One Step Formation of Heterogeneous Tubular Biomaterials and Stimulus Responsive Hydrogels

by

Haotian Chen

A thesis submitted in conformity with the requirements for the degree of MASc – Biomedical Engineering

Institute of Biomaterials and Biomedical Engineering University of Toronto

© Copyright by Haotian Chen 2014

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One Step Formation of Heterogeneous Tubular Biomaterials and

Stimulus Responsive Hydrogels

Haotian Chen

MASc – Biomedical Engineering

Institute of Biomaterials and Biomedical Engineering University of Toronto

2014

Abstract

Synthetically prepared biomaterial tubular constructs have various applications including

engineered blood vessels, vascularized biomaterials as well as soft robotic actuators. Currently

there is no one-step method for the scalable preparation of perfusable tubular soft materials that

allows the tube size, morphology, composition and tensile properties to be consistently altered.

We report a scalable microfluidic approach for the consistent formation biopolymer tubes in

sodium alginate and collagen I, with outer diameters between 0.7 mm and 2.5 mm, at extrusion

velocities between 1mm/s and 20mm/s. We report the consistent formation of homogeneous

tubes with cylindrical, crimped, and corkscrew inner wall surfaces. Heterotypic tubular

constructs possess “Janus” single-layer, axial stripe patterns, and up to three concentric wall

layers. Stimulus responsive nanoparticles were incorporated as payload to induce predictive

shape transformations. This research potentially leads to high throughput construction of

complex spatially organized 3D functional soft materials and tissues.

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Acknowledgments This work is supported by the Natural Sciences and Engineering Research Council of Canada

(NSERC) and the NSERC CREATE Program in Microfluidic Applications and Training in

Cardiovascular Health (MATCH).

The authors thank Dr. Axel Guenther for project guidance, Lian Leng for teaching

microfabrication and biopolymer extrusion technique, Mark Jeronimo and Arianna McAlister for

their contribution to the work, and all group members from the Guenther lab for helping.

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Table of Contents Acknowledgments.......................................................................................................................... iii

Table of Contents ........................................................................................................................... iv

List of Tables ................................................................................................................................. vi

List of Figures ............................................................................................................................... vii

1 INTRODUCTION AND BACKGROUND ................................................................................1

1.1 Formation of tubular constructs ...........................................................................................2

1.2 Solidification of biopolymers ..............................................................................................9

1.3 Introducing heterogeneity in biomaterials .........................................................................11

1.4 Stimulus responsive materials............................................................................................13

2 CONTRIBUTIONS ..................................................................................................................17

3 MOTIVATION .........................................................................................................................17

4 SPECIFIC AIMS .......................................................................................................................18

4.1 One-step formation of homogeneous soft material tubes ..................................................18

4.2 One-step formation of heterogeneous soft material tubes and stimulus responsive tubes ...................................................................................................................................19

4.3 Stimulus responsive hydrogel characterization and extrusion ...........................................19

5 MATERIALS AND EXPERIMENTAL...................................................................................21

5.1 Microfluidic device designs ...............................................................................................21

5.2 PDMS Device Fabrication .................................................................................................25

5.3 Homogeneous Tube Extrusion ...........................................................................................26

5.4 Multilayer Tube Extrusion .................................................................................................27

5.5 Janus Tube Extrusion .........................................................................................................28

5.6 Preparation of Stripe-Patterned Hydrogel Tubes ...............................................................28

5.7 Collagen Tube Extrusion ...................................................................................................29

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5.8 Valves Control Board Design ............................................................................................30

5.9 Labview Control Programs ................................................................................................31

5.10 Tube Dimensions Measurements .......................................................................................33

5.11 Polymer Synthesis and Hydrogel Solution Preparation .....................................................34

5.12 Stimulus Responsive Material Characterization ................................................................34

5.13 SEM Imaging of Tubular Soft Material Constructs ...........................................................35

5.14 Environmental Temperature and pH Control of Stimulus Responsive Hydrogels ............35

6 RESULTS AND DISCUSSION ...............................................................................................36

6.1 One-step formation of homogeneous soft material tube ....................................................36

6.2 Predictive control of tube dimensions................................................................................40

6.3 Soft material tubes of non-circular cross section ...............................................................46

6.4 One-step formation of heterogeneous soft material tube ...................................................50

6.5 Collagen tube formation ....................................................................................................55

6.6 Stimulus responsive tubular structures ..............................................................................58

7 SUMMARY ..............................................................................................................................68

8 REMAINING WORK AND FUTURE DIRECTION ..............................................................69

References ......................................................................................................................................70

Appendices .....................................................................................................................................75

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List of Tables Table 1. Tubular structure formation summary .............................................................................. 7

Table 2. Summary of synthetic heterogeneities in biomaterials ................................................... 13

Table 3 Recipe of stimuli-responsive particles synthesis ............................................................. 59

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List of Figures Figure 1. Microfluidic device designs. ......................................................................................... 23

Figure 2. Experimental setup of homogeneous tube extrusion. .................................................... 27

Figure 3. Experimental setup of spotted/striped tube extrusion. .................................................. 29

Figure 4. Valve control circuit. ..................................................................................................... 31

Figure 5. Labview control programs. ............................................................................................ 33

Figure 6. Device design and experimental setup. ......................................................................... 38

Figure 7. Dimension control of homogeneous tubes. ................................................................... 44

Figure 8. Homogenous tube morphologies. .................................................................................. 48

Figure 9. Heterogeneous tubular structures. ................................................................................. 53

Figure 10. Collagen tubes ............................................................................................................. 57

Figure 11. Stimuli-responsive PNIPAM-VAA nanoparticles. ..................................................... 59

Figure 12. PNIAPAM-VAA particle size and its stimuli-responsive volumetric changes. ......... 61

Figure 13. Schematic of homogeneous stimulus responsive sheet extrusion. .............................. 64

Figure 14. Homogeneous stimuli-responsive hydrogel and its response to temperature and pH. 64

Figure 15. Schematic of stimulus responsive tube assembly. ....................................................... 66

Figure 16. Temperature induced response of half-half stimulus responsive tube. ....................... 67

Figure A1. Tube extruded with no crosslinker in inner streaming fluid. ...................................... 75

Figure A2. Spotted tubes............................................................................................................... 76

Figure A3. Comparison of dimensions of single layer/Three-layer tubes. ................................... 77

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Figure A4. Heterogeneous hydrogel extrusion. ............................................................................ 80

Figure A5. Heterogeneous stimuli-responsive hydrogel sheets and their response to temperature

and pH. .......................................................................................................................................... 81

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1 INTRODUCTION AND BACKGROUND In nature, tubular structures are a common element in both plant and animal structure and

physiology. In the human body and human organs, they present in the forms of blood vessels,

airways, and ducts. Artificial biomaterial tubes were produced to mimic or replace natural

existing tubular structures. They can find applications as engineered blood vessels and micro

tissues[1], and cell-laden tubes incorporated into vascularized biomaterials [2]. Although various

efforts have been made to achieve controlled formation of tubular structures of different

materials including biomaterials, the goals of high throughput production, predictive controlled

dimensions, spatial compositional heterogeneity, flexibility of biomaterials and compatibility

with cells have not been met.

Tubular tissues in the human body exist at different scales. They can be as small as capillaries (5-

10µm in diameter) and gland ducts, and as large as the aorta (diameter approximately 30mm in

human) and bronchus (diameter approximately 25mm in human). Present approaches of

biomaterial tube formation often lack of scalability due to the small dynamic diameter range

accessible for tube formation and the inability to predictively control and maintain tube

diameters of formed tubes. Current methods are limited in terms of their compatibility with

different soft materials due to constrain on gelation time. A scalable method with precise

predictive control over dimensions of formed tubular structures, and with flexibility on material

choices, is an essential requirement towards the formation of engineered tubular tissues [1, 2].

Human tubular tissues such as bronchioles, intestinal mucosa, arteries, and blood vessels rely on

a unique spatial organization that stretches from the molecular alignments of biopolymers in the

extracellular matrix, over cells to the tubular diameter and is necessary for tissue function [3].

For instance, bronchioles and intestinal mucosa have non-circular inner walls; arteries and veins

have multiple cell layers; bronchi have gland ducts and cartilage around its circumference. This

heterogeneity in spatial composition and special morphologies cannot be mimicked by the

current tubular structure formation techniques, especially in the context of biomaterials. High

throughput formation of heterogeneous tubular structure made of biomaterials has been an

important step towards vascular graft and drug testing developments.

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To mimic biological function, nature requires that engineered tissues be responsive and adaptive

in different environments, and adjust their function accordingly. For example, blood vessels

regulate blood flow in response to temperature and several other stimuli to maintain homeostasis.

