of the
QCM-D frequency and dissipation signals to the injection of liposomes is in accordance with
previous literature [137]. Other liposome formulations such as DOPC can also be used and the
formation of more than one layer of liposomes to maximize the drug loading capacity is also
possible by alternating positively charged poly-l-lysine layers and negatively charged DOPS
liposomes [138]. QCM-D analysis of the adsorption of the above and other liposome
formulations are shown in Appendix B.
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Figure 6.16. Liposomes adsorb in an intact state to the surface of ABFs. a) Normalized resonance frequency shift
and dissipation shift of the 15 MHz detection frequency for the adsorption of DOPE/DOTAP (3:1) liposomes on
titanium coated quartz crystal microbalance chips. b) A representative image of ABFs coated with liposomes
loaded with 50 mM calcein. c) An ABF coated with NBD-Labeled liposomes before bleaching (left), after
bleaching the center of the swimming microrobot (middle), and 60 min following the bleaching (right). Since no
recovery was observed, it can be concluded that the liposomes are intact on the surface of the helical swimmers.
Reused from [66].
To determine if the liposome-coating process was also applicable on the surface of TiO2-coated
ABFs, calcein loaded vesicles were adsorbed on the surface of ABFs and images of the f-ABFs
were acquired with a CLSM. Because the curvature of the helical rods and their roughness
might affect the adsorption behavior of liposomes on helical microswimmers, it is important to
investigate if adsorption to planar substrates (QCM-D crystal) also applies to the helices. Figure
6.16b depicts a representative CLSM image of ABFs functionalized with calcein-loaded
liposomes where a clear signal can be observed throughout the f-ABF. This is evidence of the
formation of intact liposomes, as calcein would not be present in case of the formation of lipid
bilayers/monolayers on the surface of ABFs. In order to further confirm that the adsorbed
vesicles were intact and did not form lipid bilayers, FRAP was used. Figure 6.16c shows a
representative image of ABFs functionalized with NBD-labeled vesicles before
78
photobleaching, immediately after photobleaching the middle area of the ABFs, and one hour
after photobleaching. In this experiment, the fluorescence is in the bilayer membrane but not
inside the vesicle. When intact vesicles adsorb to a substrate, lipids are not mobile over
micrometer distances, thus fluorescence in a bleached spot does not recover. However,
supported lipid bilayers are mobile, and the photo bleached area recovers its fluorescent signal
within a short period [139]. The fluorescent signal of the photo bleached spot on f-ABFs did
not recover even after 1h of photobleaching (Figure 6.16c); we can thus confirm that liposomes
adsorbed to the ABF surfaces were intact.
In Figure 6.17a, a representative ABF functionalized with calcein-loaded vesicles was
subjected to a rotating magnetic field at a frequency of 8 Hz using a manual magnetic motor.
The resulting swimming speed of the f-ABF was 23 µm/s. The fact that f-ABFs could swim
without significant loss of the calcein fluorescent signal (p=0.2 in Table 1) suggests the ability
of these functionalized microswimmers to deliver their cargo to cells.
Figure 6.17. F-ABF swimming and calcein delivery to single cells. a) A representative time-lapse follow-up of
the swimming of functionalized artificial bacterial flagella (f-ABF) coated with calcein-loaded liposomes. b)
Representative calcein delivery from f-ABF to single cells at low magnification (left), high magnification of
calcein delivery transmission image (middle) and calcein delivered to cells after removing the f-ABF (right).
Reused from [66].
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Table 1. Fluorescent intensity of f-ABFs measured at varying swimming times in Figure
6.17a. One-way ANOVA showed no significant effect of swimming time on fluorescent
intensity (p=0.2). Reused from [66].
Swimming Time (sec) Fluorescent Intensity (a.u.)
0 80.6 86.1 65.7
2 89.0 81.5 68.4
5 58.5 70.0 69.9
10 52.1 69.8 69.7
15 62.1 72.8 71.3
The delivery of calcein to single C2C12 cells in contact with f-ABF is shown in Figure 6.17b
and Figure 6.18. It is noteworthy that even when two cells were in direct contact, the calcein
was only delivered to the cell in contact with the ABF and not the other. Additionally, nearby
cells showed no calcein signal (Figure 6.17b). When the ABF was removed from the surface
of the cell using an external magnetic source, the calcein signal was still clearly present. This
data indicates that f-ABFs can perform precise swimming and deliver the loaded model drugs
into single cells. The delivery of calcein might occur through fusion of the cationic vesicles
with the membrane of the cultured cells or through endocytosis [140, 141].
Figure 6.18. Calcein delivery to single cells in vitro. Merged fluorescent/transmission image (left), fluorescent
(middle) and transmission (right) of the calcein delivery to a single C2C12 cell using the DOPE/DOTAP liposome
formulation. Reused from [66].
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6.3 ABFs functionalized with lipoplexes for gene delivery in vitro
6.3.1 Introduction
In medicine, gene therapy is the medical treatment using DNA as a therapeutic drug by
delivering the DNA into a patient’s cells to treat diseases such as inherited disorders and
cancers. During the process the nucleic acids must be delivered to the defective cells, be
transfected into the cells and express the function which can cure diseases. Gene delivery
carriers, commonly called vectors, have been developed to carry DNA to improve gene therapy
due to the poor efficiency of naked DNA entering into cells. Lipoplexes, complexes of cationic
lipids and DNA, are promising tools for nucleic acid delivery to cells such as siRNA for gene
silencing and plasmid DNA (pDNA) transfection [142-145]. The delivery of nucleic acids to
targeted cells and tissues is still a challenge [142]. The combination of lipoplexes with mobile
microrobots as lipoplex carriers, such as ABFs which have the ability of controlled 3D
navigation in liquid, may create a multifunctional system which can be wirelessly controlled
and targeted delivery of the DNA to specific areas in the hard-to-reach areas, such as in human
bodies and in lab-on-a-chip environments.
In this section, we demonstrate, for the first time, the successful wirelessly targeted and single-
cell gene delivery to human embryonic kidney (HEK 293) cells using ABFs loaded with a
lipoplex in vitro. As a result of development of the reverse transfection technology [146], the
lipoplex was successfully bound to ABFs and was only released into cells by contact between
ABFs and cells. Plasmid DNA (pDNA) was first mixed and complexed with lipofectamine
2000, a cationic lipid, to form the lipoplex, and the functionalization of ABFs with the lipoplex
was characterized by fluorescent probe method. The functionalized ABFs (f-ABFs) were
steered and controlled wirelessly by low-strength rotating magnetic fields and the loaded
pDNA was delivered into targeted cells. The lipoplex containing pDNA carried by f-ABFs was
only taken by the targeted cells. The successful gene transfection and gene expression to
encoding proteins by the targeted cells was verified.
6.3.2 Functionalization of ABFs with lipoplexes
Materials and Methods
Materials: Lipofectamine 2000 (Life technologies, Carlsbad, CA), bovine fibronectin (Life
Laboratory Company, Yamagata, Japan), gelatin (Sigma-Aldrich, St. Louis, Madison, MO),
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Label IT® Plasmid Delivery Control (Fluorescein; Mirus Bio LLC, WI), HEK 293 cells
(American Type Culture Collection, Manassas, VA), fetal bovine serum (FBS; Life
technologies), Dulbecco’s modified eagle's medium (DMEM; Life technologies), DMEM
without phenol red (Life technologies), Antibiotic-antimycotic (100×; Life technologies),
HEPES sodium buffer (10 mM 4-(2-hydroxyethyl)piperazine-1-ethane-sulfonic acid and 150
mM NaCl in distilled water; MicroSelect, Fluka Chemie GmbH, Switzerland), DPBS
(Dulbecco’s phosphate-buffered saline; Life technologies), trypsin-EDTA (0.05%; Life
technologies) and sodium chloride (NaCl; Sigma-Aldrich) were purchased. The pDNA
encoding Venus protein was a gift from Advanced Industrial Science and Technology, Ibaraki,
Japan.
Fabrication of ABFs: The helical bodies were fabricated in IP-L photoresist using DLW
(Nanoscribe, from Nanoscribe GmbH) on a transparent glass substrate, followed by coating
with 25 nm Ni and 15 nm Ti using an electron beam evaporator (Plassys-II MEB550SL).
The preparation of the final lipoplex: Lipofectamine 2000 was used as cationic lipids to mix
with pDNA in this work. Lipofectamine 2000 is the commercially available cationic reagent
that provides highly efficient transfection of nucleic acids (DNAs or RNAs) into a wide range
of mammalian cells [147]. While the detailed chemical composition of the reagent is not
disclosed by the provider, the mixing of the cationic lipids of lipofectamine 2000 and the
anionic nucleic acids generates cationic lipoplexes through ionic interactions. The lipoplexes
easily fuse with the negatively-charged cell membrane or are incorporated into cells by
endocytosis, which facilitate the transport of DNA through the cell membrane and also
protecting the DNA from undesirable degradation [148, 149].
Cell adhesive proteins, fibronectin and gelatin, were added to the mixture of lipofectamine
2000 and pDNA to form the final lipoplex since previous results showed that these proteins are
very important for local and highly efficient transfection of lipoplexes in the reverse
transfection technology [146]. Fibronectin, which is an extracellular matrix component that
binds to integrins on cell membranes, gives cells scaffold and makes strong cellular adhesion
on solid surfaces, such as glass and metal etc. [146, 150]. Strong cellular adhesion on solid
surfaces promotes gene transfection by a contact between lipoplexes and cellular membrane.
Moreover, gelatin adsorbs aqueous mixtures including lipoplexes and suppresses their release
from solid surfaces [151].
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The detailed prepared steps were described as follows. First, 2 μl of a solution of pDNA (1
μg/μl; Fluorescein labeled one or Venus encoding one), 4 μl of lipofectamine 2000 and 31.5 μl
of DMEM were mixed and incubated for 20 min at room temperature. Second, 2.5 μl of 2%
(w/v) gelatin in distilled water (DW) and 10 μl of 0.4% (w/v) fibronectin in DW were added in
the mixture. The components of the final lipoplex are shown in Table 2.
Table 2. The components of the final lipoplex (50 µl in total). Adapted from [68].
Materials Amount
Lipofectamine 2000 4 µl
pDNA (pVenus-N1) (1µg/µl in DW) 2 µl
DMEM without phenol red 31.5 µl
Gelatin (20µg/µl in DW) 2.5 µl
Fibronectin (4µg/µl in PBS) 10 µl
QCM-D experiments: The TiO2-coated crystal (Q-Sense, Sweden) was treated with
UV/ozone cleaner for 30 min and loaded into the QCM-D chamber (Q-Sense E4, Gothenburg,
Sweden). HEPES sodium buffer was injected into the chamber until a stable baseline was
observed. After that, 500 l of the lipoplex (prepared as indicated in Table 2), was loaded in
the chamber with a flow rate of 0.2 ml/min. After 13h of incubation, the substrate was washed
by HEPES sodium buffer with the speed of 0.1 ml/min.
Functionalization of ABFs with lipoplexes: The fabricated ABFs (Figure 6.8) were cleaned
using UV/ozone cleaner for 30 minutes and released from the glass substrate in HEPES sodium
buffer by sonication. The ABF suspension was condensed to 50 l by centrifuging. The
lipoplex solution (50 l) and the ABF suspension were mixed and incubated at room
temperature for 3h. In order to remove the uncoated lipoplex in the solution, the mixture was
washed 3 times in 1 ml of HEPES sodium buffer and 1 time in 1 ml of DMEM before
centrifuging the solution (4,000 rpm, 3 minutes each time). The final f-ABF suspension in
DMEM was condensed to 100 l by centrifuging for further usage.
Cell culture: The DMEM supplemented with 10% fetal bovine serum and 1 × antibiotic-
antimycotic was used as cell medium. We suspended 2 × 105 cells of HEK 293 in 500 μl of
medium for seeding in 1 well of a 24-well culture dish with a glass bottom for the experiments
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of gene expression or on glass substrate for the swimming experiments.
Swimming experiments: The swimming tests were conducted using rotating magnetic fields.
The rotating magnetic fields are generated by three-orthogonal-pair Helmholtz coil setup
(Figure 4.2a). The glass substrate with HEK 293 cells was placed in a tank (3 cm height, 1.5
cm width and 3 mm height), which was filled with cell medium. The tank was placed in the
center of the coil setup. A microscope with a camera was mounted above the tank. The
swimming experiments with cells were finished within 1h after the cells were taken out from
an incubator. After f-ABFs targeted and made contact with cells, the cells were immediately
put back into the incubator.
Gene expression: The f-ABF suspension (50 µl) was added into HEK 293 cells medium, 500
µl of DMEM supplemented with 10% fetal bovine serum and 1 × antibiotic-antimycotic. The
cells were incubated in CO2 incubator at 37ºC for 16-48h for gene transfection and protein
expression. The fluorescent images were acquired using a CLSM (Carl Zeiss AG/ LSM 510
with 10×, 20×, 40× 0.6 NA objectives and a 488 nm argon laser).
Results and discussion
A flat TiO2-coated surface was used to model the outer TiO2 surface of ABFs, since the Ti was
previously oxidized to TiO2 during the UV/ozone treatment. The QCM-D results show that the
frequency shift dropped and the dissipation shift increased dramatically once the lipoplex was
injected into the chamber, which indicated good adsorption of the lipoplex onto the TiO2
surface (Figure 6.19). The adsorption reached a maximum after incubation for 3h. When
HEPES sodium buffer was loaded onto the substrate for washing after lipoplex was adsorbed
to the substrate for 18h, the frequency shift decreased and the dissipation shift increased. The
decrease in the frequency may come from the continuous adsorption of the substances
remaining in the pump tube which were pumped into the chamber when starting the washing.
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Figure 6.19. QCM-D measurement of the adsorption of the lipoplex on a TiO2 crystal. The HEPES sodium buffer
was used for washing. Only the third overtones are shown for the sake of clarity. Reused from [68].
Figure 6.20. Functionalization of ABFs with the lipoplex. (a) The schematic of the functionalization of ABFs with
the lipoplex. (b) The CLSM image of a f-ABF. The pDNA was marked using green fluorescence (fluorescein).
Reused from [68].
The functionalization of ABFs was performed by mixing the suspension of ABFs with the
lipoplex and the mixture was incubated at room temperature for 3h followed by washing to
generate functionalized ABFs (f-ABFs) (Figure 6.20a). In order to confirm the successful
loading of pDNA on the real ABF surface, we used pDNA labeled with fluorescein, a green
fluorescent (Label IT® Plasmid Delivery Control, Fluorescein; Mirus Bio LLC, Madison, WI)
as a fluorescent probe. The fluorescent signals were verified by confocal CLSM after coating
the lipoplex onto ABFs. The green fluorescent signal detected around the ABF indicates the
lipoplex loaded with pDNA was successfully coated and loaded on the ABF surface by the
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electrostatic force between the negatively charged TiO2 surface of ABFs and the positively
charged lipoplex (Figure 6.20) [149] .