Responsive polymer materials with very similar attributes have been used in a broad range of

applications; responsive biointerfaces [4], controlled drug delivery [5], interactive coating [6]

and to mimic muscle actions [7]. They have been applied to regulate transport of ions and

molecules, alter wettability and adhesion of different species or convert between chemical

stimuli and optical, thermal and mechanical responses. Currently, the synthesis and preparation

of stimulus responsive materials involves multiple steps that are often costly and time

consuming. One-step formation of patterned soft stimulus responsive materials will improve the

production throughput significantly. Incorporating stimulus responsive material in hydrogel tube

can induce predictive stimulus response of the formed tube with known environmental changes,

and these tubes can find potential applications in controlling rate and direction of flow in

vascularized hydrogel scaffold.

1.1 Formation of tubular constructs Two approaches have been widely employed for the formation of soft material tubes: extrusion

and molding.

The Takeuchi group at the University of Tokyo demonstrated the extrusion of meter-long cell-

laden microfibers [8]. A microfluidic device with double coaxial laminar flow was used to

produce the core-shell hydrogel microfibers. This extrusion method is referred as in-plane

extrusion in my thesis because the extrusion direction is in-plane with flows in the microfluidic

device. The composition of the three co-axially flowing biopolymer solutions from inside out

was: a cell-laden biopolymer solution at the core, a fast gelling shell biopolymer solution and a

crosslinker solution. The shell biopolymer solution was ionically cross-linked while the solution

at the core was cross-liked by temperature-induced gelation. Rapid gelation produced a hydrogel

shell immediately after exiting the device with a cell-laden biopolymer solution confined in the

core. The fiber was then transferred to 37°C environment for 15 minutes for the temperature-

induced gelation of the core. Compatibility of the approach with the acid-solubilized type I

collagen, pepsin-solubilized type I collagen and fibrin as core material was demonstrated. The

extruded microfibers had cores of cell-laden collagen and shells of alginate or alginate-agarose

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interpenetrating networks. In the case of endothelial cells introduced in the core, the cells aligned

themselves to form a tubular structure. The microfibers had diameters between 92±5µm and

210±5µm.

There are several other examples demonstrating extrusion of tubular structure using an in plane

extrusion method [9-11]. Examples of predictable control of the dimensions of formed tubular

structures have been demonstrated for tubes with outer diameters from 300µm to 900µm.

However, the tubes were not made with biocompatible materials[9]. The tube materials used by

Luo’s group are polyacrylonitrile, polysulfone and polystyrene. In-plane extrusion of coaxially

arranged fluid streams was utilized and solidification was achieved due to solvent loss. This

solidification method is not suitable for hydrogel formation because hydrogels are required to

remain hydrated even after solidification. The inner diameter, outer diameter and the wall

thickness of the formed tubular structure were controlled by changing the flow rates of individual

fluid streams in the coaxial flow. Keeping the flow rates of the other fluids constant, increasing

the inner fluid flow rate resulted in increased inner and outer tube diameters. Keeping the flow

rates of the other fluids constant, increasing the outer fluid flow rate resulted in a decrease in the

tube diameters. The immiscible outer and inner fluid streams extracted the solvent from the

polymer solution contained in the middle polymer stream. An analytical model was presented

based on the Navier Stokes equations, and allowed the tube dimensions to be accurately

predicted. The model neglected the solidification of the tube wall and the associated increase in

viscosity.

Other examples demonstrating ink-jet printing of alginate based tubular structures lack

predictable tube dimensions [10, 11]. Kawakami’s group [10] demonstrated the in-plane

extrusion of homogeneous alginate tubes with outer diameters between 260µm and 310µm.

Ozbolat’s group [11] demonstrated the formation of homogeneous hydrogel tubes in alginate and

chitosan, with outer diameters between 600µm and 1300µm. In both works, the dimensions of

formed biopolymer tubes were predicted based on empirical data. Heterogeneity of spatial

composition is very difficult to achieve with in-plane extrusion of tubular structures.

Nakajima’s group demonstrated out-of-plane extrusion [12]. In this approach, the alginate tube

flowed perpendicular to the direction of alginate solution flow in a microfluidic device using a

2D array of microfabricated nozzles. The nozzles had square shapes and the center regions of the

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nozzles were solid. The alginate solution was extruded along the edge of the square nozzles. The

alginate solution was extruded into a flow of crosslinker solution that approached in a direction

perpendicular to the alginate solution flow. The alginate solution flow then changed its direction

and flowed parallel with the crosslinker flow. The outer diameter of the formed tubes was

approximately 120µm.

The off-plane tubular structure extrusion technique that is further developed in this thesis was

originally demonstrated by Arianna McAlister [13]. The designs of microfluidic devices

introduced in this thesis evolved from her device designs. Homogeneous tube extrusion was

accomplished using three layer PDMS devices; the three layers, from top down, are an inner

crosslinker solution layer of calcium chloride, an alginate matrix solution layer and an outer

calcium chloride crosslinker solution layer. The channels in the device radially feed the solutions

towards a center outlet, which directs the layered fluid towards the common extrusion outlet.

The biopolymer solution is sandwiched by the crosslinker solutions, which forms a cylindrical

wall and moves downwards. A biopolymer tube with outer diameters between 600µm and

1600µm was formed and passed through a confinement tube before entering a reservoir filled

with crosslinker solution. The dimensions of the formed alginate tubes were controlled by

changing the flow rates of the streaming solutions and the biopolymer solution. Extrusions of

bilayer tubes, Janus tubes and tubes that had secondary biopolymers incorporated as spots along

the tube circumference were demonstrated. The microfluidic device designs presented in this

thesis are further developments of the ones originally presented by Arianna McAllister. The

original device designs are more challenging to fabricate with a high yield using multilayer soft

lithography. The center extrusion outlet shared the exact diameter of the tool used for punching

the hole, requiring perfect alignment during a manual punching process. Misalignment lead to

accidental shortening of microfluidic channels around the center extrusion hole, which resulted

in an imbalance of the flow resistance around the extrusion outlet and prevented the consistent

formation of intact tubes at experimental yields exceeding 20%. An analytical model was

presented that described the tube formation process by neglecting the influences of crosslinking

and solidification of the biopolymer during extrusion process. The experimental results only to a

limited extent agreed with model predictions. Bilayer and Janus tube extrusion required more

complex device, but given the low yield of device fabrication, heterogeneous tubes could not be

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formed consistently. The formation of heterogeneous tubes and their dimensions could not yet be

consistently controlled.

The other common method of forming tubular strucutres is molding. Sacrificial materials are

commonly used to form a perfusable vascular structure inside a scaffold. In one example

demonstrated by Christopher Chen’s group, a rigid 3D filament network of carbohydrate glass

was used as sacrificial template in engineered tissues to generate vascular network [14]. The

dimension of the vasculature was defined by the sacrificial material dimensions, and the limiting

dimensions of the filament network was the nozzle size of their 3D printer. A simple model,

which did not consider coaxial flow, helped predict with reasonable accuracy the diameters of

the extruded filaments. The diameter of the filament varied between 200µm and 1100µm. A

suspension of cells in extra cellular matrix (ECM) was poured on to the filament network and

was allowed to solidify. The filament was then dissolved, leaving behind an elliptical flow

conduit within the bulk gel. The scaffold was then perfused with blood under positive pressure

pulsatile flow. An advantage of this multistep strategy is its flexibility in terms of ECM material,

as it is not limited in terms of the crosslinking time. The scaffold can also be made of agarose,

alginate, fibrin, matrigel and poly(ethylene glycol) (PEG), and a wide range of cell types can be

incorporated into the scaffold. The perfusable network was not exactly a vascular system because

the materials made of the channels are the same as the bulk gel. The cultured tissues are therefore

homogeneous in spatial composition, and cannot mimic the complex tissue interactions in natural

organisms.

In other examples, extruded endothelial cell laden alginate fibers were embedded in a collagen

scaffold [15, 16]. Fibers composed of a mixture of alginate and gelatin served as the sacrificial

material. The hydrogel fibers were produced by extrusion using the in-plane method which was

discussed above. The diameters of the fibers can be controlled to a certain extent by controlling

the size of extrusion needle and flow rate of biopolymer solution, with a range of diameters from

150µm to 1200µm. An empirical relationship between biopolymer flow rates, extrusion needle

size and hydrogel fiber diameter was presented. The produced alginate/gelatin fibers were coated

with smooth muscle cells and were embedded in collagen solution, which solidified to form a

bulk gel after thermal gelation. The alginate/gelatin fibers were dissolved by alginate lyase at the

end to leave channels in the bulk collagen gel in an approach similar to the one used by Chen’s

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group [14]. In this instance, alginate/gelatin hydrogel fibers served as a sacrificial material and a

homogenous scaffold with embedded vasculature was produced using this multi-step approach.

A technique developed by Jennifer Lewis’s group demonstrated the capability of introducing

spatial composition heterogeneity into vascularized scaffold [17]. Two types of sacrificial

materials, fugitive Pluronic F127 and gelatin methacrylate (GelMA) were used to define the

channels and were thermally gelled. A vascular network made of fugitive ink was printed with a

nozzle such that the diameter of the printed filaments could be controlled by the printing

pressure, nozzle height and nozzle movement speed. The dimensions of the printed filaments

were controlled based on empirical data. The diameter of the filaments varied from 150µm to

650µm. After the vascular network was printed, a bulk volume of GelMA ink was cast around it.

As the temperature was lowered below 4°C, the GelMA ink formed a solidified bulk gel and the

fugitive ink returned to a liquid state and was removed from the bulk gel. A perfusable vascular

network remained inside the bulk gel. A heterogeneous vascular network was created by co-

printing fugitive ink and cell-laden GelMA. A bulk GelMA ink was then cast around the printed

network. GelMA inks were then photopolymerized to form a bulk gel with an internal gelled

cell-laden vascular network. The fugitive ink was removed by reducing the temperature. At the

end, heterogeneous vascular network was formed. However, the part of network formed by cell-

laden GelMA ink was not perfusable because GelMA ink cannot be removed after

photopolymerization. This approach provides some spatial composition heterogeneity to the bulk

gel and the surroundings of the tubular structure.

The previous molding approaches produced vascularized scaffold but all involve the use of

sacrificial materials to produce a vascular network. These approaches limit the control of tubular

structure dimensions due to the limited accuracy of sacrificial layer printing. Another technique,

in which the bulk mold was produced first, effectively improved the control over vascular

network dimensions [18]. Polydimethylsiloxane (PDMS) devices containing array of channels

were produced using standard soft lithography with channel sizes between 50µm and 150µm.

The channels were then filled with fibrin/collagen solution and allowed to gel. Human umbilical

vein endothelial cells were coated at the two ends of the channels, and after 3 to 4 days of

culture, the cells were able to migrate into the gel and form capillaries. The yield of the entire

process is 50%. The major limitation of this approach is the length of tubular structure being

restricted to 1600µm.

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Table 1 summarizes the dimensions of tubular structures made by various methods, including the

ones discussed previously.

Table 1. Tubular structure formation summary

Reference Method Materials Outer

Diameters(µm)

Dimension

Control

Onoe, H., et

al. [8]

In-plane extrusion Alginate,

Alginate/Agarose

92-210 N/A

Lan, W., et

al. [9]

In-plane extrusion Polyacrylonitrile,

Polysulfon,Polystyrene

300-900 Analytical

Takei, T., et

al. [10]

In-plane extrusion Alginate 260-310 Empirical

Zhang Y.,

Y.Y.[11]

In-plane extrusion Alginate 600-1200 N/A

Hwang,

S.W., et

al.[19]

In-plane extrusion Poly(lactic co-glycolic

acid) (PLGA)

20-230 Empirical

Jeong, W., et

al. [20]

In-plane extrusion Acrylic Acid 20-90 Analytical

Choi, C.-H.,

et al. [21]

In-plane extrusion Poly(ethylene glycol)

diacrylate (PEG-DA)

55-75 N/A

Cho, S.,

T.S.,et al.

[22]

In-plane extrusion PEG-DA 70-140 N/A

Dittrich, P.S.,

et al. [23]

In-plane extrusion 1,2-dilauroyl-sn-glycero-

3-phosphocholine

3.5-20 N/A

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Hu, M., et al.

[24]

In-plane extrusion gelatin-

hydroxyphenylpropionic

acid (Gtn-HPA), poly (N-

isopropyl acrylamide)

(PNIPAAM), Alginate

111-227 Analytical

Su, J., et al.

[25]

In-plane extrusion Alginate 1-10 Analytical

Iwasaki, N.,

et al. [26]

In-plane extrusion Alginate,

Alginate/Chitosan

28.6-31.3 N/A

Cheng, Y., et

al. [27]

In-plane extrusion Alginate 40-120 N/A

Sugiura, S.,

et al. [12]

Off-plane

extrusion

Alginate 120 N/A

McAllister,

A. [13]

Off-plane

extrusion

Alginate 600-1600 Analytical

Raof, N.A.,

et al. [28]

Off-plane

extrusion

Alginate 30-300 N/A

Miller, J.S.,

et al. [14]

Molding Agarose, Alginate, Fibrin,

Matrigel, PEG

200-1100 Analytical

Sakai, S., et

al. [16]

Molding Collagen 150-1200 Empirical

Kolesky,

D.B. [17]

Molding GelMA 150-650 Empirical

Yeon, J.H., et

al. [18]

Molding Fibrin, Collagen 50-150 Analytical

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1.2 Solidification of biopolymers There are various means to trigger solidification of polymer solutions, including temperature,

pH, electric or magnetic field, catalyst, photopolymerization, solvent extraction and ionic

concentration gradient [29, 30]. These solidification processes can be reversible or irreversible.

Sol-gel transition of biomaterials is widely used in applications in tissue graft, immobilization of

biological materials and cells, and biosensor production [29, 31-33]. In the scenario of

biopolymer soft material extrusion, the extruded biomaterials should keep their mechanical

strength for a reasonably long period and should be tolerant to external environmental change to

a certain extent. Due to the need of high throughput extrusion, the reaction time for solidification

of biomaterials in extrusion process is limited. Therefore, an irreversible rapid gelation process

will be ideal for soft biomaterial extrusion.

As discussed previous, alginate is the candidate best fit in the context of soft biomaterial

extrusion process. It was used by many research groups as extrusion material for microfibers and

hydrogel tubes [34, 35]. The main advantage of using alginate is its rapid gelation process. The

sol-gel transition of alginate is induced by ionic crosslink. Rapid ion exchange kinetics results in

quick crosslink reaction. Alginates are block copolymers consist of α-L guluronic acid (G) and β-

D-mannuronic acid (M). Homopolymeric regions of M and G blocks, and interspersed regions of

alternating M and G are present in alginate. Gelation of alginates is originated from the affinity

of alginates towards divalent metal ions such as Ca2+. Calcium ions selectively bind G blocks of

two alginate chains. The process is irreversible and almost instant. The rate of gelation reaction

is dominated by relative diffusion rate of calcium ions in alginate solution [36]. Therefore, the

gelation rate can be predicted by modeling the system as diffusion limited reaction system. The

alginate gel properties strongly depend upon the length of G blocks, monomeric composition,

molecular size and concentration of alginate molecules, as well as the concentration of calcium

ions [37]. Moreover, alginates have been proved to be compatible to culture of many cell types.

They have already been widely used in applications in the fields of tissue engineering, cell

immobilization and drug delivery.

Collagen is another ECM biomaterial which is commonly used as 3D tissue culture scaffold. It is

a better material comparing to alginate in terms of mechanical strength and cell adhesion [38].

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Collagen attracts great interest from biomedical engineers for several reasons: i) Collagen is one

of the major structural materials of animal bodies. ii) It produces tissues with a wide range of

mechanical strengths based on its organizational modulation, such as bone, cartilage, tendon,

skin, intervertebral disk and ligament. iii) Load bearing structures made of collagen covers a

huge range of sizes. Blue whale Balaenoptera musculus has the length of 30m while female

Paedocypris micromegethes has the length of 0.008m. iv) Collagen-related diseases largely

impact life quality, for instance, intervertebral disk degeneration.

As discussed before, thermal gelation of collagen can only be used in molding of collagen gel

but not extrusion processes. It is because thermal gelation of collagen is a very slow process

which takes a minimum of 15 minute [8]. In hydrogel extrusion systems, the coaxial flow of

biomaterial solution and crosslinker solution cannot be maintained for such a long period.

In vivo, collagen present in the form of procollagen consisting of a right-handed helical region

which is flanked by two non-helical propeptides. The propopetides take large space around

collagen molecules and keep them apart, hence inhibit self-assembly of collagen molecules [39].

After cleaving the propetides from the tropocollagen using enzymes, spontaneous self-assembly

of highly organized fibers will start [40]. It is believed that some physical factors also exert

significant effect on fibrilogenesis via liquid crystal phasing of collagen [41]. Physicochemical

effect drives collagen fiber formation and molecular crowding plays an important role in this

process. Molecular crowding refers to volume excluded by one soluble molecule to another.