6.3.3 Swimming performance of f-ABFs
Figure 6.21a shows the wirelessly controlled swimming performance of f-ABFs in cell medium
in a 5mT rotating magnetic field. The forward swimming speed of f-ABFs increased almost
linearly as the rotating frequency of the magnetic fields increased and then reached the peak at
the step-out frequency. When the input frequency became higher than the step-out frequency
of the swimmers, the swimming speed of the f-ABFs decreased dramatically. The average
maximum forward speed of three tested f-ABFs (F-ABF1, F-ABF2 and F-ABF3) was 43.9 ±
1.3 m/s, and the average step-out frequency was 31 ± 3 Hz. Figure 6.21b shows the wireless
targeting ability of an f-ABF in vitro. The f-ABF was controlled to swim towards the target
cell (cell 3) and made contact with the cell (the insert of Figure 6.21b) by wireless means.
Figure 6.21. Swimming performance of f-ABFs in cell medium. (a) The forward swimming speeds of three tested
f-ABFs (F-ABF1, F-ABF2 and F-ABF3) as a function of input rotating frequency at 5mT magnetic fields. (b)
Time-lapse photo of the controlled actuation of an f-ABF to the target cell (cell 3, the contour is the circled in
red). Cell 3 was under cell division and was considered as one single cell in this case. Cell 2 and cell 4 are blurred
in the image due to the drift of cells in the medium during the time-lapse image. The movement of the f-ABF was
marked with blue rectangles. The interval of each movement was four seconds. The insert shows an f-ABF in
contact with cell 3. Reused from [68].
6.3.4 Gene transfection and protein expression
The transfection efficiency of the lipoplex was tested using HEK 293 cells. Two different
amounts of lipoplexes including pDNA encoding yellow-green fluorescent Venus protein [0.2
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µg of pDNA (Figure 6.22a) and 1 µg of pDNA (Figure 6.22b)] were prepared. Then each
lipoplex was added to 500 µl of DMEM with HEK 293 cells in a 24-well culture plate for gene
transfection. The results show that a 1 µg amount of pDNA as a component of the lipoplex is
enough to transfect most cells in 500 µl cell medium in 23h (Figure 6.22b).
Figure 6.22. The transfection efficiency tests of the lipoplex after incubation for 23h with two different pDNA
amounts. (a) 0.2 g pDNA was added into the cell medium. (b) 1 g pDNA was added into the cell medium.
Reused from [68].
Gene transfection of f-ABFs was then conducted. The f-ABF suspension was added to 500 l
of DMEM with the HEK 293 cells in one well of a glass-bottom 24-well plate. The f-ABFs
deposited on HEK 293 cells on the bottom of the plate. After 16h of incubation, the two cells
contacting and/or neighboring with one f-ABF became transfected with pDNA and presented
a strong fluorescent signal from Venus proteins (Figure 6.23 and Figure 6.24). The two
transfected cells were daughter cells divided from the parent cell which had had contact with
and been transfected with pDNA from the f-ABF, as indicated by the morphologies of two cells
expressing Venus protein which migrated separately after the division (Figure 6.23e). The
successful protein expression indicates that the pDNA carried by ABFs was transferred into
the membrane of the parent cells, imported into the nuclei through the temporary broken
nuclear membrane during the cell division and Venus proteins were expressed in the daughter
cells.
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Figure 6.23. pDNA transfection and protein expression by the cells that were in contact with f-ABFs. (a,b) and
(c,d) show the fluorescent and transmission images of cells with the f-ABF at 10× and 40× magnifications,
respectively. (e) The zoomed overlay of fluorescent and transmission images. The f-ABF (the red helix) is false-
colored in red, processed using ImageJ software. Reused from [68].
Figure 6.24. Gene transfection and protein expression in the cells contacting f-ABFs. (a,d), (b,e) and (e,f) are
fluorescent, transmission and overlay images, respectively. (a-c) show the cells incubated with f-ABFs for 16h
and (d-f) show the cells incubated with f-ABFs for 48h. Reused from [68].
The absence of fluorescent signal from cells surrounding the daughter cells shows that the
transfection from f-ABFs was locally limited to the cells which had been in contact (Figure
88
6.23a-d). Figure 6.24 shows the two daughter cells migrated and were separated by a greater
distance after 48 hours of incubation (Figure 6.24d-f) than after 16 hours (Figure 6.23 and
Figure 6.24a-c), which indicates that the f-ABF was able to transfect the contact cells and did
not influence the cell division and migration. The cells underneath or surrounding f-ABFs seem
healthy from the transmission images (Figure 6.24a,d) and the f-ABFs seem nontoxic to human
embryonic kidney (HEK 293) cells after two days.
6.4 ABFs functionalized with near-infrared dyes for in vivo
tracking and actuation
6.4.1 Introduction
In our previous work, the functionalization of ABF with lipid-based drug carriers containing
fluophores, drug models and DNA was achieved and a cargo could be delivered in vitro to cells
[66-68] to demonstrate the possibility of using ABFs as active drug delivery devices. The f-
ABFs were tracked/localized by normal optical microscopes or CLSM, however, for in vivo
applications the ABFs should be tracked and the moving direction should be controlled. In this
work, we describe the surface functionalization of ABFs with near-infrared probes (NIR-797)
that allowed whole-body optical (fluorescence) imaging to track for the first time in vivo the
magnetically controlled navigation of a swarm of functionalized ABFs (f-ABFs) in the
peritoneal cavity of a mouse. This is a collaborative work between our lab and Nanomedicine
lab in University of Manchester, UK. The functionalization of ABFs with near-infrared dye
was mainly finished in the collaborator’s lab by Dr. Ania Servant. The results were recently
accepted by the journal of Advanced Materials [152].
6.4.2 Functionalization with near-infrared dyes
Materials and methods
Materials: All the chemicals were ordered from Sigma Aldrich.
Fabrication of ABFs: The polymeric bodies of ABFs were fabricated using IP-L with DLW
and were coated with 50 nm Ni and 5 nm Ti by e-beam evaporation (Plassys-II MEB550SL)
with a rotation speed of 4 rpm and 15° tilt angle.
Actuation and tracking of f-ABFs in vitro: The three pairs of Helmholtz coils were used to
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control the movements of the f-ABFs (Figure 6.25A and Figure 4.2a). For the optical tracking,
the f-ABF suspension was contained in a micro-well (1.8 mm diameter, 2 mm height) (Figure
6.25B). The microchip was made from PDMS on a glass slide, and it was obtained from Yun
Ding in the deMello group at ETH Zurich. An optical microscope equipped with 10×, 20× and
50× objectives and a CCD camera were placed on the top of the well.
Figure 6.25. Actuation of a swarm of ABFs in vitro tracked by an optical microscope. (A) The three pairs of
Helmholtz coils used for applying rotating magnetic fields. The suspension of ABFs was placed in the center of
the three pairs. (B) The suspension was injected into a micro-well (1.8 mm in diameter and 2 mm in height) on a
microchip made from PDMS. Reused from [152].
In addition, a small fraction of residual impurities from the coated glass substrate,
functionalized with NIR-797 dyes, generated during the untethering of the f-ABFs, were also
present in the f-ABF solution. Such impurities acted as magnetic rods rotating with the
magnetic field, however, moving perpendicularly to the swimming direction of f-ABFs. This
induced some degree of inaccuracy in the calculation of the cloud centroid location and should
be further improved.
A
B
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ABF functionalization with FITC and NIR-797 fluorescent dyes: The surfaces of ABFs
were oxidized before untethering them from the glass slide substrate using an anodic oxidation
process according to published procedures [153, 154]. Briefly, the Ni/Ti coated glass slide was
cleaned with acetone and ethanol and dried. The sample was then immersed in an electrolyte
solution prepared by dissolving 10g of NaNO3 in 1L of methanol containing less than 1% of
water. The formation of a uniform TiO2 oxide layer was obtained by anodization (the cathode
being a graphite rod) after 4h of applying a current density of 0.03 mA/mm2 at room
temperature and was characterized by SEM. Following the anodic oxidation, the sample was
left in the oven at 200°C for one hour. The thickness of the surface oxide layer was
characterized using SEM and the composition profile was characterized by energy-dispersive
X-ray spectrometry (EDS). The presence of nickel was only detectable on the Ni/Ti coated
substrate and on the anodized samples confirming the successful Ni/Ti coating of the glass
substrate. The sample surface was then cleaned with acetone and ethanol, followed by
treatment with concentrated NaOH at 60 °C for 15 min in order to insure a high concentration
of hydroxyl groups on the surface. After rinsing the sample surface with distilled water, the
sample was treated with a 5% (w/w) 3-aminopropyltriethoxysilane (3-APS) solution in
Ethanol/Water (95/5 vol%) at pH 4 for 15 min and then cured at 110°C for 40min in order to
cross-link the silane film. Aminosilanes covalently attached to the oxide layer of the anodized
ABF surface by a condensation reaction with the hydroxyl groups on the anodized substrate
that were obtained by treatment with concentrated NaOH. The derivatization of the hydroxyl
groups onto amino groups was characterized using the Kaiser test. This method uses the
properties of the ninhydrin molecule to change color on reaction with primary amines and is
widely used for characterization of amino-functionalized nanomaterials [155]. The
determination of NH2 group concentration was conducted using the following equation:
(6-2)
where C is the amino group concentration, A570 the absorbance at 570 nm, and ε the ninhydrin
extinction coefficient and l the length of the quartz cuvette (1 cm).
The functionalization of the sample surface with NIR-797 was conducted by reacting NIR-797
with the amino groups introduced on the sample surface according to published procedures
[156]. In brief, the sample was immersed in 4 ml of a 1mg/ml solution of FITC/ NIR-797 in
chloroform/methanol (50/50 vol%), 15 µl of trimethylamine solution was added to the reaction
l
dilutionAlmolC
Ninhydrin
nmNH
570
2 /
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and the reaction was allowed to proceed for 48 hours at room temperature under constant
shaking. After functionalization with FITC/NIR-797, the ABFs were untethered from the
surface by sonication.
NIR-797 functionalized ABF (f-ABFs) characterization: The fluorescent signal of the f-
ABFs was characterized using a LS-50B fluorimeter (PerkinElmer) and a correlation between
the number of f-ABFs and the fluorescence signal could be determined. The functionalized
ABF surfaces were characterized using an epifluorescent microscope equipped with 10×, 20×
and 40× objectives.
In vitro and in vivo live imaging (IVIS) of the f-ABFs: The fluorescence signal of the f-ABFs
was characterized in vitro and in vivo using the IVIS Lumina III and the living Image software
(Xenogen, Caliper Life Sciences). Imaging parameters: λbkg= 570 nm, λexc=745 nm, λem= 810
nm~875 nm, exposure time 6s and f/stop 2. Different solutions of f-ABFs dispersed in
isopropyl alcohol (IPA) with decreasing concentrations were prepared and analyzed in vitro in
a quartz cell in order to determine a calibration curve. For the IVIS tracking, a suspension of
120, 000 f-ABFs in 1ml of IPA was placed in a rectangular quartz vial (1 cm × 1 cm) in the
center of the three pairs of coils. The coils were placed inside of the chamber of the IVIS under
a camera (Figure 6.26A). The f-ABFs were placed in the center of the coils, and the rotating
magnetic fields were controlled by changing the input current using a current controller.
For the in vivo experiments, all experiments were performed in accordance with the approved
recommendations and policies of the UK Home Office (Animal Scientific Procedures Act
1986, UK). Balb-C mice (4 weeks old, Harlan, UK) were used for the in vivo experiments.
Anesthesia was induced by inhalation of 2% isoflurane. Animals were administered in the
intra-peritoneal cavity with a dispersion of f-ABFs (80000 f-ABFs in 400 µl) in 5% dextrose;
the volume was adjusted according to the animal body weight to a maximum of 20 mL/kg as
recommended on the LASA (Laboratory Animal Science Association, UK) good practice
guidelines. The anesthetized animal was placed in the center of the coils (Figure 6.26A) and
the rotating magnetic fields were controlled by changing the input currents using a current
controller. The IVIS image of a mouse injected with 5% dextrose solution served as control
(Figure 6.26B). Images of f-ABFs before (Figure 6.26C) and every minute after actuation by
the magnetic field, were recorded at the settings described above.
Image processing: The contours and the centroids of the clouds were measured using Icy
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1.4.3.5 version, an open source community software for bio-imaging
(http://icy.bioimageanalysis.com). The contours and the centroids of the cloud in Figure 6.31
and Figure 6.33 were analyzed by ‘Active Contour’ function. The contours in Figure 6.34 and
Figure 6.35 were measured by a custom-made Matlab code.
Figure 6.26. Actuation and in vivo tracking of a swarm of f-ABFs. (A) Experimental setup for actuation and in
vivo tracking of f-ABFs. The right insert shows the three pairs of Helmholtz coils with an anesthetized 4 week old
Balb-C mouse. (B) The IVIS image of a mouse injected with 5% dextrose solution, served as control. (C) The
IVIS image of a mouse injected with about 100,000 f-ABFs in 5% dextrose. Reused from [152].
Results and discussion
Figure 6.27A shows the fabricated ABFs. The helical bodies of ABFs could be obtained in
different sizes, such as 8 µm (Figure 6.27Aa) and 16 µm (Figure 6.27Ab) in length. Following
DLW, the polymeric bodies were coated with a 50 nm-thick layer of Ni and a 5 nm layer of Ti.
Magnetic material such as Ni enabled us to wirelessly control ABF, and Ti was used to improve
Coils
Mouse
A
42 cm
37 cm
30 cm
Height
adjustable
platform
Camera
B C
5 mm
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the biocompatibility of the devices [157, 158]. A large number of ABFs (10,080 in each batch)
was produced on a glass slide, which served as a substrate, and subsequently coated with a
Ni/Ti layer (Figure 6.27Ac-d).
Figure 6.27. Fabrication and functionalization of ABFs with NIR-797 dyes (f-ABFs). (A) ABFs of different
lengths (8 m (a) and 16 m (b)) were fabricated by DLW, scale bars 4 m. (c) One array of ABFs with 16 m
length (10,080 ABFs) was produced on the glass substrate (dark square inside the red box). (d) SEM image of
part of the arrays of ABFs. (B) Functionalization of ABFs with NIR-797 dyes. After fabrication (i), the Ni/Ti
surfaces of ABFs were anodized (ii) and covalently coupled with NIR-797 dyes (iii). After functionalization, the
f-ABFs were detached from the substrate by sonication (iv). (C) The fluorescent image of f-ABFs after
functionalization. (D) The f-ABF suspension in a centrifuge tube (left) and near (right) a permanent magnet (500
mT). Reused from [152].