Molecular crowding produces molecule confinement such that the volume occupied by one

molecule sets a fixed boundary to another. In the crowded scenario, molecules with high aspect

ratio such as collagen tend to self-align along the long axial direction. This effect is driven by the

fact that the translational entropy gain from aligning the molecules is greater than the rotational

entropy loss [42]. The free energy of the isotropic state is therefore lower than the anisotropic

state. The system where the collagen molecules are self-assembled is more stable [43].

Molecular crowding condition can significantly accelerate the gelation of collagen therefore

reduce the gelation time. By taking advantage of the effect of molecular crowding, extrusion of

collagen gel can be achieved. Crowded solutions of PEG or hyaluronic acid (HA) can induce

molecule crowding condition to collagen solution. Using PEG or HA solutions as coaxial flow

with collagen solution can potentially shorten the time required for collagen gelation, and makes

collagen gel extrusion possible.

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1.3 Introducing heterogeneity in biomaterials Spatial composition heterogeneities usually present in natural tissues but are hardly achieved by

synthetic methods especially in the context of soft biomaterials. Current tissue engineering

technologies cannot mimic many of the complex organs with heterogeneous composition where

the organ function largely relies on interaction among different cell types. As simple as tissues,

heterogeneous structures also present. For instance, blood vessel is a multilayer tubular structure

where the layers are made of collagen and elastin respectively.

Heterogeneities can be introduced into tubular structures in the axial, radial and circumferential

directions. Axial heterogeneity refers to the compositional change along the length of the tubular

structure. Radial heterogeneity refers to multilayer tubular structure with a material difference

among layers. Circumferential heterogeneity refers to the presence of multiple materials around

the circumference of the cross section perpendicular to the length of the tube. Radial, acial and

circumferential heterogeneities individually has been demonstrated by using both extrusion and

molding methods for tubular structure formation. Incorporating all types of heterogeneities using

a single platform remains as a challenge for producing soft material tubes.

Heterogeneities have been introduced to biomaterials in various methods. The basic techniques

are similar to those producing homogeneous tubular structures discussed before. Extrusion and

molding are the two main methods.

Leng et al. have demonstrated a technique to introduce heterogeneity in biomaterials in a

continuous extrusion process[44]. A microfluidic device was used to incorporate one or multiple

secondary biopolymer solutions into a layer of a base polymer. The spatial arrangement of the

secondary biopolymer retains after leaving the microfluidic device by fast gelation of the base

polymer. Alginate solution was pumped into the microfluidic device by a syringe pump. It

formed hydrogel sheet immediately after exiting from the device and exposing to calcium

chloride solution. After exiting the device, the formed hydrogel sheet was collected by a rotating

drum. The thickness of the hydrogel sheet formed can by controlled by both the infusion rate of

the pump and the speed of the collecting drum. Computer-controlled solenoid valves initiated the

outflow of secondary biopolymers from the on-chip reservoirs. These secondary biopolymers are

loaded on the hydrogel sheet made of a primary biopolymer in the form of spots. The position

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and size of the spots can be controlled by programming the opening sequence and duration of the

solenoid valves. The heterogeneity can be formed along both lateral and axial directions.

Other examples extruding heterogeneity in hydrogel fibers have been demonstrated [22]. The

methods are based upon the stop-flow lithography technique developed by the Doyle group.

Heterogeneous poly(ethylene glycol) diacrylate (PEGDA) fibers were successfully produced

using extrusion method. Mosaic patterned microfibers were produced using a microfluidic

device. The spatial composition of the polymer solutions were controlled by microfluidic

channels prior to photo crosslink of the polymers. Fibers with width of 200µm to 1000µm and

thickness of 70µm to 140µm can be extruded. The size and aspect ratio of the fiber was

controlled by varying the fractional volume of PEGDA solution with respect to carrier liquid

flows. The carrier liquid flows were inert PEG solutions extruded from the device and enclosed

PEGDA solution from top and bottom. This method produced axial and lateral heterogeneities to

extruded fibers but cannot be directly transferred to formation of extrusion of heterogeneous

tubular structures.

An example demonstrated formation of tubular structure with heterogeneity along the radial

direction using molding [45]. Different cells were used to cover the surface of a stretched

Polydimethylsiloxane (PDMS) layer. The stretched layer was then rolled up to form a multilayer

tubular structure with different cells in different layers. The cell layers were not in direct contact

with each other due to the separation form the PDMS layer. The tube diameters can be controlled

between 100µm to 2mm. The predictive control of tube dimensions was based on empirical data.

Control over the tube diameter is based on changing the spinning speed, which determines the

thickness of PDMS layer. The tubes had finite length due to the limitation of the size of casted

PDMS layer. The multi-steps nature of the process limits the throughput of tubular structure

production.

A multilayer tubular structure has been successfully molded using collagen [46]. The

heterogeneity of the tubular structure is along the radial direction. The collagen tubular structures

were molded by using a pipe with a needle in the center. The second layer of tube was produced

by seeding endothelium on the inner wall of the formed collagen tube. The collagen tube

produced had the outer diameter of 830 µm.

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Table 2 summarizes the heterogeneities been introduced to different biomaterials, including the

ones discussed previously and some others.

Table 2. Summary of synthetic heterogeneities in biomaterials

Reference Method Structure Heterogeneity

Onoe, H., et al. [8] In-plane extrusion Fiber Radial

Dittrich, P.S., et al.

[22]

In-plane extrusion Fiber Axial, Lateral

Leng, L., et al. [44] In-plane extrusion Sheet Axial, Lateral

Cheng, Y., et al. [27] In-plane extrusion Tube Radial,

Circumferential

Yuan, B., et al. [45] Molding Tube Radial

Takei, T., et al. [46] Molding Tube Radial

1.4 Stimulus responsive materials As mentioned before, stimulus responsive materials have many valuable applications. One

example of the applications of stimulus responsive materials will be reconstructable surface

which involves formation of stimuli-responsive thin films which macromolecules are grafted

chemically to a surface at sufficiently high grafting density. The responsive behavior of these

thin films originates from the properties of the grafted polymer chains. Poly(N-

isopropylacrylamide) (PNIPAM) thin film that possess a lower critical solution temperature

(LCST) undergo a phase transition in response to temperature [47]. Polyelectrolyte thin films

respond with large conformational changes to pH [48]. Some zwitterionic thin films possess an

upper critical solution temperature changes their wetting behavior with temperature[49]. These

responsive surfaces can be used for biointerfaces and bioseparation. The possibility of switching

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adhesion between stimuli-responsive materials and proteins and cells has been explored for

control of cell and protein adhesion [50, 51].

Other common applications of stimuli-responsive materials are micro- and nano-actuators and

sensors [52]. These applications are based on the swelling and de-swelling, wetting and de-

wetting, or adsorption and desorption of organic suface layers. An example will be use internal

stresses caused by conformational transformations within thin film layers to design temperature

sensitive microcantilevers by grafting PNIPAM [53].

Moreover, stimuli-responsive materials can be used to control catalyzed chemical reaction and

drug delivery. By encapsulate catalyst particles in a thermoresponsive polymer shell, temperature

dependent swelling and shrinking of the shell can alternatively expose and hide the catalyst and

hence control chemical reaction rate [54]. Nanosized capsules could store and protect drugs and

release them inside cells after the capsule has been internalized [55]. Stimuli-responsive carries

of drugs attract a great attention because of its broad opportunities of controlled drug delivery.

Temperature and pH are common triggers of stimulus responsiveness in both nature and

synthetic materials. In the context of stimulus responsive tubular structure, a stimulus responsive

tube can regulate its diameter hence the flow rate through the tube by reacting to the pH or

temperature change of the fluid inside the tube. The working mechanism is similar to blood

vessels which regulate the flow rate of blood by changing their diameters.

Temperature responsive polymers are polymers that exhibit a drastic and discontinuous change

of their physical properties with temperature. PNIPAM is a well-known thermoresponsive

polymer. It has a LCST around 32 ̊C. The response generated by PNIPAM is due to the change

of inter and intra molecular interactions between its pendant groups and water molecules.

PNIPAM exhibits hydrophilic properties due to the hydrogen boding between the pendant group

and water molecules. When the temperature is greater than the LCST, the bonding breaks and the

polymer becomes hydrophobic [56]. The polymer collapses and thus shrinks in volume in

temperature higher than its LCST. It converts the thermal signal into a mechanical signal.

PNIPAM based particles are chosen for secondary biopolymer doping due to its excellent

responsiveness to temperature changes. Its LCST is around 32 ̊C which is in the temperature

range suitable for most biological system. The process of polymerization of PNIPAM particle is

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suitable to copolymerize many additives. Therefore the particles can be functionalized to have

more complicated stimuli-responses.

A pH responsive polymer is a material which will respond to the changes in the pH of the

surrounding medium by varying their dimensions. Many pH responsive polymers have acidic

groups which swell in basic pH [57]. As the pH in the solution increases, acidic groups such as

carboxylic group releases hydrogel ion and ionized to carry a negative charge. As the polymer

possesses many similar charged groups repelling each other, it swells and expands in

dimensions. Polymer of vinyl acetic acid (VAA) is rich in carboxylic group hence possesses pH

responsiveness. VAA group has a relatively low reaction ratio comparing to NIPAM [58].

Therefore, it can form a shell encapsulate PNIPAM when added during the polymerization of

PNIPAM. By functionalize PNIPAM particles with VAA, the resulting particle will become pH

responsive.

Stimuli-responsive materials are often patterned in order to better serve their purposes in varies

systems. For example, stimuli-responsive patterns can control the flow in microchannels as

pumps, valves, or mixers through their volumetric actuation or the change of surface properties

[59]. There are two main categories of stimuli-responsive material patterning techniques. The

first category will be photolithograph[60]. A selective UV light exposure through a photomask is

used for patterning. UV exposure selectively crosslink or cleave polymer. This technique is

usually used for making 2D patterns. For 3D patterning, multiphoton lithography can be

used[61]. This method selectively crosslink or cleave polymer on the focal spot of light.

The second category is the micromolding techniques. Replica molding makes use of thermal

crosslinking of prepolymer filling the cavity of the master mold. UV molding use UV to

crosslink prepolymer filling the cavity of the master mold[62]. Nano-imprinting technique makes

patterns by pressure-induced deformation of polymer above glass transition temperature[63].

Capillary force lithography takes advantage of capillary rise of polymer in the cavity of the

master mold [64].

These techniques can produce micro and nanoscale patterns of stimuli-responsive materials.

However, due to the requirement of photomasks or master molds, these methods can be

expensive especially for small quantity productions. As making photomasks and master molds

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are expensive and time consuming, these methods are not tolerable to frequent design changes.

They are not cost effective for prototyping and laboratory uses which produce small quantity and

are projected to frequent design changes. Moreover, as patterning process is combined with

polymerization process, the flexibility of chemistry and properties of stimuli-responsive

materials are limited due to constricted reaction conditions. The patterning process can be time

consuming because of involvement of polymerization. In my thesis, the one-step formation of

heterogeneous stimulus responsive 3D soft material can effectively reduce time cost of

patterning process.

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2 CONTRIBUTIONS The idea of off-plane extrusion and prototype designs of tube extrusion devices was first

introduced by Arianna McAlister. Arianna designed and fabricated microfluidic devices and also

conducted the initial characterization of devices and formed alginate tubes during her master’s

thesis [13].

The presented work on the one-step formation of homogeneous tubular structure was carried out

in collaboration with Mark Jeronimo, which includes modifications to microfluidic device

designs and fabrication protocol, an analytical model describing tube formation, a thorough

analysis of the operating conditions associated with the successful formation of homogeneous

alginate tubes, and the investigation of different not previously demonstrated morphologies of

homogeneous tubes, in particular, section 6.1, 6.2 and 6.3. Mark Jeronimo’s contribution is also

reflected in Figure 1, Figure 6, Figure 7 and Figure 8.

Lian Leng contributed to the designs of extrusion devices for stimulus responsive sheets shown

in Figure 13 and Figure A4.

For myself, I did the studies on heterogeneous alginate tube formation, collagen tube formation,

stimulus responsive particle synthesis and characterization of stimulus responsive hydrogels, in

particular, section 6.4, 6.5 and 6.6. I lead the research on special morphologies of homogeneous

tube in section 6.3. I also contributed to the collaborative work of microfluidic device designs

and fabrication, analytical model construction and validation, exploring operating parameter

space of tube extrusion system, in particular, section 6.1 and 6.2.

The tube formation and patterning technology demonstrated in this thesis is protected under US

Patent PCT/CA2014/050413.

3 MOTIVATION The routine formation of perfusable soft material tubes and tubular assemblies is a key

requirement for recapitulating organ structure and function in a wide range of tissue engineering

and regenerative medicine applications. Various strategies for the formation of perfusible cell-

laden tubular constructs have been demonstrated. However, the impact of these strategies is still

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limited by a lack of understanding of how heterogeneities provided by cells and their niches

impact remodeling and long-term properties of these engineered tissues. The availability of a

routine technology that allows the formation of heterotypic soft material tubes is required for

such systematic studies. Current approaches for the formation of perfusable tubular soft materials

are generally highly manual in nature, discontinuous, lack scalability, are limited in the range of

tube dimensions possible, cross-linking mechanisms, and mechanical properties produced. In

addition, the spatial heterogeneities commonly associated with the structure and function of

intact tissues and organs cannot be recapitulated achieved in present approaches.

Incorporating stimulus responsive particles into tubular hydrogels will provide additional

flexibility in the control of tube mechanical properties. The present patterning techniques of

stimulus responsive soft materials involve multiple steps which make the process lengthy and

costly. There is no technology available to introduce stimulus responsiveness into soft material

tubular structures in one-step extrusion tube formation process. Stimulus responsive

homogeneous and heterogeneous tubes should lead to new tubular structure assembly techniques

and smart tubes which predictively regulate flow rate and direction.

4 SPECIFIC AIMS

4.1 One-step formation of homogeneous soft material tubes

• Optimize the existing designs of off-plane extrusion devices to improve the yield of

device fabrication.

• Fabricate and characterize microfluidic devices for the formation of homogeneous

hydrogel tubes.

• Demonstrate successful one-step off-plane extrusion of homogeneous alginate tubes and

determine the operating envelope for successful sodium alginate tube formation.

• Develop an analytical model to predict the dimensions of extruded tubes as a function of

flow rates of crosslinker solutions and biopolymers solutions.

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• Measure the wall thickness, inner and outer diameters of formed alginate tube at different

alginate, and inner and outer streaming flow conditions.

• Compare the measured tube dimensions with expected dimensions calculated from the

analytical model.

• Demonstrate the extrusion of collagen I tube using the same system.

• Measure the dimensions of collagen tubes and compare with alginate tubes extruded with

the same conditions.

4.2 One-step formation of heterogeneous soft material tubes and stimulus responsive tubes

• Design, fabricate and test microfluidic devices that promote the formation of

heterogeneous tubes, specifically single layer “Janus” tubes, tubes with axially aligned

stripe patterns, and multilayer tubes with up to three uniform thickness layers.

• Measure the uniformity of the two halves of Janus tubes.

• Measure the thickness and uniformity of the layers in the multilayer tubes.

• Explore the relationship between the stripe width and the driving pressure of secondary

biopolymer solution.

• Measure the dimensions of Janus/multilayer/striped tubes and compare them with

homogeneous tubes extruded in the same conditions.

4.3 Stimulus responsive hydrogel characterization and extrusion

• Synthesize and characterize stimulus responsive nanoparticles which are responsive to

temperature, pH or both.

• Dope nanoparticles into alginate solution to extrude stimulus responsive hydrogel sheets.

• Characterize the formed stimulus responsive hydrogel including its responsive time,

reversibility, volume change and critical temperature/pH.

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• Demonstrate extrusion of stimulus responsive hydrogel tube using the same system as

before.

• Show stimulus responsiveness of extruded hydrogel tubes.

• Extrude heterogeneous (Janus) stimulus responsive hydrogel tube and demonstrate

stimulus induced shape transformation (bending).

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5 MATERIALS AND EXPERIMENTAL

5.1 Microfluidic device designs The feature layers of the multilayered microfluidic devices are arranged such that one or more

biopolymer matrix layers (blue color) are surrounded by layers delivering the crosslinking

solution at the top and bottom. Each layer contains the same key features: the solution, for a

given layer, was distributed evenly from a single inlet such that all channels have the same

resistance. The channels were arranged in a radial configuration around the common outlet hole.

The smallest feature size of 300µm was the width of the microchannels that radially delivered

the fluid streams to the outlet hole located at the center. All the channels in all the fluid layers

shared the same depth of 150µm. All fluid paths in a given layer had the same length and

corresponding widths and therefore identical flow resistance, causing uniform flow distribution

around the circumference of the center outlet. Within each layer there is a 1.2mm diameter pillar

feature was located in the center hole. The post was sacrificed when defining the common outlet

hole and ensured the center outlet did not collapse during device fabrication. The device design

was compliant with outlet diameters between 1.2mm (equivalent to the pillar diameter) and

4.0 mm (equivalent to the surrounded by the radially distributed channels). During device

fabrication, the outlet size should not approach the limits as the process is manually done which

carries human error. The size of the common outlet hole was 3.2mm, unless specified otherwise.