Fluorescence-based in vivo imaging in the near infrared (NIR) spectral region has proven to be
a beneficial tool for the development of novel drug delivery vectors and in particular for
monitoring drug targeting. This technique offers the possibility of generating data with a high
signal-to-noise ratio, and is able to achieve deep tissue penetration as a result of the low photon
OHOH OH
OHOH
O
A
B
C D
OHOH OHOHOH
O
SubstrateSwimmer
(a)
(c)
(d)
(b)
OH OH OHOHOH
OH
20 m
i
ii iii
iv
20 m
6 mm
94
adsorption of endogenous biomolecules in the range of 650 to 1000 nm wavelength [159]. NIR-
797 was selected as a NIR emitting fluorophore in order to monitor the magnetically controlled
in vivo navigation of the ABFs using whole-body optical (fluorescence) imaging (IVIS system).
The ABFs were functionalized with NIR-797 prior to untethering from the glass substrate in a
three step procedure: i) formation of oxide layer; ii) derivatization of the hydroxyl groups into
amino groups; and iii) coupling with NIR-797 molecule via the reaction between
isothiocyanate groups present in NIR-797 molecule and the amino groups on the ABF surface
(Figure 6.27B-C).
The first step, the formation of a titanium oxide layer, was performed by the anodization
process. This method has commonly been used to produce a Ti-rich oxide layer on surfaces of
super-elastic NiTi alloys for biomedical applications in order to improve the corrosion
resistance of NiTi-coated materials in physiological environments [160]. In our case, the
formation of an oxide layer by anodization was used for the introduction of hydroxyl groups to
facilitate the coupling of the NIR-797 molecules onto ABF surfaces while leaving the Ni atoms,
responsible for the magnetic properties of the swimmers, unaffected. A cross section of the
Ni/Ti coated glass slide before and after the anodization process is shown in Figure 6.28A. The
thickness of the titanium oxide layer could be roughly estimated at 120 ± 30 nm by measuring
the difference in thickness before and after anodization with Image J using 4 different samples.
The atom composition of the ABF surface substrate before and after anodization was
characterized using energy-dispersive X-ray spectrometry (EDS) (Figure 6.28B) using an
uncoated glass slide as a control. The data confirmed the presence of Ni/Ti atoms on the coated
substrates and on the anodized surfaces. The presence of oxygen, aluminum, sodium,
potassium and silane in high proportion could be observed on all samples. These atoms are
characteristic of atoms present on a glass slide. The atomic percentage content of oxygen (O),
titanium (Ti) and nickel (Ni) atoms on the surface of the different samples, normalized to the
control glass substrate, are shown in Figure 6.28C and indicate that oxygen content
significantly increased after the anodization step, while the content of titanium and nickel
remained stable, confirming the formation of the TiO2 layer on the coated glass substrate.
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Figure 6.28. The f-ABFs characterization. (A) Thickness of the TiO2 layer after anodization of a Ni/Ti coated
array of ABFs. (i) A cross-section of a Ni/Ti coated glass substrate. Thickness of a Ni/Ti coated glass substrate
(ii) before anodization and (iii) after anodization. (B) EDS Spectra of (i) glass substrate, (ii) Ni/Ti coated glass
substrate, and (iii) anodized Ni/Ti coated glass substrate. The red arrows highlight the presence of Ti and Ni on
the coated glass substrate. (C) The comparison of elements (O, Ti and Ni) before and after anodization. (D) (i)
ABFs and (ii) optical image of an untethered ABF before functionalization, (iii) epifluorescent images of FITC-
functionalized ABFs, and (iv) untethered FITC-functionalized ABF. (E) Characterization of f-ABF fluorescence:
(i) fluorescence signal at 820 nm of f-ABFs dispersed in IPA at increasing concentrations (10 000, 20 000, 30 000,
60 000, and 100 000 f-ABFs/ml) (λexc: 795 nm); (ii) Calibration curve of f-ABFs dispersed in IPA. The volume
of the tested sample was 10 ml. (F) IVIS fluorescence signal of the f-ABFs at different concentrations in IPA (60
000, 30 000, 15 000 f-ABFs/ml). Reused from [152].
96
The addition of reactive amine groups onto the coated ABF surface for the subsequent reaction
with the isothiocyanate NIR-797 was performed by the introduction of an aminosilane layer.
This method was previously reported by Bakhshi et al., in order to coat a polyhedral oligomeric
silsesquioxane (POSS)-nanocomposite polymer onto NiTi surfaces [153]. The
functionalization of the ABF surface with the aminosilane was found to be highly reproducible
and the loading of primary amine groups was around 0.94 nmol/cm2 (standard deviation 0.08
nmol/cm2) and considering the surface area of a single ABF (251.33 µm2), the functionalization
was found to be around 2.36 × 10-7 nmol/ABF (1 amine group per 17.7 Å2) (Figure 6.29).
Figure 6.29. Derivatization of the hydroxyl groups into primary amine groups by Kaiser test. (A) UV-Visible
spectra of NH2-modified ABF substrates following reaction with the amino silane onto the oxidized substrate. A
Ni/Ti coated sample that was anodized was used as a control. The presence of primary amine groups is
characterized by the presence of a peak at 570 nm. (B) Determination of the NH2 functionalization onto the
substrate surface by the Kaiser tests. The quantification of the primary amine groups is determined by the
maximum absorbance of the peak at 570 nm. Reused from [152].
0
5
10
15
20
25
sample 1 sample 2 sample 3 sample 4 sample 5
NH
2q
uan
tity
(n
mo
l/g)
0
0.1
0.2
0.3
0.4
0.5
0.6
0.7
0.8
0.9
1
300 400 500 600 700 800
Ab
sorb
ance
Wavelength (nm)
Control
sample 1
sample 2
sample 3
sample 4
A
B
97
Following this step, coupling with NIR-797 isothiocyanate was performed at room temperature
for 2 days in chloroform/methanol (50/50 vol%). A control reaction was conducted under the
same conditions using a fluorescent isothiocyanate (FITC) (λexc 488 nm and λem 568 nm) that
allowed the observation of a fluorescently labelled ABF under an epifluorescent microscope.
The resulting FITC-functionalized ABFs surfaces were characterized by fluorescence
microscopy and the images are shown in Figure 6.28D-iii. The data demonstrated a
homogenous fluorescence labelling of the functionalized ABF surfaces. The fluorescent signal
of the dispersed FITC/NIR-797 functionalized ABFs was characterized by spectrofluorimetry
and fluorescence microscopy (Figure 6.28D-iv). In order to quantify the reaction yield of the
coupling with NIR-797, a calibration curve was generated using different f-ABF concentrations
in isopropyl alcohol (IPA) (Figure 6.28E). The amount of NIR-797 on each f-ABF could be
estimated and was found to be 1.09×10-5 nmol per f-ABF. The yield of the NIR-797/NH2-
functionalized ABF coupling reaction was therefore around 5%. This low yield was explained
by the fact the NIR 797 is a bulky molecule with limited accessibility to the amino groups
located on the surface of the ABF.
The fluorescent signal of the NIR 797/ABF coupling was then studied in aqueous and
physiological media such as HEPES buffer (25mM, pH 7.4), 5% dextrose and mouse serum
(Figure 6.30A). The intensity of the fluorescent signal of the f-ABFs varied in the different
solvent systems (Figure 6.30B), this was consistent with the extinction coefficient of NIR-797
in the same solvents. NIR-797 extinction coefficient was found to be higher in mouse serum
than in all the other solvents. Mouse serum appeared to be the best solvent for an optimal NIR-
797 fluorescent signal. The stability of the fluorescent signal of the f-ABFs in 5% dextrose and
50% mouse serum was also studied overtime (Figure 6.30C). The fluorescent signal remained
unchanged over a period of 2 weeks in all media used.
In order to investigate the swimming ability of the f-ABFs, the intensity of the NIR fluorescent
signal was determined using the IVIS optical imaging system at an excitation wavelength of
745 nm, at which the adsorption of endogenous biomolecules is expected to be low. The f-
ABFs dispersed in 5% dextrose at different concentrations (60,000; 30,000; 15,000 f-ABFs/ml)
were prepared and the fluorescent signal was investigated. The results are shown in Figure
6.28F. The fluorescent signal decreased when the number of f-ABFs was decreased, and a
fluorescent signal could be detected with a concentration of f-ABFs as low as 15,000 f-
ABFs/ml.
98
Figure 6.30. Fluorescence signals of f-ABFs in different solvents and signal stability. (A) Fluorescence signal of
f-ABFs (120 000 f-ABFs/ml) in HEPES buffer (25 mM, pH 7.4), 5% dextrose, IPA and 50% mouse serum; (B)
Extinction coefficient of NIR-797 in different solvent systems; (C) Fluorescence signal stability in 5% dextrose
and 50% mouse serum; Fluorescence signal of 120 000 f-ABFs/ml dispersed in 5% dextrose and 50% mouse
serum over a period of 2 weeks. Reused from [152].
0
5
10
15
20
25
30
35
40
45
50
800 850 900
Flu
ore
scen
ce in
ten
sity
(a.
u.)
Wavelength (nm)
5% dextrose
50% mouse serum
IPA
HEPES buffer
y = 0.0072xR² = 0.9708
y = 0.1832xR² = 0.9868
y = 0.1383xR² = 0.9813
y = 0.476xR² = 0.9947
0
50
100
150
200
250
300
350
0 1000 2000 3000
Flu
ore
scen
ce in
ten
sity
(a.
u)
Concentration (nM)
HEPES buffer
IPA
5% dextrose
50% mouse serum
0
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t In
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50% mouse serum
5% dextrose
A
B
C
99
6.4.3 In vitro tracking and actuation of f-ABFs
The f-ABF in vitro swimming capabilities by wireless magnetic actuation were then
investigated. Figure 6.31A shows the controlled motion of an f-ABF swarm (of around 20, 000
ABFs) observed by optical microscopy using three pairs of Helmholtz coils that provided a
rotational magnetic field (Figure 6.25). The application of a rotational magnetic field allowed
the f-ABFs to translate a rotational movement into propulsion [6]. Since the intensity of the
magnetic field inside the coil was uniform, the movement of the whole swarm of f-ABFs could
be precisely controlled. Depending on the input magnetic fields, ABF can assemble and
disassemble. The assembly between two ABFs can however alter the swimming speed,
depending on the angle of assembly. Nevertheless, when this event occurs, it was previously
demonstrated in our group that the assembled ABFs are still able move in a same direction, and
swimming can still be controlled [161].
Altering the rotating magnetic field parameters such as field strength and frequency allowed
the tuning and control the swimming direction and speed of the swarm. Yaw and pitch were
the steering parameters in order to control the orientation of the ABF in the horizontal plane
and out-of-plane, respectively. In addition, the f-ABFs are able to swim with six degree of
freedoms (left, right, forward, backward, up, down and free rotation about three perpendicular
axes). When ABFs swim on an in-plane surface, the ABFs have both forward and drift speeds
due to the drag force imbalance of the wall effect [6]. In order to reduce the drift speed, the
swimming experiments were conducted with 10 degree out-of-plane (pitch 10). The average
speed of the f-ABFs was 70.4 µm/s (the forward speed 69.1 µm/s and the drift speed 13.3 µm/s)
under a magnetic field of 9 mT and 90 Hz (Figure 6.31A and Figure 6.32). The swimming of
f-ABF swarms was then studied by IVIS optical imaging. Figure 6.31B shows the controlled
motion of an f-ABF swarm (about 110,000 ABFs dispersed in 1 ml of 5% dextrose) in a
1cm×1cm quartz cuvette. The three Helmholtz coils were placed inside the IVIS chamber under
the camera to allow direct visualization of the f-ABF swarm swimming (Figure 6.26A). The
fluorescence signal generated by the f-ABF dispersion in 5% dextrose appeared as a red cloud
in the IVIS images (Figure 6.31B) and indicated the location and distribution of the f-ABF
swarm. The rotation of the magnetic field was programmed for a precise trajectory: from their
initial position, the f-ABF swarm was exposed to a rotating magnetic field for a propelling
motion to the right (9 mT, 90 Hz, yaw 90, pitch 10), 1 min exposure time) (Figure 6.31B1-2),
then the direction of the rotating magnetic field was changed for a propelling motion in a
100
diagonal direction towards the upper left corner of the quartz cell (9 mT, 90 Hz, 1 min exposure
time) (Figure 6.31B3). Images of the f-ABF swarm before the application of the field were first
recorded, and then every minute after the application of the field.
Figure 6.31. Controlled swimming of a f-ABF swarm in vitro at 9 mT and 90 Hz. (A) A swarm of ABFs swam
on a polished Si wafer in distilled water tracked by optical microscopy. (B) A swarm of ABFs swam in a 1 x 1
cm vial tracked by the IVIS lumina III and Live Imaging software. The color lines around the clouds represent
the contours of the clouds, measured by Icy software and the number ‘●’ indicates the centroids of the clouds.
Reused from [152].
101
Figure 6.32. In vitro forward velocity of ABFs under a rotating magnetic field (9 mT and 90 Hz). (A) An ABF
swarm under an optical microscope. (B) The mean forward speed of different numbered ABFs in (A), and the
error bars present the standard deviations of three measurements. The swimming tests were conducted in distilled
water on a clean Si wafer. The average forward speed of nine ABF was 69.1 m/s. The variability of the speeds
may come from the release of the ABFs from the substrate by sonication and the impurities on the surface on
which ABFs swam. Reused from [152].
0
20
40
60
80
100
1 2 3 4 5 6 7 8 9
Fo
rwar
d S
pee
d /
m
/s
ABFs indicated in figure A
1
2
3
4
56
7
8
9
A
B
50 m
102
Figure 6.33. Data analysis of the swarm swimming in vitro. (A) The overlay of the fluorescent signals in Figure
6.31B1 and Figure 6.31B2. The ‘●’ and ‘■’ present the centroids of the red clouds in Figure 6.31B1 and Figure
6.31B2, respectively. (B) The overlay of the fluorescent signals in Figure 6.31B2 and Figure 6.31B3, and the ‘♦’
presents the centroid of the red cloud in Figure 6.31B3. The inserts show the applied magnetic fields and moving
direction of ABF. The total movement of the centroid was found to be 1.18 mm towards the right in image A and
1.10 mm towards the left diagonal in image B. Reused from [152].
The swimming pattern of the f-ABF swarm was demonstrated by the changing shape of the
cloud (Figure 6.31B). In order to statistically study the movement of the swarm, the movement
of the centroid of the cloud was studied and quantified using Icy software 1.4.3.5 to determine
A
B
103
the centroid position by an active contour function and a custom-made Matlab code (Figure
6.33). The centroid moved 1182 m towards the right from (1) to (2) with a total speed of 19.7
m/s (forward speed 18.5 m/s, drift speed 6.8 m/s) (Figure 6.33A). Afterwards, the centroid
appeared to be displaced by 1102 m diagonally, towards the upper left corner from (2) to (3)
(at total speed of 18.3 m/s, forward speed 18.0 m/s, drift speed 3.6 m/s) (Figure 6.33B).
The movement of the cloud centroid was found to be consistent with the actuation trajectory.