As the diameter of the outlet configuration was larger than the outlet hole, the space between the

outlet and channels allowed the liquids to distribute evenly around the outlet, even if the center

holes were not perfectly aligned. These selected device design and fabrication protocol renders

devices robust, reusable and insensitive with respect to small misalignment between layers. The

maximum tolerance of the misalignment is 800µm (while two layers are misaligned, the

maximum distance between the central points of the two layers which allows punching a 3.2mm

hole without cutting channels). In the older versions of designs made by Arianna McAllister

[13], the maximum outlet size is same as the punched hole size. Therefore, any human error such

as misalignment of layers or tilted punching will result in cutting channels, which leads to

imbalanced flow resistance around the outlet. The biopolymer containing layer only differed

from the layers carrying the inner and outer fluid (crosslinker) stream in the channel length by

2.6mm less. The footprints of different layers ensured that adjacent layers have as little channel

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overlap as possible. Overlap area of adjacent layer (matrix layer and streaming layers) is

23.3mm2.

Figure 1(a-d) show various example designs of microfluidic devices configured for preparing

different tubular structures in terms of geometry, composition and throughput. Inlets are labeled

as follows: streaming fluid inlet (1), matrix solution 1 inlet (2), matrix solution 2 inlet (3). In

Figure 1 b, all the 3 inlet of matrix solution 2 are labeled the same. However, they can be

different matrix solutions. The same conditions apply to matrix solution 1 in Figure 1d. The

solutions for the 3 matrix inlets can be different. The device in Figure 1a may be employed, for

example, for the formation of the tubular structures shown in Figure 7(c, d, e), Figure 8(a, e) and

Figure 10 a. The device in Figure 1b may be employed, for example, for the formation of the

tubular structures with spots and stripes around the circumference, shown in Figure 9(d, e) A2(a-

d). The device in Figure 1c may be employed, for example, for the formation of the tubular

structures shown in which the two halves of the tube wall made of different materials as shown

in Figure 9g. The device shown in Figure 1d was used to form three-layered tubes, for example,

for the formation of the tubular structures shown in Figure 9 (a, b).

Figure 1a shows the design of the basic device design which is used for homogeneous tube

extrusion. This device consists of 3 layers, from top down, inner streaming (Figure 1 a 2nd from

left), matrix (Figure 1 a 3rd from left) and outer streaming (Figure 1 a 4th from left). The inner

streaming and outer streaming layer mentioned below are referring to these layers. All the

devices share the same designs of inner and outer streaming layers.

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Figure 1. Microfluidic device designs. (a) Device design for homogeneous tube formation consisting of (top down): inner streaming layer (green color), matrix layer (blue color) and outer streaming layer (red color). The inner and outer streaming layers mentioned in the later captions are referring to the layers shown here. The inlets of the channels are labeled as (1) streaming

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fluid, (2) matrix solution 1 inlet, and (3) secondary matrix solution. The channels radially feed the solutions from the inlets to the center outlet. (b) Design for formation of heterogeneous tubes with at 3 equidistant locations along the circumference consisting of (top down): matrix solution feeding layer (black color), inner streaming layer, secondary biopolymer layer with 3 locations for insertion along the circumference and outer streaming layer. Secondary biopolymer solutions are fed into the biopolymer layer through the feeding layer at the locations for insertion. (c) Device design for formation of “Janus” tubes, consisting of 3 layers (top down): inner streaming layer, half-half biopolymer layer and outer streaming layer. There are two inlets feed the biopolymer layer from the two sides. (d) Device design for forming three-layered tubular structure consisting of (top down): inner streaming layer, matrix layer 1 in Figure 1a, matrix layer 2, matrix layer 3 and outer streaming layer. The center layer is matrix layer 2. The right layer is matrix layer 3.

Figure 1(b-d) serve the formation of heterogeneous tubes and are modifications to the device

design for the formation of single layer homogeneous tubes (Figure 1a). To avoid repetition,

Figure 1(b-d) only shows the additional layers that are added to the layers already shown in

Figure 1a. In each panel, the first drawing from left side corresponds to the top view of all the

layers stacked. Overlapping channels were not shown in the panels.

A single layer soft material tube accommodating three stripes of a secondary biopolymer was

produced using the configuration shown in Figure 1b. The device shown in Figure 1b is suitable

for forming tubular structures with 3 axially aligned stripes with equidistant spacing along the

tube circumference. The devices consisted of 4 feature layers from top down, formed layers

including a matrix solution feeding layer (Figure 1b 3rd from left), an inner streaming layer, an

additional matrix layer for introducing the secondary biopolymer solution at 3 locations (Figure

1b 2nd from left) and an outer streaming layer. Stripe-pattered tubes were formed by introducing

a secondary biopolymer solution from pressurized wells while supplying the primary biopolymer

at a fixed flow rate.

Figure 1c shows the device design that allows the formation of “Janus” tubes, i.e., a single-layer

tubular structure with a varying composition along its circumference. For example, one half of its

wall is composed of soft material A and the other half is composed of soft material B, with two

corresponding biomaterial inlets for each side. The corresponding microfluidic device may be

employed for forming a heterogeneous tubular structure with different compositions on each half

of the tubular structure and consists of 3 layers: a layer for the distribution of the inner fluid at

the top, the center layer distributing the two matrix streams A and B (Figure 1c 2nd from left),

and an outer fluid layer at the bottom.

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Multi-layered soft material tubes were obtained using the microfluidic device design shown in

Figure 1d. Since the microfluidic device design is modular in the vertical direction, it is possible

to incorporate multiple matrix layers and thereby obtain multi-walled tubular structures by

stacking more matrix layers in between the inner (top layer) and outer streaming layers (bottom

layer). Figure 1d shows the design of a microfluidic device for the formation of tubes with three

co-axial matrix layers. The device consists of the following layers (top down): an inner

streaming layer, a layer for biopolymer solution 1 (Figure 1a 3rd from left), a layer for

biopolymer solution 2 (Figure 1d 2nd from left), a layer for biopolymer solution 3 (Figure 1d 3rd

from left) and an outer streaming layer. The approach could be further extended towards larger

numbers of matrix layers.

5.2 PDMS device fabrication Devices were prepared using multilayer soft lithography. Transparency mask designs shown in

Figure 1 were prepared using a computer aided design program (AutoCAD version 2013,

Autodesk, San Rafael, CA, USA) and photomasks were printed at a resolution of 20,000 DPI

(CAD/ART Services, OR, USA). Masters for replica molding were prepared by rinsing

7.62mm×10.16mm glass substrates (Corning Inc., Corning, NY, USA) with isoproponol,

acetone, and then isoproponol and subsequent dehydration on a hot plate (model HP30A, Torrey

Pines Scientific, San Marcos, CA, USA) at 200ºC for 30 min. Slides were allowed to cool to

65ºC and then oxygen plasma treated for 30 s (model PDC-32G, Harrick Plasma, Ithaca, NY,

USA). A seed layer of SU-8 25 negative photoresist (Microchem, Newton, MA, USA) was spun

at 2000 rpm for 30 s (model SCS G3 spin coater, Specialty Coating Systems, Indianapolis, IN,

USA) and soft baked at 65ºC and 95ºC for a total of 10 min. The seed layer was exposed to UV

light (365 nm) for 13 s (mask aligner model 200, OAI, San Jose, CA, USA), and baked for 6 min

at 95 ºC. Feature heights of 150μm were defined by spinning two 75μm thick layers of a negative

photoresist (SU8 2050, Microchem, Newton, MA, USA) at 1900 RPM for 30 s, and soft baking

in between spins for 5 min at 65 ºC and 15 min at 95 ºC. The substrate was baked for 15 min at

65 ºC, and for 45 min at 95 ºC. The substrate was exposed through the transparency mask with

365 nm UV light at 250 J, hard baked for 20 min at 95 ºC, and developed for 10 min (SU-8

Developer, Microchem, Newton, MA, USA). The masters were rinsed with isoproponol, dried

under N2, and baked for 15 min at 80 ºC. Separate masters with the same feature layer depth

were prepared using the outlined fabrication sequence. The masters were degassed in -25 mmHg

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at room temperature for 1 hour.

To obtain reliably bonded multilayer devices, a multilayer partial curing and bonding technique

was adopted from previously established protocols [65] in combination with fully cured PDMS

layers. Pre-polymer and curing agent were mixed in a ratio of 10:1, spun onto masters at 400 rpm

for 30 s, resulting in a 200 μm thick PDMS layer. All layers except for the top layer were

prepared as described. During the preparation of the top layer, an approximately 5mm thick layer

of uncured PDMS was poured onto the master, and the layer was fully baked at 80ºC for 20 min.