Both the optical (Figure 6.31A) and IVIS tracking (Figure 6.31B) showed that the swarm of f-
ABFs exhibited controlled motion by actuation of the rotational magnetic field. The speed of
movement of the centroid in 5% dextrose was found to be slower than the speed measured in
IPA in Figure 6.32. This could be explained by the possible aggregation of the f-ABFs due to
the surface changes of f-ABFs, the change of media from IPA to an aqueous medium or the
magnetic attraction between f-ABFs. When aggregated, the ABFs are still able to move
forward, however the speed of the aggregated ABFs decreases [161].
6.4.4 In vivo tracking and actuation of f-ABFs
The controlled propelling of the f-ABF swarm in vivo was then investigated. A swarm of f-
ABFs dispersed in 5% dextrose (80,000 f-ABFs in 400 µl) was injected in the intra-peritoneal
cavity of a 4-week old Balb/C mouse. The peritoneal cavity is the space between the parietal
peritoneum and visceral peritoneum [162] and was selected as a site of injection that would
allow enough space for navigation of the swimming f-ABFs. Immediately after injection, the
animals were anesthetized and placed upside down with the lower abdominal part located at
the center of the three pairs of the Helmholtz coils (Figure 6.34-i). This experimental setup is
shown in Figure 6.26A. A fluorescent signal could be detected in the lower part of the
abdominal cavity consistent with the location of the injected f-ABFs (Figure 6.34-ii).
Intraperitoneal administration with a 5% dextrose solution showed no background fluorescent
signal with the optical setup used (Figure 6.26B), confirming that the fluorescent signals
detected originated from the swarm of f-ABFs (Figure 6.26C). A rotational magnetic field was
applied (9 mT, 90 Hz, yaw 0, pitch 10) for 5 min to induce actuation of the f-ABFs in a direction
towards the lower part of the animal body. Images were captured every minute. As shown in
Figure 6.34B, the f-ABF swarm represented by the red and yellow cloud appeared to move
towards the direction of the magnetic field overtime. This suggested that the f-ABFs responded
to the rotational magnetic field and could swim within the intraperitoneal cavity.
104
Figure 6.34. A swarm of f-ABFs tracked and actuated in the intra peritoneal cavity of a Balb-C mouse under 9mT
and 90 Hz. (A) The scheme of in vivo experiment (i) and the original data taken by the IVIS Lumina III using
Live Imaging software; (ii)Image of an anesthetized 4 week old Balb-C mouse inside the magnetic coils. The red
spots represent the fluorescent signal of the injected f-ABFs. (B) A swarm of f-ABFs (80 000 f-ABFs in 400µl)
swimming downward under the actuation of a rotating magnetic field (9mT, 90Hz). The upper array of three
images shows the swarm of f-ABFs swimming downwards (total movement of the swarm center of mass: 1.3
mm) and the bottom images are the magnified image from the upper array. The contour of the yellow cloud was
determined using a custom-made Matlab code and was used to determine the movement of the swarm of f-ABFs.
Reused from [152].
Similar to the in vitro studies, the location of the f-ABFs cloud centroid was monitored and
quantified (Figure 6.35 and Table 3). The downward motion of the cloud centroid was
estimated to be 1.25 mm after 5 min of field exposure and the speed of the cloud was found to
be 6.8 µm/s (forward speed 4.2 µm/s and drift speed 5.4 µm/s), which was slower than the
speed measured in the in vitro studies. The decreased speed of the propelling f-ABF motion
can be explained by the increased fluid viscosity, the potential viscoelasticity of the fluid, the
105
change of surface properties of f-ABFs and the shift of the step-out frequency of f-ABFs inside
the intraperitoneal cavity due to the presence of proteins and other endogenous
biomacromolecules. Additionally, the tissue in vivo presents various anatomical barriers to f-
ABF swimming which could reduce swimming speed. Another explanation of the reduction of
the swimming speed is that during the in vitro experiments under the optical microscope, the
ABFs operate at the step-out frequency. At that frequency they reached their maximum
swimming speed. For the in vitro and in vivo experiments in biological media, this remains
unclear as the step out frequency was determined in water or IPA. As a consequence the f-
ABFs may operate below the step-out frequency.
Figure 6.35. In vivo movement of a swarm of f-ABFs. The images show the swarm of f-ABFs swimming
downwards. The contour of the yellow cloud was determined using a custom-made Matlab code and was used to
determine the movement of the swarm of f-ABFs. The code identifies the center of the yellow contoured area.
The table displays the movement of the center of the yellow contoured area that was determined with a custom-
made Matlab code. Reused from [152].
Table 3. The coordinates of the center of the clouds. Reused from [152].
X (mm) Y (mm) Total movement (mm)
0 min (A) 0 0 0
1 min (B) -0.38 0.16 0.38
3 min (C) -0.51 0.17 0.54
5 min (D) -1.61 1.25 2.04
A B
C D
106
Recent interest has focused on the design and utilization of magnetically propelled constructs
to transport therapeutic agents to precisely targeted locations of the body by active navigation,
despite this remaining challenging [163-167]. Significant progress has been achieved in
manipulating larger scale materials in liquids and tissue. For example, the navigation of a 1cm
ferromagnetic screw in tissue with a rotating magnetic field, as well as the navigation of a
1.5mm bead in the carotid artery using magnetic field gradients were reported [168, 169]. The
use of the magnetic field as an external source of power for micro-swimmers is the common
route for powering swimming motion, as magnetic fields are known to be less invasive than
other forms of actuation. In addition, the utilization of magnetic fields in medicine and in
biological environments is well accepted as the widespread use of magnetic resonance imaging
(MRI) can demonstrate. One of the approaches to navigate magnetic micro and nanoparticles
is to use external magnetic field gradients to provide a translational motion. However, an
important consideration to be taken into account when utilizing magnetic field gradients, is the
system scale size. It has been demonstrated that producing a magnetic field gradient to propel
objects over long distances, typically required for in vivo applications, becomes unattainable
when considering the capabilities of available sources of magnetic fields [14].
Another approach to achieve navigation of magnetic micro and nanoparticles using magnetic
fields is based on the exploitation of the mechanism of travelling wave propulsion. This method
of navigation relies on the creation of a travelling wave to generate propulsion the same way
as that of eukaryotic flagella [170]. This method is found to be a very effective means of
propulsion, possibly more effective than helical propulsion. However, the development of such
systems in terms of fabrication, power, and control appears to be challenging with regards to
the type of distributed actuation seen in eukaryotic flagella [1]. The advantage of using helical
and screw-like structures, as in the case of ABFs, is their ability to operate with the application
of a weak (< 10 mT) and homogenous external magnetic field. At such field intensities
spherical or symmetrical ferromagnetic nanoparticles cannot be propelled, nor be dragged.
When the magnetic field rotates in a plane orthogonal to the main axis of the helix, the ABF
rotate and as a result a propelling motion is achieved (similar to a screw turning), going forward
or backward depending on the direction of rotation of the field [171].
Various micro-robot designs, using different forms of actuation have been reported to swim in
vitro in physiological environments [26, 172-176]. Magnetically actuated
micro/nanoswimmers have been studied in biological fluids such as fetal bovine serum [57],
107
human serum [58], undiluted human blood [70] and saliva [177]. However, the present study
reports previously unattained in vivo navigation of a swarm of microscale, magnetically-
actuated swimmers. The ‘swimming’ motion of the surface-functionalized ABFs as a result of
the application of a mild external rotating magnetic field in the peritoneal cavity of an
anesthetized mouse could be tracked and monitored in real time using live fluorescence
imaging. For many applications, as well as being able to potentially retrieve the swimmers, the
ability to visualize and monitor the ‘swimming’ motion of micro-robots will be essential. This
study demonstrates the ability of navigating these microrobots in an in vivo set-up for the first
time and therefore achieving a major milestone in the field; however for the translation of this
technology for clinical applications there are still important issues to be addressed: such as a)
the interactions of biomolecules with ABF surfaces and their effect of their swimming
properties; b) the number of ABFs to be used in a swarm for optimal delivery of drugs to the
targeted site. c) a thorough investigation of the in vivo toxicity of the ABFs is required in order
to fully understand the excretion mechanism of these entities after the accomplishment of their
mission. Non-optical methods, such MRI imaging, have been found to be useful, in particular
combined with paramagnetic and ferromagnetic microparticles that enhance the MRI signal
contrast. However, MRI imaging becomes unfeasible because high magnetic fields, typically
associated with MRI instrumentation, will interfere with the magnetization of the ABF
swimmers, particularly in ferromagnetic systems actuated with homogenous magnetic fields
[178]. In this case, there is a need to incorporate alternative imaging signals to enable tracking
and visualization.
6.5 Conclusion
The functionalization of ABFs with drugs is essential to enhance their biomedical
performances for targeted drug therapy. Three types of lipid-based nanoscale drug carriers
(DPPC-based liposomes, DOTAP/DOPE liposomes and lipoplexes) were successfully
functionalized on the ABFs, confirmed by QCM-D and CLSM results. The ABFs
functionalized with temperature-sensitive DPPC-based liposomes show the ability to load both
hydrophilic and hydrophobic drugs, and to release calcein (a drug model) by increasing
temperature. The results show that calcein was quickly released at 39ºC, and the release
efficiency of calcein reached 73 ± 15% at 41ºC. The ABFs functionalized with cationic
liposomes DOTAP/DOPE were able to deliver the hydrophilic model drug calcein to C2C12
mouse myoblasts in vitro by direct contact with cells. The ABFs functionalized with lipoplexes
108
loaded with pDNA showed the ability to conduct targeted gene delivery to HEK 293 cells in
vitro. The cells in contact with f-ABFs were successfully transfected by the carried pDNA and
expressed the encoding Venus protein.
Furthermore, the chemical functionalization of the ABFs with a NIR fluorophore, NIR-797,
provides adequate imaging to allow tracking of an f-ABF swarm trajectory in the body of the
animal. Although the speed of the f-ABF swimming motion was found to be lower in vivo than
in vitro due to the possible reasons of higher fluid viscosity, blood and internal fluid flow, the
biomolecules in the medium and the aggregation of ABFs in the medium, the 3-dimensional
directionality of the f-ABF swarm actuation could still be controlled. This represents significant
progress for the development of microscale robots that could be navigated, imaged and tracked
in vivo that could constitute the technology platform for advanced, wirelessly controlled and
imaged vehicles for transport of therapeutics and tools for micro-surgery.
109
7 Summary and future work
7.1 Summary
Inspired by the flagellar propulsive motion of bacteria such as E. Coli, magnetic helical
microrobots, artificial bacterial flagella (ABFs), can perform controlled, micrometer precision,
3D navigation in low Re number environments under low-strength rotating magnetic fields (<
10 mT). They are promising tools for biomedical applications, such as minimally invasive
surgery, cell manipulation and analysis, and targeted therapy.
This dissertation has focused on using these ABFs for potential targeted therapies. The
fabrication, motion control, cytotoxicity, biomedical functionalization for in vitro applications
and in vivo tracking and actuation of these ABFs were studied, and the key findings are
summarized in Figure 7.1.
First, a straightforward and reproducible fabrication process of polymer-based ABFs using 3D
laser direct writing (DLW) and electron beam evaporation was developed. Helical bodies with
lengths ranging from 4 m to 300 m were fabricated by DLW after which a magnetic layer,
such as iron or nickel, and titanium was deposited onto the bodies.
Second, motion control of ABFs in a uniform rotating magnetic field generated by three-pairs
of Helmholtz coils and the swimming behavior of ABFs were explained. The swimming
behavior of ABFs is frequency-dependent. The ABFs wobble at a low frequency, and move in
a stable corkscrew motion in the corkscrew region. In this region, the swimming speed
increases linearly as the frequency increases. The speed reaches the maximum value when the
step-out frequency is reached. The velocity of the ABFs drops dramatically when the frequency
is higher than the step-out frequency. Multiple ABFs can be controlled as a swarm
synchronously or individually.
Third, by use of the preliminary but important tests of biocompatibility, the cytotoxicity of
ABFs in vitro was examined. Cell proliferation, cell viability and cell morphology of mouse
myoblasts C2C12 cells interaction with ABFs reveal little cytotoxicity for up to three days in
culture. Cells contacted with and adhered to the surface of ABFs via lamellipodia and filopodia
interactions.
110
Figure 7.1. Summary of the key findings. The flow of the work is: Fabrication (Adapted from [57].), Motion
control, Cytotoxicity (Reused from [63].), In vitro application, In vivo tracking and actuation of ABFs (Adapted
from [152].).
Fourth, lipid-based nanoscale drug carriers were successfully functionalized on the ABFs. For
biomedical applications such as targeted drug therapy, the functionalization of ABFs with
drugs is essential to enhance their biomedical performances. The functionalized ABFs (f-
ABFs) showed the abilities to be wirelessly steered to the specific sites and to release the carried
drugs models (calcein, a green fluorescent probe, and DNA) into targeted cells in vitro.
Finally, for in vivo applications, ABFs have to be tracked when they are steered inside a living
body. The ABFs were functionalized with a near-infrared dye NIR-797 for in vivo tracking.
We demonstrated the simultaneous injection of over 80,000 ABFs into a mouse peritoneal
cavity, in vivo tracking using near infrared fluorescence, and integrated wireless control of
ABFs using rotating magnetic fields within the mouse. As only weak magnetic fields are
required for actuation with feedback provided by in vivo tracking, the approach can be used
deep within tissue relatively far from an organism’s surface.
111
7.2 Future work
Several challenges remain before realizing biomedical application of these magnetic helical
microrobots. Several possible research directions are possible.
In the area of motion control, we currently control and tune the swimming direction of ABFs
by changing the magnetic fields manually. An algorithm for automatic control of ABFs with
gravity compensation should be integrated into the software of the magnetic setup to facilitate
the control of ABFs.
In the area of material properties, besides in vivo cytotoxicity, a detailed examination of
biocompatibility of ABFs (including in vitro and in vivo cell and tissue compatibility [90])
according to defined standards should be conducted. These tests should be carried out
depending on the environment where ABFs are used and the period the ABFs are used for,
especially for in vivo applications. Additionally, after ABFs complete their task in vivo, a new
challenge is how best to remove the microdevices from the human body. One way is to guide
the ABFs to one area and remove them by minimally invasive surgery. Perhaps the best solution
is to make ABFs biodegradable or bio-absorbable, so new biocompatible and biodegradable
materials, such as some biodegradable hydrogels, are needed to achieve this.
In the area of functionalization, ABFs currently swim mostly in DI-water. In biological and
medical environments, the physiological fluids are more complex than DI-water and there are
various protein and macromolecules inside. So further functionalization of ABFs is needed to
improve the movement of ABFs in these heterogeneous viscous environments.
In the area of in vivo tracking, since the ABFs move in 3D in liquid, tracking of ABFs in 3D
is needed when they move in human bodies, for example, the human blood stream, since the
blood vascular system is made up of 3D tubular structures. The 3D tracking of ABFs is
necessary, which will help the precise control of ABFs to targeted areas in vivo.
112
References
[1] B. J. Nelson, I. K. Kaliakatsos, and J. J. Abbott, "Microrobots for Minimally Invasive
Medicine," Annual Review of Biomedical Engineering, vol. 12, pp. 55-85, 2010.
[2] S. Fusco, M. S. Sakar, S. Kennedy, C. Peters, R. Bottani, F. Starsich, et al., "An
Integrated Microrobotic Platform for On-Demand, Targeted Therapeutic Interventions,"
Advanced Materials, vol. 26, pp. 952-957, 2014.
[3] W. Gao and J. Wang, "The environmental impact of micro/nanomachines: a review,"
ACS Nano, vol. 8, pp. 3170-80, 2014.
[4] J. Orozco, G. Cheng, D. Vilela, S. Sattayasamitsathit, R. Vazquez-Duhalt, G. Valdés-
Ramírez, et al., "Micromotor-Based High-Yielding Fast Oxidative Detoxification of Chemical
Threats," Angewandte Chemie International Edition, vol. 52, pp. 13276-13279, 2013.
[5] V. Magdanz, S. Sanchez, and O. G. Schmidt, "Development of a Sperm-Flagella Driven
Micro-Bio-Robot," Advanced Materials, vol. 25, pp. 6581-6588, 2013.
[6] L. Zhang, K. E. Peyer, and B. J. Nelson, "Artificial bacterial flagella for
micromanipulation," Lab on a Chip, vol. 10, pp. 2203-2215, 2010.
[7] H. C. Berg and R. A. Anderson, "Bacteria swim by rotating their flagellar filaments,"
Nature, vol. 245, pp. 380-382, 1973.
[8] T. Baba, T. Ara, M. Hasegawa, Y. Takai, Y. Okumura, M. Baba, et al., "Construction
of Escherichia coli K-12 in-frame, single-gene knockout mutants: the Keio collection,"
Molecular Systems Biology, vol. 2, p. 2006.0008, 2006.
[9] W. R. DiLuzio, L. Turner, M. Mayer, P. Garstecki, D. B. Weibel, H. C. Berg, et al.,
"Escherichia coli swim on the right-hand side," Nature, vol. 435, pp. 1271-1274, 2005.
[10] D. J. Bell, S. Leutenegger, K. M. Hammar, L. X. Dong, and B. J. Nelson, "Flagella-like
propulsion for microrobots using a nanocoil and a rotating electromagnetic field," in IEEE
International Conference on Robotics and Automation, 2007, pp. 1128-1133.
[11] L. Zhang, J. J. Abbott, L. X. Dong, B. E. Kratochvil, D. Bell, and B. J. Nelson,
113
"Artificial bacterial flagella: Fabrication and magnetic control," Applied Physics Letters, vol.
94, p. 064107, 2009.
[12] L. Zhang, J. J. Abbott, L. Dong, K. E. Peyer, B. E. Kratochvil, H. Zhang, et al.,
"Characterizing the Swimming Properties of Artificial Bacterial Flagella," Nano Letters, vol.
9, pp. 3663-3667, 2009.
[13] K. E. Peyer, L. Zhang, and B. J. Nelson, "Localized non-contact manipulation using
artificial bacterial flagella," Applied Physics Letters, vol. 99, 2011.
[14] J. J. Abbott, K. E. Peyer, M. C. Lagomarsino, L. Zhang, L. X. Dong, I. K. Kaliakatsos,
et al., "How Should Microrobots Swim?," International Journal of Robotics Research, vol. 28,
pp. 1434-1447, 2009.
[15] E. M. Purcell, "Life at low Reynolds number," American Journal of Physics, vol. 45,
pp. 3-11, 1977.
[16] J. Wang and W. Gao, "Nano/Microscale Motors: Biomedical Opportunities and
Challenges," ACS Nano, vol. 6, pp. 5745-5751, 2012.
[17] T. R. Kline, W. F. Paxton, T. E. Mallouk, and A. Sen, "Catalytic Nanomotors: Remote-
Controlled Autonomous Movement of Striped Metallic Nanorods," Angewandte Chemie
International Edition, vol. 44, pp. 744-746, 2005.
[18] S. Fournier-Bidoz, A. C. Arsenault, I. Manners, and G. A. Ozin, "Synthetic self-
propelled nanorotors," Chemical Communications, pp. 441-443, 2005.
[19] W. F. Paxton, A. Sen, and T. E. Mallouk, "Motility of Catalytic Nanoparticles through
Self-Generated Forces," Chemistry – A European Journal, vol. 11, pp. 6462-6470, 2005.
[20] J. G. Gibbs and Y. P. Zhao, "Autonomously motile catalytic nanomotors by bubble
propulsion," Applied Physics Letters, vol. 94, 2009.
[21] A. A. Solovev, Y. F. Mei, E. B. Urena, G. S. Huang, and O. G. Schmidt, "Catalytic
Microtubular Jet Engines Self-Propelled by Accumulated Gas Bubbles," Small, vol. 5, pp.
1688-1692, 2009.
[22] W. Gao and O. C. Farokhzad, "Self-Propelled Microrockets to Capture and Isolate
114
Circulating Tumor Cells," Angewandte Chemie International Edition, vol. 50, pp. 7220-7221,
2011.
[23] M. S. Sakar, E. B. Steager, D. H. Kim, A. A. Julius, M. Kim, V. Kumar, et al.,
"Modeling, control and experimental characterization of microbiorobots," International
Journal of Robotics Research, vol. 30, pp. 647-658, 2011.
[24] B. J. Williams, S. V. Anand, J. Rajagopalan, and M. T. A. Saif, "A self-propelled
biohybrid swimmer at low Reynolds number," Nature communications, vol. 5, 2014.
[25] M. H. Oddy and J. G. Santiago, "A method for determining electrophoretic and
electroosmotic mobilities using AC and DC electric field particle displacements," Journal of
Colloid and Interface Science, vol. 269, pp. 192-204, 2004.
[26] S. T. Chang, V. N. Paunov, D. N. Petsev, and O. D. Velev, "Remotely powered self-
propelling particles and micropumps based on miniature diodes," Nature Materials, vol. 6, pp.
235-240, 2007.
[27] G. Hwang, S. Haliyo, and a. S. Régnier, "Remotely powered propulsion of helical
nanobelts," in Robotics Science and Systems, Zaragoza, Spain, 2010.
[28] G. Hwang, R. Braive, L. Couraud, A. Cavanna, O. Abdelkarim, I. Robert-Philip, et al.,
"Electro-osmotic propulsion of helical nanobelt swimmers," International Journal of Robotics
Research, vol. 30, pp. 806-819, 2011.
[29] D. G. Grier, "A revolution in optical manipulation," Nature, vol. 424, pp. 810-816,
2003.
[30] S. Ahmed, W. Wang, L. O. Mair, R. D. Fraleigh, S. Li, L. A. Castro, et al., "Steering
Acoustically Propelled Nanowire Motors toward Cells in a Biologically Compatible
Environment Using Magnetic Fields," Langmuir, vol. 29, pp. 16113-16118, 2013.
[31] W. Wang, L. A. Castro, M. Hoyos, and T. E. Mallouk, "Autonomous Motion of
Metallic Microrods Propelled by Ultrasound," ACS Nano, vol. 6, pp. 6122-6132, 2012.
[32] W. Wang, S. Li, L. Mair, S. Ahmed, T. J. Huang, and T. E. Mallouk, "Acoustic
Propulsion of Nanorod Motors Inside Living Cells," Angewandte Chemie International
Edition, vol. 53, pp. 3201-3204, 2014.
115
[33] W. Wang, W. Duan, S. Ahmed, T. E. Mallouk, and A. Sen, "Small power: Autonomous
nano- and micromotors propelled by self-generated gradients," Nano Today, vol. 8, pp. 531-
554, 2013.
[34] D. Jiles, "Introduction to magnetism and magnetic materials," London: Chapman and
Hall, 1991.
[35] T. Honda, K. I. Arai, and K. Ishiyama, "Micro swimming mechanisms propelled by
external magnetic fields," IEEE Transactions on Magnetics, vol. 32, pp. 5085-5087, 1996.
[36] Z. Ye, S. Regnier, and M. Sitti, "Rotating Magnetic Miniature Swimming Robots With
Multiple Flexible Flagella," IEEE Transactions on Robotics, vol. 30, pp. 3-13, 2014.
[37] J. J. Abbott, O. Ergeneman, M. P. Kummer, A. M. Hirt, and B. J. Nelson, "Modeling
magnetic torque and force for controlled manipulation of soft-magnetic bodies," IEEE
Transactions on Robotics, vol. 23, pp. 1247-1252, 2007.
[38] R. Dreyfus, J. Baudry, M. L. Roper, M. Fermigier, H. A. Stone, and J. Bibette,
"Microscopic artificial swimmers," Nature, vol. 437, pp. 862-865, 2005.
[39] M. Suter, L. Zhang, E. Siringil, C. Peters, T. Luehmann, O. Ergeneman, et al.,
"Superparamagnetic microrobots: fabrication by two-photon polymerization and
biocompatibility," Biomedical Microdevices, pp. 1-7, 2013.
[40] K. Peyer, E. Siringil, L. Zhang, M. Suter, and B. Nelson, "Bacteria-Inspired Magnetic
Polymer Composite Microrobots," in Biomimetic and Biohybrid Systems, ed: Springer Berlin
Heidelberg, 2013, pp. 216-227.
[41] K. E. Peyer, L. Zhang, and B. J. Nelson, "Bio-inspired magnetic swimming microrobots
for biomedical applications," Nanoscale, vol. 5, pp. 1259-1272, 2013.
[42] W. Gao, S. Sattayasamitsathit, K. M. Manesh, D. Weihs, and J. Wang, "Magnetically
Powered Flexible Metal Nanowire Motors," Journal of the American Chemical Society, vol.
132, pp. 14403-14405, 2010.
[43] L. Zhang, T. Petit, Y. Lu, B. E. Kratochvil, K. E. Peyer, R. Pei, et al., "Controlled
Propulsion and Cargo Transport of Rotating Nickel Nanowires near a Patterned Solid Surface,"
ACS Nano, vol. 4, pp. 6228-6234, 2010.
116
[44] C. E. Sing, L. Schmid, M. F. Schneider, T. Franke, and A. Alexander-Katz, "Controlled
surface-induced flows from the motion of self-assembled colloidal walkers," Proceedings of
the National Academy of Sciences, vol. 107, pp. 535-540, 2010.
[45] A. W. Mahoney and J. J. Abbott, "Managing magnetic force applied to a magnetic
device by a rotating dipole field," Applied Physics Letters, vol. 99, p. 134103, 2011.
[46] T. Y. Huang, F. Qiu, H. W. Tung, X. B. Chen, B. J. Nelson, and M. S. Sakar,
"Generating mobile fluidic traps for selective three-dimensional transport of microobjects,"
Applied Physics Letters, vol. 105, pp. 114102 - 114102-5, 2014.
[47] R. S. M. Rikken, R. J. M. Nolte, J. C. Maan, J. C. M. van Hest, D. A. Wilson, and P. C.
M. Christianen, "Manipulation of micro- and nanostructure motion with magnetic fields," Soft
Matter, vol. 10, pp. 1295-1308, 2014.
[48] S. Floyd, C. Pawashe, and M. Sitti, "An untethered magnetically actuated micro-robot
capable of motion on arbitrary surfaces," in 2008 IEEE International Conference on Robotics
and Automation ( ICRA), 2008, pp. 419-424.
[49] D. R. Frutiger, K. Vollmers, B. E. Kratochvil, and B. J. Nelson, "Small, Fast, and Under
Control: Wireless Resonant Magnetic Micro-agents," International Journal of Robotics
Research, vol. 29, pp. 613-636, 2010.
[50] H.-W. Tung, D. R. Frutiger, S. Pane, and B. J. Nelson, "Polymer-based Wireless
Resonant Magnetic microrobots," in 2012 IEEE International Conference on Robotics and
Automation (ICRA), 2012, pp. 715-720.
[51] M. P. Kummer, J. J. Abbott, B. E. Kratochvil, R. Borer, A. Sengul, and B. J. Nelson,
"OctoMag: An Electromagnetic System for 5-DOF Wireless Micromanipulation," IEEE
Transactions on Robotics, vol. 26, pp. 1006-1017, 2010.
[52] S. Schurle, K. E. Peyer, B. E. Kratochvil, and B. J. Nelson, "Holonomic 5-DOF
magnetic control of 1D nanostructures," in 2012 IEEE International Conference on Robotics
and Automation (ICRA), 2012, pp. 1081-1086.
[53] K. E. Peyer, S. Tottori, F. Qiu, L. Zhang, and B. J. Nelson, "Magnetic Helical
Micromachines," Chemistry – A European Journal, vol. 19, pp. 28-38, 2013.
117
[54] K. Kikuchi, A. Yamazaki, M. Sendoh, K. Ishiyama, and K. I. Arai, "Fabrication of a
spiral type magnetic micromachine for trailing a wire," IEEE Transactions on Magnetics, vol.
41, pp. 4012-4014, 2005.
[55] A. Ghosh and P. Fischer, "Controlled Propulsion of Artificial Magnetic Nanostructured
Propellers," Nano Letters, vol. 9, pp. 2243-2245, 2009.
[56] A. W. Mahoney, J. C. Sarrazin, E. Bamberg, and J. J. Abbott, "Velocity Control with
Gravity Compensation for Magnetic Helical Microswimmers," Advanced Robotics, vol. 25, pp.
1007-1028, 2011.
[57] S. Tottori, L. Zhang, F. Qiu, K. K. Krawczyk, A. Franco-Obregón, and B. J. Nelson,
"Magnetic Helical Micromachines: Fabrication, Controlled Swimming, and Cargo Transport,"
Advanced Materials, vol. 24, pp. 811-816, 2012.
[58] W. Gao, X. Feng, A. Pei, C. R. Kane, R. Tam, C. Hennessy, et al., "Bioinspired Helical
Microswimmers Based on Vascular Plants," Nano Letters, vol. 14, pp. 305-310, 2013.
[59] J. X. Li, S. Sattayasamitsathit, R. F. Dong, W. Gao, R. Tam, X. M. Feng, et al.,
"Template electrosynthesis of tailored-made helical nanoswimmers," Nanoscale, vol. 6, pp.
9415-9420, 2014.
[60] K. E. Peyer, L. Zhang, B. E. Kratochvil, and B. J. Nelson, "Non-ideal Swimming of
Artificial Bacterial Flagella Near a Surface," in 2010 IEEE International Conference on
Robotics and Automation, 2010, pp. 96-101.
[61] S. Tottori and B. J. Nelson, "Artificial helical microswimmers with mastigoneme-
inspired appendages," Biomicrofluidics, vol. 7, p. 61101, 2013.
[62] K. E. Peyer, F. Qiu, L. Zhang, and B. J. Nelson, "Movement of Artificial Bacterial
Flagella in Heterogeneous Viscous Environments at the Microscale," in IEEE/RSJ
International Conference on Intelligent Robots and Systems, New York, 2012, pp. 2553-2558.