The second layer was baked at 80ºC for approximately 9 min. When partially cured, the thicker

top layer was aligned over the sticky second layer and air bubbles were carefully removed. The

combined layers were baked for 18 min to ensure a strong bond. The sequence was successively

repeated until all layers were attached to each other. After bonding, inlet holes were punched for

all layers and both top. The bottom was sealed with partially cured PDMS sheets with no

features. The center outlet hole was then punched and sealed the top end.

5.3 Homogeneous tube extrusion Alginate sodium salts (Sigma Aldrich), 0.1 µm carboxylate-modified microsphere (FluoSphere,

Life Technology, ON, Canada), glycerol (Bioshop Canada, ON, Canada) and calcium chloride

(Bioshop Canada, ON, Canada) were used as received. The matrix solution was prepared by

dissolving 1 g of alginate in 30 mL of distilled water. 20 mL of glycerol was then added to the

mixture. The mixture was stirred for 10 min and then sonicated for 60 min. The streaming

solution was prepared by mixing 600mL of distilled water and 400mL of glycerol. 11.1g of

calcium chloride was then added. The concentration of Ca2+ in the streaming fluid is 100mmol. 1

L of reservoir solution was mixed with 254 mL of glycerol and 746 mL of water. The streaming

solution was prepared by mixing 20 mL of glycerol with 30 mL of water. The resulting density

of both the matrix solution and the streaming fluid was 1.12g/mL. The density of reservoir

solution was 1.08g/mL. The streaming fluid was fed to the microfluidic device by a syringe

pump (model Nemesys, Centoni, Korbussen, Germany). The inner and outer streaming fluids

were controlled independently. The matrix solution was fed by a second syringe pump (model

PHD 22/2000, Harvard Apparatus, Holliston, MA, USA). The reservoir was made of

polycarbonate sheets (thickness 1mm) and was 12cm tall, 20cm long and 15cm wide. The device

was placed horizontally within the reservoir at a distance 10 cm from the bottom. The

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confinement was 2.8cm long and was made of Tygon tubing. Figure 2 illustrates the schematic

of the experimental setup of homogeneous tube extrusion.

Figure 2. Experimental setup of homogeneous tube extrusion. The biopolymer solution (blue), inner (green) and outer (red) streaming fluids are stored in glass syringes and are driven by syringe pumps into the microfluidic device. The microfluidic device guided the solutions so they flow out from the device at the center outlet facing downwards. The solutions and extruded tube are then released into a reservoir filled up by streaming solution.

5.4 Multilayer tube extrusion Streaming and reservoir solutions were made the same way as homogeneous tube extrusion

experiment. 2%wt Alginate solutions were made the same way as before but with different 0.1

µm carboxylate-modified microspheres in solutions for the different tube layers. Alginate

solutions with different fluorescence were driven by syringe pumps with synchronized infusion

rate.

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5.5 Janus tube extrusion Streaming and reservoir solutions were made the same way as homogeneous tube extrusion

experiment. 2%wt Alginate solutions were made the same way as before but with different 0.1

µm carboxylate-modified microspheres in solutions for the two halves of the tube. Alginate

solutions with different fluorescence were driven by syringe pumps with synchronized infusion

rate.

5.6 Preparation of stripe-patterned hydrogel tubes The setup of spotted/striped tube extrusion is similar to homogeneous tubes that the inner/outer

streaming fluids and matrix feeding to the microfluidic device is the same. As the matrix layer

has additional inlets for spotting/striping materials, on-chip wells are molded on the top of the

device to contain the spotting/striping fluids. The fluids in wells can be different. A schematic

drawing of this setup is shown in Figure 3. The matrix layer of the device shown in Figure 1b

was modified such that 3 channels for the insertion of a secondary biopolymer solution were

inserted between the radially oriented channels delivering the primary biopolymer solution

equidistantly around the circumference of the center outlet. A second matrix layer served to feed

the secondary biopolymer solution and was added to the top of the device. The solutions up to

1.5mL are stored within wells that were directly attached to the top layer of the device. When

pressurized, the fluids in the well will enter the device and form spots/stripes on the extruded

tubes. The pressures in the wells are controlled independently by a magnetic valves array. The

magnetic valves array in connected to a gas pressure regulator which regulates the pressure goes

in to the valves. Head pressures varied between atmospheric pressure and a pressure level above

atmospheric pressure. When the valve is open, the pressure in the corresponding well is elevated

to the level of the gas pressure regulator output. When the valve is closed, the pressure in the

well drops to atmosphere level. The upper pressure level was adjusted with a 0-15psi digital

pressure regulator (model 3410, Marsh Bellofram, Newell, WV, USA) that was connected to

compressed house air at its inlet to individual solenoid valves hosted in a manifold (The Lee

Company, Westbrook, CT, USA) at its outlet. The solenoid valves were electronically addressed

by a and a control circuit shown in Figure 4 and a custom software program shown in Figure 5

(Labview, National Instruments, Austin, TX, USA). By activating a solenoid valve, the pressure

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in the particular well increased from atmospheric pressure to a pressure level above atmospheric

pressure the secondary biopolymer to enter from the well into the microfluidic device.

Figure 3. Experimental setup of spotted/striped tube extrusion. The primary biopolymer solution (blue), inner (green) and outer (red) streaming solutions are driven by syringe pumps into the microfluidic device. The secondary spotting/striping biopolymer solution (pink) is stored in on-chip wells. It is driven by a head pressure Ph. The pressure can be switched on/off by an array of magnetic valves where each valve controls one well.

5.7 Collagen tube extrusion Collagen I (Rat Tail, BD Science, Mississauga, ON, Canada), sodium chloride (Bioshop Canada,

ON, Canada) and PEG (Mn 35,000, Sigma Aldrich) were used as received. 100g of PEG powder

was added to 900mL of deionized (DI) water. 20g of sodium chloride was then added to the

solution. The end product is 1L of 10%wt PEG and 2%wt NaCl solution. This solution was then

used to fill up the reservoir and was used as streaming solutions. Collagen solution was used as

the biopolymer matrix solution and was infused using a syringe pump. The inner and outer

streaming fluids were controlled independently by another syringe pump. The Tygon tubing

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confinement was 18.0cm long. Collagen tubes were extruded into the reservoir and were allowed

to stay in the reservoir solution for 15min. They were then transferred to petri dish and left in

room temperature for a day for complete gelation.

5.8 Valve control board design The schematic diagram of the circuit design is shown in Figure 4a. The board layout is shown in

Figure 4b. The layout was transfer printed on to a PCB board. There are several jumpers needed

on the top of the board which is shown in Figure 4c. Two SN754410 H-bridge (Texas

Instruments) and one 74HCT04 hex inverter (NXP Semiconductors) were used to construct the

circuit. The board is powered by an external power supply of +9V. The power supply can be

switched on/off by a power switch signal which give +5V or 0V. The circuit is designed to

control up to 4 valves. There are 4 signal pins on the board that connect to a digital signal source

(0V and +5V). The digital source USB6008 (National Instrument, TX, USA) is controlled by

Labview. Each digital ping controls one valve. The circuit convert digital signal to voltage

supply to the magnetic valves. The valves are opened by +9V and are closed by -9V. The boards

reserved pins sharing power supply, power switch signal and ground between boards, so they can

be stacked to control more valves (e.g. stacking 2 boards will control 8 valves).

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Figure 4. Valve control circuit. (a) Schematic drawing of valves control circuit board. This board uses motor controllers to supply and invert the voltage to magnetic valves. SWITCH: connection to external analog input for H bridge power control. VALVEIN: voltage supply to valves. VALVEOUT: inverted voltage supply to valves. HBRIDGE: H bridge unit. INVERTER: hex inverter. SIG: connection to external digital signal which controls the valves. GND: ground. 5V: connection to external 5V power source. 12V: connection to external 12V power source. (b) PCB board blueprint of the valves control circuit board. This layout is transfer printed on the back of the circuit board. The printed wires connect the pins of the electronic components on the board. Red: wire transfer printed wire. Green: through holes on board. (c) Top view of valve control board. The wires are the jumpers used to complement connects of the printed circuit on the back of the board. Inlet for power supply and electronic components are mounted on the top of the board.