[63] F. Qiu, L. Zhang, S. Tottori, K. Marquardt, K. Krawczyk, A. Franco-Obregon, et al.,
"Bio-inspired microrobots," Materials Today, vol. 15, pp. 463-463, 2012.
[64] F. Qiu, R. Mhanna, L. Zhang, Y. Ding, K. Sugihara, M. Zenobi-Wong, et al., "Artificial
bacterial flagella functionalized with temperature-sensitive liposomes for biomedical
118
applications," in The 17th International Conference on Solid-State Sensors, Actuators and
Microsystems (TRANSDUCERS & EUROSENSORS XXVII), Barcelona, Spain 2013, pp. 2130-
2133.
[65] F. Qiu, L. Zhang, K. E. Peyer, M. Casarosa, A. Franco-Obregon, H. Choi, et al.,
"Noncytotoxic artificial bacterial flagella fabricated from biocompatible ORMOCOMP and
iron coating," Journal of Materials Chemistry B, vol. 2, pp. 357-362, 2014.
[66] R. Mhanna, F. Qiu, L. Zhang, Y. Ding, K. Sugihara, M. Zenobi-Wong, et al., "Artificial
Bacterial Flagella for Remote-Controlled Targeted Single-Cell Drug Delivery," Small, vol. 10,
pp. 1953-1957, 2014.
[67] F. Qiu, R. Mhanna, L. Zhang, Y. Ding, S. Fujita, and B. J. Nelson, "Artificial bacterial
flagella functionalized with temperature-sensitive liposomes for controlled release," Sensors
and Actuators B: Chemical, vol. 196, pp. 676-681, 2014.
[68] F. Qiu, S. Fujita, R. Mhanna, L. Zhang, B. R. Simona, and B. J. Nelson, "Magnetic
Helical Microswimmers Functionalized with Lipoplexes for Targeted Gene Delivery,"
Advanced Functional Materials, vol. 25, pp. 1666-1671, 2015.
[69] D. Schamel, A. G. Mark, J. G. Gibbs, C. Miksch, K. I. Morozov, A. M. Leshansky, et
al., "Nanopropellers and Their Actuation in Complex Viscoelastic Media," ACS Nano, vol. 8,
pp. 8794-801, 2014.
[70] P. L. Venugopalan, R. Sai, Y. Chandorkar, B. Basu, S. Shivashankar, and A. Ghosh,
"Conformal Cytocompatible Ferrite Coatings Facilitate the Realization of a Nanovoyager in
Human Blood," Nano Letters, vol. 14, pp. 1968-75, 2014.
[71] T. W. R. Fountain, P. V. Kailat, and J. J. Abbott, "Wireless control of magnetic helical
microrobots using a rotating-permanent-magnet manipulator," in 2010 IEEE International
Conference on Robotics and Automation (ICRA), 2010, pp. 576-581.
[72] J. Lighthill, "FLAGELLAR HYDRODYNAMICS - NEUMANN,JV LECTURE,
1975," Siam Review, vol. 18, pp. 161-230, 1976.
[73] L. Liu, L. Zhang, S. M. Kim, and S. Park, "Helical Metallic Micro- and Nano-structures:
Fabrication and Application," Nanoscale, vol. 6, pp. 9355-65, 2014.
119
[74] V. Y. Prinz, V. A. Seleznev, A. K. Gutakovsky, A. V. Chehovskiy, V. V.
Preobrazhenskii, M. A. Putyato, et al., "Free-standing and overgrown InGaAs/GaAs
nanotubes, nanohelices and their arrays," Physica E: Low-dimensional Systems and
Nanostructures, vol. 6, pp. 828-831, 2000.
[75] O. G. Schmidt and K. Eberl, "Nanotechnology - Thin solid films roll up into
nanotubes," Nature, vol. 410, pp. 168-168, 2001.
[76] L. Zhang, E. Deckhardt, A. Weber, C. Schonenberger, and D. Grutzmacher,
"Controllable fabrication of SiGe/Si and SiGe/Si/Cr helical nanobelts," Nanotechnology, vol.
16, pp. 655-663, 2005.
[77] S. Schuerle, S. Pané, E. Pellicer, J. Sort, M. D. Baró, and B. J. Nelson, "Helical and
Tubular Lipid Microstructures that are Electroless-Coated with CoNiReP for Wireless
Magnetic Manipulation," Small, vol. 8, pp. 1498-1502, 2012.
[78] M. A. Zeeshan, R. Grisch, E. Pellicer, K. M. Sivaraman, K. E. Peyer, J. Sort, et al.,
"Hybrid Helical Magnetic Microrobots Obtained by 3D Template-Assisted Electrodeposition,"
Small, vol. 10, pp. 1284-1288, 2014.
[79] S. H. Wu, J. Serbin, and M. Gu, "Two-photon polymerisation for three-dimensional
micro-fabrication," Journal of Photochemistry and Photobiology a-Chemistry, vol. 181, pp. 1-
11, 2006.
[80] K. Anatoli, Nanoelectronics and photonics : from atoms to materials, devices, and
architectures. New York: New York : Springer, 2008.
[81] S. Maruo and S. Kawata, "Two-photon-absorbed near-infrared photopolymerization for
three-dimensional microfabrication," Journal of Microelectromechanical Systems, vol. 7, pp.
411-415, 1998.
[82] A. Ovsianikov, A. Ostendorf, and B. N. Chichkov, "Three-dimensional
photofabrication with femtosecond lasers for applications in photonics and biomedicine,"
Applied Surface Science, vol. 253, pp. 6599-6602, 2007.
[83] Z. Wang and H. Y. Zheng, "Femtosecond laser direct writing microfluidic channels
inside photosensitive glass," presented at the 2008 IEEE Photonicsglobal@Singapore, 2008.
120
[84] A. Doraiswamy, A. Ovsianikov, S. D. Gittard, N. A. Monteiro-Riviere, R. Crombez, E.
Montalvo, et al., "Fabrication of Microneedles Using Two Photon Polymerization for
Transdermal Delivery of Nanomaterials," Journal of Nanoscience and Nanotechnology, vol.
10, pp. 6305-6312, 2010.
[85] R. J. Narayan, A. Doraiswamy, D. B. Chrisey, and B. N. Chichkov, "Medical
prototyping using two photon polymerization," Materials Today, vol. 13, pp. 42-48, 2010.
[86] S. Kawata, H. B. Sun, T. Tanaka, and K. Takada, "Finer features for functional
microdevices - Micromachines can be created with higher resolution using two-photon
absorption," Nature, vol. 412, pp. 697-698, 2001.
[87] I. S. M. Khalil, H. C. Dijkslag, L. Abelmann, and S. Misra, "MagnetoSperm: A
microrobot that navigates using weak magnetic fields," Applied Physics Letters, vol. 104, pp.
-, 2014.
[88] S. Kim, F. Qiu, S. Kim, A. Ghanbari, C. Moon, L. Zhang, et al., "Fabrication and
Characterization of Magnetic Microrobots for Three-Dimensional Cell Culture and Targeted
Transportation," Advanced Materials, vol. 25, pp. 5863-5868, 2013.
[89] K. E. Peyer, A. W. Mahoney, L. Zhang, J. J. Abbott, and B. J. Nelson, "Chapter 7 -
Bacteria-Inspired Microrobots," in Microbiorobotics, M. Kim, A. A. Julius, and E. Steager,
Eds., ed Boston: William Andrew Publishing, 2012, pp. 165-199.
[90] D. R. Buddy, Biomaterials science : an introduction to materials in medicine, 3rd ed.
ed. Amsterdam: Amsterdam : Elsevier, 2013.
[91] D. F. Williams, "On the mechanisms of biocompatibility," Biomaterials, vol. 29, pp.
2941-2953, 2008.
[92] S. Schlie, A. Ngezahayo, A. Ovsianikov, T. Fabian, H. A. Kolb, H. Haferkamp, et al.,
"Three-dimensional cell growth on structures fabricated from ORMOCER (R) by two-photon
polymerization technique," Journal of Biomaterials Applications, vol. 22, pp. 275-287, 2007.
[93] X. Hu, S. H. Park, E. S. Gil, X. X. Xia, A. S. Weiss, and D. L. Kaplan, "The influence
of elasticity and surface roughness on myogenic and osteogenic-differentiation of cells on silk-
elastin biomaterials," Biomaterials, vol. 32, pp. 8979-89, 2011.
121
[94] M. Ermolli, C. Menne, G. Pozzi, M. A. Serra, and L. A. Clerici, "Nickel, cobalt and
chromium-induced cytotoxicity and intracellular accumulation in human hacat keratinocytes,"
Toxicology, vol. 159, pp. 23-31, 2001.
[95] M. Taira, M. S. Toguchi, Y. Hamada, M. Okazaki, J. Takahashi, R. Ito, et al., "Studies
on cytotoxicity of nickel ions using C3H10T1/2 fibroblast cells," J Oral Rehabil, vol. 27, pp.
1068-72, 2000.
[96] M. Mikulewicz and K. Chojnacka, "Cytocompatibility of Medical Biomaterials
Containing Nickel by Osteoblasts: a Systematic Literature Review," Biological Trace Element
Research, vol. 142, pp. 865-889, 2011.
[97] P. P. Mueller, S. Arnold, M. Badar, D. Bormann, F. W. Bach, A. Drynda, et al.,
"Histological and molecular evaluation of iron as degradable medical implant material in a
murine animal model," Journal of Biomedical Materials Research Part A, vol. 100, pp. 2881-
2889, 2012.
[98] M. Peuster, P. Wohlsein, M. Brugmann, M. Ehlerding, K. Seidler, C. Fink, et al., "A
novel approach to temporary stenting: degradable cardiovascular stents produced from
corrodible metal - results 6-18 months after implantation into New Zealand white rabbits,"
Heart, vol. 86, pp. 563-569, 2001.
[99] M. Peuster, C. Hesse, T. Schloo, C. Fink, P. Beerbaum, and C. von Schnakenburg,
"Long-term biocompatibility of a corrodible peripheral iron stent in the porcine descending
aorta," Biomaterials, vol. 27, pp. 4955-4962, 2006.
[100] F. Klein, T. Striebel, J. Fischer, Z. X. Jiang, C. M. Franz, G. von Freymann, et al.,
"Elastic Fully Three-dimensional Microstructure Scaffolds for Cell Force Measurements,"
Advanced Materials, vol. 22, pp. 868-871, 2010.
[101] S. H. Yoon, Y. K. Kim, E. D. Han, Y. H. Seo, B. H. Kim, and M. R. K. Mofrad, "Passive
control of cell locomotion using micropatterns: the effect of micropattern geometry on the
migratory behavior of adherent cells," Lab on a Chip, vol. 12, pp. 2391-2402, 2012.
[102] C. Schizas and D. Karalekas, "Mechanical characteristics of an Ormocomp (R)
biocompatible hybrid photopolymer," Journal of the Mechanical Behavior of Biomedical
Materials, vol. 4, pp. 99-106, 2011.
122
[103] A. Doraiswamy, C. Jin, R. J. Narayan, P. Mageswaran, P. Mente, R. Modi, et al., "Two
photon induced polymerization of organic–inorganic hybrid biomaterials for microstructured
medical devices," Acta Biomaterialia, vol. 2, pp. 267-275, 2006.
[104] F. Klein, B. Richter, T. Striebel, C. M. Franz, G. v. Freymann, M. Wegener, et al.,
"Two-Component Polymer Scaffolds for Controlled Three-Dimensional Cell Culture,"
Advanced Materials, vol. 23, pp. 1341-1345, 2011.
[105] S. L. Tao, K. C. Popat, J. J. Norman, and T. A. Desai, "Surface modification of SU-8
for enhanced biofunctionality and nonfouling properties," Langmuir, vol. 24, pp. 2631-2636,
2008.
[106] L. Kupcsik, "Estimation of cell number based on metabolic activity: the MTT reduction
assay," Methods in molecular biology, vol. 740, pp. 13-9, 2011.
[107] A. Abarrategi, M. C. Gutierrez, C. Moreno-Vicente, M. J. Hortiguela, V. Ramos, J. L.
Lopez-Lacomba, et al., "Multiwall carbon nanotube scaffolds for tissue engineering purposes,"
Biomaterials, vol. 29, pp. 94-102, 2008.
[108] D. Fischer, Y. Li, B. Ahlemeyer, J. Krieglstein, and T. Kissel, "In vitro cytotoxicity
testing of polycations: influence of polymer structure on cell viability and hemolysis,"
Biomaterials, vol. 24, pp. 1121-1131, 2003.
[109] S. Sanchez, A. N. Ananth, V. M. Fomin, M. Viehrig, and O. G. Schmidt, "Superfast
Motion of Catalytic Microjet Engines at Physiological Temperature," Journal of the American
Chemical Society, vol. 133, pp. 14860-14863, 2011.
[110] J. van Meerloo, G. L. Kaspers, and J. Cloos, "Cell Sensitivity Assays: The MTT Assay,"
in Cancer Cell Culture. vol. 731, I. A. Cree, Ed., ed: Humana Press, 2011, pp. 237-245.
[111] T. Mosmann, "RAPID COLORIMETRIC ASSAY FOR CELLULAR GROWTH AND
SURVIVAL - APPLICATION TO PROLIFERATION AND CYTO-TOXICITY ASSAYS,"
Journal of Immunological Methods, vol. 65, pp. 55-63, 1983.
[112] K. Ishizaki, Y. Sugita, F. Iwasa, H. Minamikawa, T. Ueno, M. Yamada, et al.,
"Nanometer-thin TiO2 enhances skeletal muscle cell phenotype and behavior," International
Journal of Nanomedicine, vol. 6, pp. 2191-2203, 2011.
123
[113] M. Moravej, F. Prima, M. Fiset, and D. Mantovani, "Electroformed iron as new
biomaterial for degradable stents: Development process and structure-properties relationship,"
Acta Biomaterialia, vol. 6, pp. 1726-1735, 2010.
[114] A. M. Schmidt and D. S. Azambuja, "Corrosion Behavior of Ti and TI6Al4V in Citrate
Buffers Containing Fluoride Ions," Materials Research-Ibero-American Journal of Materials,
vol. 13, pp. 45-50, 2010.
[115] R. Waksman, R. Pakala, R. Baffour, R. Seabron, D. Hellinga, and F. O. Tio, "Short-
term effects of biocorrodible iron stents in porcine coronary arteries," Journal of Interventional
Cardiology, vol. 21, pp. 15-20, 2008.
[116] D. van Swaay and A. deMello, "Microfluidic methods for forming liposomes," Lab on
a Chip, vol. 13, pp. 752-767, 2013.
[117] M. Bally, K. Bailey, K. Sugihara, D. Grieshaber, J. Vörös, and B. Städler, "Liposome
and Lipid Bilayer Arrays Towards Biosensing Applications," Small, vol. 6, pp. 2481-2497,
2010.
[118] M. E. Lynge, M. Baekgaard Laursen, L. Hosta-Rigau, B. E. B. Jensen, R. Ogaki, A. A.
A. Smith, et al., "Liposomes as Drug Deposits in Multilayered Polymer Films," Acs Applied
Materials & Interfaces, vol. 5, pp. 2967-2975, 2013.
[119] S. Bibi, E. Lattmann, A. R. Mohammed, and Y. Perrie, "Trigger release liposome
systems: local and remote controlled delivery?," Journal of Microencapsulation, vol. 29, pp.
262-276, 2012.
[120] A. M. Ponce, Z. Vujaskovic, F. Yuan, D. Needham, and M. W. Dewhirst,
"Hyperthermia mediated liposomal drug delivery," International Journal of Hyperthermia, vol.
22, pp. 205-213, 2006.
[121] D. Needham and M. W. Dewhirst, "The development and testing of a new temperature-
sensitive drug delivery system for the treatment of solid tumors," Advanced Drug Delivery
Reviews, vol. 53, pp. 285-305, 2001.
[122] T. Tagami, M. J. Ernsting, and S.-D. Li, "Optimization of a novel and improved
thermosensitive liposome formulated with DPPC and a Brij surfactant using a robust in vitro
124
system," Journal of Controlled Release, vol. 154, pp. 290-297, 2011.
[123] R. R. Sawant and V. P. Torchilin, "Liposomes as 'smart' pharmaceutical nanocarriers,"
Soft Matter, vol. 6, pp. 4026-4044, 2010.
[124] J. K. Mills and D. Needham, "Lysolipid incorporation in
dipalmitoylphosphatidylcholine bilayer membranes enhances the ion permeability and drug
release rates at the membrane phase transition," Biochimica Et Biophysica Acta-
Biomembranes, vol. 1716, pp. 77-96, 2005.
[125] F. Olson, C. A. Hunt, F. C. Szoka, W. J. Vail, and D. Papahadjopoulos, "Preparation of
liposomes of defined size distribution by extrusion through polycarbonate membranes,"
Biochimica et Biophysica Acta, vol. 557, pp. 9-23, 1979.
[126] R. C. Macdonald, R. I. Macdonald, B. P. M. Menco, K. Takeshita, N. K. Subbarao, and
L. R. Hu, "Small-volume extrusion apparatus for preparation of large, unilamellar vesicles,"
Biochimica et Biophysica Acta, vol. 1061, pp. 297-303, 1991.
[127] N.-J. Cho, C. W. Frank, B. Kasemo, and F. Höök, "Quartz crystal microbalance with
dissipation monitoring of supported lipid bilayers on various substrates," Nature Protocols,
vol. 5, pp. 1096-1106, 2010.
[128] E. Reimhult, F. Hook, and B. Kasemo, "Intact vesicle adsorption and supported
biomembrane formation from vesicles in solution: Influence of surface chemistry, vesicle size,
temperature, and osmotic pressure," Langmuir, vol. 19, pp. 1681-1691, 2003.
[129] M. Afadzi, C. d. L. Davies, Y. H. Hansen, T. Johansen, Ø. K. Standal, R. Hansen, et
al., "Effect of Ultrasound Parameters on the Release of Liposomal Calcein," Ultrasound in
Medicine & Biology, vol. 38, pp. 476-486, 2012.
[130] A. Fortunelli and S. Monti, "Simulations of Lipid Adsorption on TiO2 Surfaces in
Solution," Langmuir, vol. 24, pp. 10145-10154, 2008.
[131] D. Felnerova, J. F. Viret, R. Gluck, and C. Moser, "Liposomes and virosomes as
delivery systems for antigens, nucleic acids and drugs," Current Opinion in Biotechnology, vol.
15, pp. 518-529, 2004.
[132] N. Sakaguchi, C. Kojima, A. Harada, K. Koiwai, and K. Kono, "The correlation
125
between fusion capability and transfection activity in hybrid complexes of lipoplexes and pH-
sensitive liposomes," Biomaterials, vol. 29, pp. 4029-36, 2008.
[133] D. Nicolosi, M. Scalia, V. M. Nicolosi, and R. Pignatello, "Encapsulation in fusogenic
liposomes broadens the spectrum of action of vancomycin against Gram-negative bacteria,"
International journal of antimicrobial agents, vol. 35, pp. 553-8, 2010.
[134] V. Budker, V. Gurevich, J. E. Hagstrom, F. Bortzov, and J. A. Wolff, "pH-sensitive,
cationic liposomes: a new synthetic virus-like vector," Nature biotechnology, vol. 14, pp. 760-
4, 1996.
[135] J. W. Liu, A. Stace-Naughton, X. M. Jiang, and C. J. Brinker, "Porous Nanoparticle
Supported Lipid Bilayers (Protocells) as Delivery Vehicles," Journal of the American
Chemical Society, vol. 131, pp. 1354-5, 2009.
[136] A. Bahia, E. G. Azevedo, L. A. M. Ferreira, and F. Frezard, "New insights into the
mode of action of ultradeformable vesicles using calcein as hydrophilic fluorescent marker,"
European Journal of Pharmaceutical Sciences, vol. 39, pp. 90-96, 2010.
[137] E. Reimhult, F. Hook, and B. Kasemo, "Vesicle adsorption on SiO2 and TiO2:
Dependence on vesicle size," Journal of Chemical Physics, vol. 117, pp. 7401-7404, 2002.
[138] N. Graf, E. Thomasson, A. Tanno, J. Voeroes, and T. Zambelli, "Spontaneous
Formation of a Vesicle Multilayer on Top of an Exponentially Growing Polyelectrolyte
Multilayer Mediated by Diffusing Poly-L-lysine," Journal of Physical Chemistry B, vol. 115,
pp. 12386-12391, 2011.
[139] K. Sugihara, J. Vörös, and T. Zambelli, "A Gigaseal Obtained with a Self-Assembled
Long-Lifetime Lipid Bilayer on a Single Polyelectrolyte Multilayer-Filled Nanopore," ACS
Nano, vol. 4, pp. 5047-5054, 2010.
[140] Duzgune, scedil, and S. Nir, "Mechanisms and kinetics of liposome-cell interactions,"
Advanced drug delivery reviews, vol. 40, pp. 3-18, 1999.
[141] R. Leventis and J. R. Silvius, "Interactions of mammalian cells with lipid dispersions
containing novel metabolizable cationic amphiphiles," Biochimica et Biophysica Acta (BBA) -
Biomembranes, vol. 1023, pp. 124-132, 1990.
126
[142] C. Tros de Ilarduya, Y. Sun, and N. Düzgüneş, "Gene delivery by lipoplexes and
polyplexes," European Journal of Pharmaceutical Sciences, vol. 40, pp. 159-170, 2010.
[143] Z. u. Rehman, D. Hoekstra, and I. S. Zuhorn, "Mechanism of Polyplex- and Lipoplex-
Mediated Delivery of Nucleic Acids: Real-Time Visualization of Transient Membrane
Destabilization without Endosomal Lysis," ACS Nano, vol. 7, pp. 3767-3777, 2013.
[144] A. Elouahabi and J. M. Ruysschaert, "Formation and intracellular trafficking of
lipoplexes and polyplexes," Molecular Therapy, vol. 11, pp. 336-47, 2005.
[145] M. S. Al-Dosari and X. Gao, "Nonviral Gene Delivery: Principle, Limitations, and
Recent Progress," AAPS Journal, vol. 11, pp. 671-681, 2009.
[146] S. Fujita, R. Onuki-Nagasaki, J. Fukuda, J. Enomoto, S. Yamaguchi, and M. Miyake,
"Development of super-dense transfected cell microarrays generated by piezoelectric inkjet
printing," Lab on a Chip, vol. 13, pp. 77-80, 2013.
[147] B. Dalby, S. Cates, A. Harris, E. C. Ohki, M. L. Tilkins, P. J. Price, et al., "Advanced
transfection with Lipofectamine 2000 reagent: primary neurons, siRNA, and high-throughput
applications," Methods, vol. 33, pp. 95-103, 2004.
[148] E. Delyagina, W. Li, N. Ma, and G. Steinhoff, "Magnetic targeting strategies in gene
delivery," Nanomedicine, vol. 6, pp. 1593-1604, 2011.
[149] P. L. Felgner, T. R. Gadek, M. Holm, R. Roman, H. W. Chan, M. Wenz, et al.,
"Lipofection: a highly efficient, lipid-mediated DNA-transfection procedure," Proceedings of
the National Academy of Sciences of the United States of America, vol. 84, pp. 7413-7, 1987.
[150] T. Yoshikawa, E. Uchimura, M. Kishi, D. P. Funeriu, M. Miyake, and J. Miyake,
"Transfection microarray of human mesenchymal stem cells and on-chip siRNA gene
knockdown," Journal of Controlled Release, vol. 96, pp. 227-232, 2004.
[151] J. Ziauddin and D. M. Sabatini, "Microarrays of cells expressing defined cDNAs,"
Nature, vol. 411, pp. 107-110, 2001.
[152] A. Servant, F. Qiu, M. Mazza, K. Kostarelos, and B. J. Nelson, "Controlled in vivo
Swimming of a Swarm of Bacteria-Like Microrobotic Flagella," Advanced Materials, vol. 27,
pp. 2981-2988, 2015.
127
[153] R. Bakhshi, A. Darbyshire, J. E. Evans, Z. You, J. Lu, and A. M. Seifalian, "Polymeric
coating of surface modified nitinol stent with POSS-nanocomposite polymer," Colloids and
surfaces. B, Biointerfaces, vol. 86, pp. 93-105, 2011.
[154] W. Haider, N. Munroe, V. Tek, C. Pulletikurthi, P. K. S. Gill, and S. Pandya, "Surface
modifications of nitinol," Journal of long-term effects of medical implants, vol. 19, pp. 113-22,
2009.
[155] C. Menard-Moyon, C. Fabbro, M. Prato, and A. Bianco, "One-Pot Triple
Functionalization of Carbon Nanotubes," Chemistry- A European Journal, vol. 17, pp. 3222-
3227, 2011.
[156] K.-H. Chuang, H.-E. Wang, F.-M. Chen, S.-C. Tzou, C.-M. Cheng, Y.-C. Chang, et al.,
"Endocytosis of PEGylated Agents Enhances Cancer Imaging and Anticancer Efficacy,"
Molecular Cancer Therapeutics, vol. 9, pp. 1903-1912, 2010.
[157] S. K. Misra, M. Munoz-Ubeda, S. Datta, A. L. Barran-Berdon, C. Aicart-Ramos, P.
Castro-Hartmann, et al., "Effects of a Delocalizable Cation on the Headgroup of Gemini Lipids
on the Lipoplex-Type Nanoaggregates Directly Formed from Plasmid DNA,"
Biomacromolecules, vol. 14, pp. 3951-3963, 2013.
[158] C. Montis, P. Baglioni, and D. Berti, "Monitoring the interaction of nucleolipoplexes
with model membranes," Soft Matter, vol. 10, pp. 39-43, 2014.
[159] G. Tresset, "The multiple faces of self-assembled lipidic systems," PMC Biophysics,
vol. 2, pp. 1-25, 2009.
[160] J. C. Kuo, H. W. Huang, S. W. Tung, and Y. J. Yang, "A hydrogel-based intravascular
microgripper manipulated using magnetic fields," Sensors and Actuators a-Physical, vol. 211,
pp. 121-130, 2014.
[161] S. Tottori, L. Zhang, K. E. Peyer, and B. J. Nelson, "Assembly, Disassembly, and
Anomalous Propulsion of Microscopic Helices," Nano Letters, vol. 13, pp. 4263-4268, 2013.
[162] P. Tank, "The abdomen," in Grant's Dissector, 15th Edition, 15th ed Philadelphia:
LWW, 2012, p. 99.
[163] A. Ito, M. Shinkai, H. Honda, and T. Kobayashi, "Medical application of functionalized
128
magnetic nanoparticles," Journal of Bioscience and Bioengineering, vol. 100, pp. 1-11, 2005.
[164] Q. A. Pankhurst, N. T. K. Thanh, S. K. Jones, and J. Dobson, "Progress in applications
of magnetic nanoparticles in biomedicine," Journal of Physics D-Applied Physics, vol. 42,
2009.
[165] N. Tran and T. J. Webster, "Magnetic nanoparticles: biomedical applications and
challenges," Journal of Materials Chemistry, vol. 20, pp. 8760-8767, 2010.
[166] R. Tietze, S. Lyer, S. Durr, and C. Alexiou, "Nanoparticles for cancer therapy using
magnetic forces," Nanomedicine, vol. 7, pp. 447-457, 2012.
[167] R. Banerjee, Y. Katsenovich, L. Lagos, M. McIintosh, X. Zhang, and C. Z. Li,
"Nanomedicine: Magnetic Nanoparticles and their Biomedical Applications," Current
Medicinal Chemistry, vol. 17, pp. 3120-3141, 2010.
[168] K. Ishiyama, M. Sendoh, A. Yamazaki, and K. I. Arai, "Swimming micro-machine
driven by magnetic torque," Sensors and Actuators a-Physical, vol. 91, pp. 141-144, 2001.
[169] S. Martel, J. B. Mathieu, O. Felfoul, A. Chanu, E. Aboussouan, S. Tamaz, et al.,
"Automatic navigation of an untethered device in the artery of a living animal using a
conventional clinical magnetic resonance imaging system," Applied Physics Letters, vol. 90,
2007.
[170] D. M. Woolley, R. F. Crockett, W. D. I. Groom, and S. G. Revell, "A study of
synchronisation between the flagella of bull spermatozoa, with related observations," Journal
of Experimental Biology, vol. 212, pp. 2215-2223, 2009.
[171] P. Fischer and A. Ghosh, "Magnetically actuated propulsion at low Reynolds numbers:
towards nanoscale control," Nanoscale, vol. 3, pp. 557-563, 2011.
[172] O. J. Sul, M. R. Falvo, R. M. Taylor, S. Washburn, and R. Superfine, "Thermally
actuated untethered impact-driven locomotive microdevices," Applied Physics Letters, vol. 89,
2006.
[173] B. R. Donald, C. G. Levey, C. D. McGray, I. Paprotny, and D. Rus, "An untethered,
electrostatic, globally controllable MEMS micro-robot," Journal of Microelectromechanical
Systems, vol. 15, pp. 1-15, 2006.
129
[174] C. Pawashe, S. Floyd, and M. Sitti, "Multiple magnetic microrobot control using
electrostatic anchoring," Applied Physics Letters, vol. 94, 2009.
[175] B. Watson, J. Friend, and L. Yeo, "Piezoelectric ultrasonic resonant motor with stator
diameter less than 250 mu m: the Proteus motor," Journal of Micromechanics and
Microengineering, vol. 19, 2009.
[176] Y. Osada, H. Okuzaki, and H. Hori, "A POLYMER GEL WITH ELECTRICALLY
DRIVEN MOTILITY," Nature, vol. 355, pp. 242-244, 1992.
[177] V. Garcia-Gradilla, J. Orozco, S. Sattayasamitsathit, F. Soto, F. Kuralay, A. Pourazary,
et al., "Functionalized Ultrasound-Propelled Magnetically Guided Nanomotors: Toward
Practical Biomedical Applications," ACS Nano, vol. 7, pp. 9232-9240, 2013.
[178] S. Martel, O. Felfoul, and M. Mohammadi, "Flagellated Bacterial Nanorobots for
Medical Interventions in the Human Body," in 2008 2nd IEEE Ras & Embs International
Conference on Biomedical Robotics and Biomechatronics, 2008, pp. 264-269.
130
Appendix
Appendix A
A.1. The Matlab code to generate the ‘.gwl’ file of helices
A.2. The GWL file to fabricate one ABF from IP-L photoresist using DLW
%%%%% Head file %%%%%
PiezoScanMode
ContinuousMode
connectpointson
PointDistance 50
131
PerfectShapeOff
Timestamp 1
Monitoring 1
MeasureTilt 4
tiltcorrection 1
PowerScaling 1
LaserPower 20
Scanspeed 50
ResetInterface
%%%%% write one helix %%%%%
findinterfaceat 0
linenumber 1
scanspeed 25
laserpower 16
zoffset 0.5
xoffset 0
include one_helix.gwl
A.3. The code of GWL file of one helix:
write
0.000 2.500 4.750
0.069 2.696 4.741
0.139 2.891 4.716
0.208 3.082 4.673
0.278 3.270 4.614
0.347 3.451 4.539
0.417 3.625 4.449
0.486 3.791 4.343
0.556 3.946 4.224
0.625 4.091 4.091
0.694 4.224 3.946
0.764 4.343 3.791
0.833 4.449 3.625
0.903 4.539 3.451
0.972 4.614 3.270
1.042 4.673 3.082
1.111 4.716 2.891
1.181 4.741 2.696
1.250 4.750 2.500
1.319 4.741 2.304
1.389 4.716 2.109
1.458 4.673 1.918
1.528 4.614 1.730
1.597 4.539 1.549
1.667 4.449 1.375
1.736 4.343 1.209
132
1.806 4.224 1.054
1.875 4.091 0.909
1.944 3.946 0.776
2.014 3.791 0.657
2.083 3.625 0.551
2.153 3.451 0.461
2.222 3.270 0.386
2.292 3.082 0.327
2.361 2.891 0.284
2.431 2.696 0.259
2.500 2.500 0.250
2.569 2.304 0.259
2.639 2.109 0.284
2.708 1.918 0.327
2.778 1.730 0.386
2.847 1.549 0.461
2.917 1.375 0.551
2.986 1.209 0.657
3.056 1.054 0.776
3.125 0.909 0.909
3.194 0.776 1.054
3.264 0.657 1.209
3.333 0.551 1.375
3.403 0.461 1.549
3.472 0.386 1.730
3.542 0.327 1.918
3.611 0.284 2.109
3.681 0.259 2.304
3.750 0.250 2.500
3.819 0.259 2.696
3.889 0.284 2.891
3.958 0.327 3.082
4.028 0.386 3.270
4.097 0.461 3.451
4.167 0.551 3.625
4.236 0.657 3.791
4.306 0.776 3.946
4.375 0.909 4.091
4.444 1.054 4.224
4.514 1.209 4.343
4.583 1.375 4.449
4.653 1.549 4.539
4.722 1.730 4.614
4.792 1.918 4.673
4.861 2.109 4.716
4.931 2.304 4.741
5.000 2.500 4.750
5.069 2.696 4.741
5.139 2.891 4.716
5.208 3.082 4.673
5.278 3.270 4.614
5.347 3.451 4.539
5.417 3.625 4.449
5.486 3.791 4.343
5.556 3.946 4.224
5.625 4.091 4.091
5.694 4.224 3.946
5.764 4.343 3.791
5.833 4.449 3.625
5.903 4.539 3.451
5.972 4.614 3.270
6.042 4.673 3.082
6.111 4.716 2.891
6.181 4.741 2.696
6.250 4.750 2.500
6.319 4.741 2.304
6.389 4.716 2.109
6.458 4.673 1.918
6.528 4.614 1.730
6.597 4.539 1.549
6.667 4.449 1.375
6.736 4.343 1.209
6.806 4.224 1.054
6.875 4.091 0.909
6.944 3.946 0.776
7.014 3.791 0.657
7.083 3.625 0.551
7.153 3.451 0.461
7.222 3.270 0.386
7.292 3.082 0.327
7.361 2.891 0.284
7.431 2.696 0.259
7.500 2.500 0.250
7.569 2.304 0.259
7.639 2.109 0.284
7.708 1.918 0.327
7.778 1.730 0.386
7.847 1.549 0.461
7.917 1.375 0.551
7.986 1.209 0.657
8.056 1.054 0.776
8.125 0.909 0.909
8.194 0.776 1.054
8.264 0.657 1.209
8.333 0.551 1.375
8.403 0.461 1.549
8.472 0.386 1.730
8.542 0.327 1.918
8.611 0.284 2.109
8.681 0.259 2.304
8.750 0.250 2.500
8.819 0.259 2.696
133
8.889 0.284 2.891
8.958 0.327 3.082
9.028 0.386 3.270
9.097 0.461 3.451
9.167 0.551 3.625
9.236 0.657 3.791
9.306 0.776 3.946
9.375 0.909 4.091
9.444 1.054 4.224
9.514 1.209 4.343
9.583 1.375 4.449
9.653 1.549 4.539
9.722 1.730 4.614
9.792 1.918 4.673
9.861 2.109 4.716
9.931 2.304 4.741
10.000 2.500 4.750
10.069 2.696 4.741
10.139 2.891 4.716
10.208 3.082 4.673
10.278 3.270 4.614
10.347 3.451 4.539
10.417 3.625 4.449
10.486 3.791 4.343
10.556 3.946 4.224
10.625 4.091 4.091
10.694 4.224 3.946
10.764 4.343 3.791
10.833 4.449 3.625
10.903 4.539 3.451
10.972 4.614 3.270
11.042 4.673 3.082
11.111 4.716 2.891
11.181 4.741 2.696
11.250 4.750 2.500
11.319 4.741 2.304
11.389 4.716 2.109
11.458 4.673 1.918
11.528 4.614 1.730
11.597 4.539 1.549
11.667 4.449 1.375
11.736 4.343 1.209
11.806 4.224 1.054
11.875 4.091 0.909
11.944 3.946 0.776
12.014 3.791 0.657
12.083 3.625 0.551
12.153 3.451 0.461
12.222 3.270 0.386
12.292 3.082 0.327
12.361 2.891 0.284
12.431 2.696 0.259
12.500 2.500 0.250
12.569 2.304 0.259
12.639 2.109 0.284
12.708 1.918 0.327
12.778 1.730 0.386
12.847 1.549 0.461
12.917 1.375 0.551
12.986 1.209 0.657
13.056 1.054 0.776
13.125 0.909 0.909
13.194 0.776 1.054
13.264 0.657 1.209
13.333 0.551 1.375
13.403 0.461 1.549
13.472 0.386 1.730
13.542 0.327 1.918
13.611 0.284 2.109
13.681 0.259 2.304
13.750 0.250 2.500
13.819 0.259 2.696
13.889 0.284 2.891
13.958 0.327 3.082
14.028 0.386 3.270
14.097 0.461 3.451
14.167 0.551 3.625
14.236 0.657 3.791
14.306 0.776 3.946
14.375 0.909 4.091
14.444 1.054 4.224
14.514 1.209 4.343
14.583 1.375 4.449
14.653 1.549 4.539
14.722 1.730 4.614
14.792 1.918 4.673
14.861 2.109 4.716
14.931 2.304 4.741
15.000 2.500 4.750
Write
134
Appendix B
QCM-D analysis of the adsorption of various liposome formulations on Ti coated crystals and
poly-l-lysine (PLL) coating of ABFs.
Materials: 1,2-dioleoyl-sn-glycero-3-phosphocholine (DOPC), 1,2-di-(9Z-octadecenoyl)-sn-
glycero-3-phosphoethanolamine (DOPE), 1,2-dioleoyl-sn-glycero-3-phospho-L-serine
(DOPS), Poly(L-lysine)-g-poly(ethylene glycol) (PLL-PEG), Poly(L-lysine)-fluorescein
isothiocyanate (PLL-FITC), two layers of PLL/DOPS ((PLL/DOPS)2)
Figure C1. Normalized resonance frequency shift and dissipation shift of the 15 MHz detection frequency for the
adsorption of (a) DOPC liposomes (b) DOPC/DOPE (1:1 w/w) and (d) (PLL/DOPS)2/PLL-PEG. (c) Merged
fluorescent/transmission CLSM image of an ABF coated with PLL-FITC. Reused from [66].
135
CV
136
137
Publications
International Journal (Accepted/Published)
1. F. Qiu, B. J. Nelson, “Magnetic helical micro- and nanorobots: towards their biomedical
applications,” Engineering, Vol. 1, 2015, pp. 21-26.
2. A. M. Lindo, E. Pellicer, M. A. Zeeshan, R. Grisch, F. Qiu, J. Sort, et al., "The
biocompatibility and anti-biofouling properties of magnetic core-multishell Fe@C NWs-
AAO nanocomposites," Physical Chemistry Chemical Physics, Vol. 17, 2015, pp. 13274-
13279.
3. A. Servant, F. Qiu (Co-first author), M. Mazza, K. Kostarelos, B. J. Nelson，
“Controlled in vivo Swimming of a Swarm of Bacteria-Like Microrobotic Flagella,”
Advanced Materials, Vol, 27(19), 2015, pp. 2949-3091 (the front cover).
4. F. Qiu, S. Fujita, R. Mhanna, L. Zhang, B. Simona, B. J. Nelson，”Magnetic helical
microswimmers functionalized with lipoplexes for targeted gene delivery”, Advanced
Functional Materials, Vol. 25, 2015, pp.1666-1671 (the back cover).
5. T. Huang, F. Qiu, H. Tung, X. Chen, B. J. Nelson, M. S. Sakar, “Generating mobile
fluidic traps for selective three-dimensional transport of microobjects”, Applied Physics
Letters, vol. 105, 2014, 114102
6. T. Huang, F. Qiu, H. Tung, K. E. Peyer, N. Shamsudhin, J. Pokki, L. Zhang, B. J. Nelson,
M. S. Sakar, "Cooperative manipulation and transport of microobjects using multiple
helical microcarriers", RSC Advances, Vol. 4, 2014, pp.26771-26776 T.
7. F. Qiu, R. Mhanna, L. Zhang, Y. Ding, S. Fujita, and B. J. Nelson, "Artificial bacterial
flagella functionalized with temperature-sensitive liposomes for controlled release,"
Sensors and Actuators B: Chemical, vol. 196, 2014, pp. 676-681.
8. R. Mhanna, F. Qiu, L. Zhang, Y. Ding, K. Sugihara, M. Zenobi-Wong, and B. J. Nelson,
"Artificial Bacterial Flagella for Remote-Controlled Targeted Single-Cell Drug
Delivery," Small, vol. 10, 2014, pp. 1953-1957.
9. F. Qiu, L. Zhang, K. E. Peyer, M. Casarosa, A. Franco-Obregon, H. Choi, B. J. Nelson,
138
"Noncytotoxic artificial bacterial flagella fabricated from biocompatible ORMOCOMP
and iron coating," Journal of Materials Chemistry B, vol. 2, 2014, pp. 357-362 (the front
cover).
10. S. Kim, F. Qiu, S. Kim, A. Ghanbari, C. Moon, L. Zhang, B.J. Nelson, and H. Choi,
“Fabrication and Characterization of Magnetic Microrobots for Three-Dimensional Cell
Culture and Targeted Transportation,” Advanced Materials, Vol. 25, 2013, pp. 5863-
5868 (the front cover).
11. K. E. Peyer, S. Tottori, F. Qiu, L. Zhang, and B. J. Nelson, "Magnetic Helical
Micromachines," Chemistry - A European Journal, Vol. 19, December 2012, pp. 28-38
(the inside cover).
12. F. Qiu, L. Zhang, S. Tottori, K. Marquardt, K. Krawczyk, A. Franco-Obregón, and B. J.
Nelson, "Bio-inspired microrobots," Materials Today, Vol. 15, October 2012, pp. 463-
463 (the front cover).
13. S. Tottori, L. Zhang, F. Qiu, K. Krawczyk, A. Franco-Obregón, and B. J. Nelson,
"Magnetic Helical Micromachines: Fabrication, Controlled Swimming, and Cargo
Transport," Advanced Materials, Vol. 24, No. 6, February 2012, pp. 811-816 (the front
cover).
Conference
1. F. Qiu, R. Mhanna, L. Zhang, Y. Ding, Bradley J. Nelson, “Artificial bacterial flagella
functionalized with liposomes for potential targeted therapies”, The Annual Meeting of
Swiss Society for Biomedical Engineering 2014, Zurich, Switzerland, 27-28 August
2014, Poster and oral presentation.
2. S. Kim, S. Lee, S. Ha, S. J. Jung, F. Qiu, S. Yu, B. J. Nelson, and H. Choi, "Stem Cell
Transportation using Magnetically Actuated Scaffold-type Microrobots," Keystone
symposia on molecular and cellular biology, Engineering cell fate and functions (Z3),
California, USA, April 2014, poster presentation.
3. S. Kim, F. Qiu, S. Kim, A. Ghanbari, C. Moon, L. Zhang, B. J. Nelson, and H. Choi,
"Fabrication of Magnetically Actuated Microstructure for Targeted Cell Transportation,"
2013 IEEE International Conference on Nanotechnology, Beijing, China, August 2013,
139
oral presentation.
4. F. Qiu, R. Mhanna, L. Zhang, Y. Ding, K. Sugihara, M. Zenobi-Wong, and B. J. Nelson,
"Artificial Bacterial Flagella Functionalized with Temperature-sensitive Liposomes for
Biomedical Applications," Proc. in IEEE International Conference on Solid-State
Sensors, Actuators and Microsystems (Transducers 2013), Barcelona, Spain, June 2013,
pp. 2130-2133, poster presentation.
5. S. Kim, F. Qiu, L. Zhang, B. J. Nelson, and H. Choi, "Fabrication and control of
microrobots for targeted cell transportation," International Symposium on Nature
Inspired Technology, Kangwon, Korea, January 2013, oral presentation.
6. K. E. Peyer, F. Qiu, L. Zhang, and B. J. Nelson, "Movement of Artificial Bacterial
Flagella in Heterogeneous Viscous Environments at the Microscale," Proc. in IEEE/RSJ
International Conference on Intelligent Robots and Systems (IROS 2012), Vilamoura,
Portugal, October 2012, pp. 2553-2558, oral presentation (Best Student Paper Award).
7. F. Qiu, L. Zhang, S. Tottori, and B. J. Nelson, "Batch release and characterization of
artificial bacterial flagella using sonication," Proc. in 38th International Conference on
Micro & Nano Engineering (MNE 2012), Toulouse, France, September 2012, poster
presentation.
Journal covers
(2012) (2012) (2013)
140
(2013) (2014) (2015)
(2015)