5.9 Labview control programs A Labview program was written to control the magnetic valves. Signals are sent to control board

in batches. Each batch of signals defines a status of all the valves. All the valves are switched

from one status to another when the next batch of signals is sent. The number of status and

number of valves been controlled are determined in the Step 1 of the program. The statuses of

the valves can be defined on the interface or import from a record file which is in the Step 2 of

the program. The status can be manually input in Step 3. The time frame of each status can be

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defined in this step too. Upon complete the 3 steps, the program can be launched to execute

predetermined patterns. During the execution, the signal pattern, the duration of each status and

number of repetitions of the pattern can be modified. The opening/closing durations of the valves

determines the amount of fluids goes into the spotting/striping channels and the distance between

adjacent spots/stripes. The pattern of the valve activities can be defined so the valves repeat the

pattern and produce repeating pattern on the tube. The program is designed in the way that the

variables are defined by the user in an order. Variables are disabled until the user defines the pre-

requisite variable for them. To prevent overheating of the magnetic valves, power is only

supplied to a valve when the status of the valve needs to be changed. The program can

automatically detect a change of status. At all time, the program can be stopped by pressing the

‘stop all’ button on top. All the valves will be closed and the circuit board will be disabled. The

interface of the program is shown in Figure 5a.

Another Labview program is made to control the gas pressure output of the gas pressure

regulator. The program converts gas pressure into analog signal which is sent to the pressure

regulator via control unit USB6008. The target pressure can be input in the program and can be

changed in situ. At the same time, the pressure in the regulator can be monitored. The interface

of this program is shown in Figure 5b.

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Figure 5. Labview control programs. (a) Labview interface for magnetic valves control. There are 3 preparation steps which defines the pattern of the output digital array. The program can be stopped at any time by pressing the ‘stop all’ button during operation. (b) Labview interface for gas pressure control. The output pressure of pressure regulator can be defined by set a target in the program. The chart tracks the output pressure at real time.

5.10 Tube dimension measurements 3 segments of tube of 1mm length were cut from random positions on an extruded tube. The 3

segments were transferred to Nikon A1 confocal microscope (Nikon, Japan) for imaging. All the

confocal images of tube cross sections were processed using ImageJ software. For each image,

the cross sectional area enclosed by the inner tube wall was selected. The perimeter of this area

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was measured. The inner diameter of a tube DI was calculated based on the perimeter value. The

outer diameter DO of a tube was calculated based on the perimeter of the area enclosed by the

outer tube wall. The wall thickness was calculated as (DI – DO)/2. The experiment was repeated

using tubes extruded from 3 different devices for each extrusion condition (inner/outer streaming

flow rates and biopolymer matrix solution flow rate). Total of 9 measurements were obtained for

each extrusion condition. The plotted results were the average value of the 9 measurements and

the error bars represents the standard deviation. There was error associated with manually cutting

of the tube cross sections. Cuttings may not be perfectly perpendicular to tube axial direction

therefore the tube will not stand perfectly vertical while confocal image is taking. This error is

reflected in error bars of plots and cannot be isolated or eliminated.

5.11 Polymer synthesis and hydrogel solution preparation Poly(N-isopropylacrylamide-co-vinylacetic acid) (PNIPAM-VAA) particles were synthesized by

precipitation polymerization. Ammonium persulfate (APS), N,N-methylene bisacrylamide

(MBA), sodium dodecyl sulfate (SDS) and vinyl acetic acid (VAA) were used as received

(Sigma Aldrich). NIPAM (Sigma Aldrich) was recrystalized to remove inhibitors. According to

the recipe in Table 1, NIPAM, MBA, SDS and VAA were dissolved in 150 mL of water in a

three-necked flask which was fitted with a condenser and a magnetic stirrer. The solution was

heated to 70 ºC under a nitrogen atmosphere. APS was dissolved in 10 mL water and injected to

the solution after 30 min. Polymerization was carried out for 6 hours while stirring at 250 rpm.

The solution was cooled and collected for ultracentrifugation during 45 min at 120000g. The

condensed particle solution was mixed with 1.5% wt alginate salt and sonicated (Branson 5510,

Branson, Danbury, CT, USA) for 60 min. Bubbles produced during mixing were removed by

centrifuging the solution at 3500 rpm for 15 min. The solution was further condensed by heating

to 50 ºC for 2 hours and removing water of 50% of the total volume. The resulting solution was

stirred for 10 min and sonicated for 1 hour.

5.12 Stimulus responsive material characterization The polymer solutions used for particle size characterization were collected after polymerization

and before centrifugation. The solutions were diluted for 16x before doing measurements. In the

case where Ca2+ ions were present, calcium chloride was added to the diluted solution so the

Ca2+ concentration was 100mmol. The pH of the diluted solution was controlled by titrating

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hydrochloric acid (Bioshop Canada, ON, Canada) and soldium hydroxide (Bioshop Canada, ON,

Canada). Polydispersity indexes (PDI) and hydrodynamic diameters of particles were obtained

by dynamic light scattering using Zetasizer Nano (Malvern Instruments). Stimulus responsive

sheets were used for some of the characterizations of the material. They were extruded using in-

plane extrusion method via a PDMS device. The flow rates of hydrogel solution were controlled

by a syringe pump. The extruded sheet was collected using a rotating drum. The linear speed of

the drum is controlled by computer. The shape changes of the sheets/tubes were recorded as

videos by a camera above the container of the sheets. Images were cropped from the videos for

measurements of length and curvature change of the sheets/tubes using ImageJ.

5.13 SEM imaging of tubular soft material constructs Scanning electron microscope (SEM) samples were prepared by fixing the hydrogel samples in

2% glutaraldehyde in a 0.05M sodium cacodylate buffer at pH 7.4 for 1hr at room temperature,

followed by the incremental substitution of the liquid phase with 100% ethanol. Dehydration of

the samples was achieved with liquid carbon dioxide (CO2) at 10°C in a critical point dryer.

Samples were then sputtered with gold. Samples were subsequently heated to 31°C while

increasing the pressure to 7.2MPa, transitioning the CO2 to supercritical fluid conditions. An

environmental SEM (20000X, FEI XL30 ESEM, FEI, Hillsboro, OR, USA) was used to take the

images at low vacuum mode in order to avoid charge accumulation on the sample surface.

5.14 Environmental temperature and pH control of stimulus responsive hydrogels

Temperature and pH control of the hydrogel tubes and sheets were achieved by switching water

baths, and complete stimuli-responsive shape changes were observed after maintaining a

constant environment for 15 minutes. In the case where the responsive times of the hydrogel

tubes/sheets were measured, water with different temperature and pH were flowed on the

stimulus responsive sheets. The sheet was fixed at one end in a plastic half tube. The direction of

the sheet is the same as the half tube. Water with different temperature and pH were flowed

along the direction of the sheet from the fixed end.

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6 RESULTS AND DISCUSSION 6.1 One-step formation of homogeneous soft material tube

The presented method for extruding single-layer homogeneous tubes uses a three-layer

microfluidic device, as shown in Figure 6. Figure 6a illustrates the continuous formation of a soft

tubular material in the device-normal direction (off-plane extrusion).

Biopolymer solution, inner streaming fluid and outer streaming fluid enter the microfluidic

device at well-defined flow rates and are distributed to separate device layers. Within each

device layer, the solutions are first circumferentially distributed around the common central

outlet and then radially directed to the outlet through an array of uniformly long and wide

microchannels. The three miscible fluid streams form sheath flows that meet at the 3.2mm

diameter common central outlet where the sheaths are redirected in the device-normal direction

to form concentric fluid layers that enter a confinement tube. The inner fluid stream occupies the

center region the outer fluid stream occupies the region close to the wall of the confinement tube

and the matrix fluid is confined in between. The inner diameter of the confinement tube was

identical to the diameter of the common outlet. The biopolymer occupying the coaxial matrix

layer polymerized upon contact with the inner and outer fluids, producing a continuously flowing

soft material tube.

The confining tube (Tygon, inner diameter 3.175mm, length LC=28mm, unless otherwise

specified) retains the co-axial organization of the three fluid layers while polymerization takes

place and the flowing soft material solidifies. In cases where the gelation is rapid enough, the

confinement is not necessary for successful tube formation. However it significantly extends the

operational envelope available for successful tube formation and also improves the uniformity

and predictability of tube diameters and thicknesses. The high degree of predictability is a

prerequisite for the consistent formation of tubular structures with increased geometrical

complexity. Examples are heterotypic tubes with tailored heterogeneities that predictably affect

the physicochemical properties of the produced tubes. In the case that the density of tube is

higher than the reservoir fluid, the required minimum flow rates and the minimum total flow rate

are calculated that prevent the unwanted effect of a backflow from the reservoir into the outer

fluid layer of the confinement tube. The as-produced tube exits into a reservoir that is filled with

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a cross-linker and has a density that is lower than the three fluids. The diameter of the confining

tube may be altered with the inner diameter of the common outflow hole of the microfluidic

device in order to expand the range of diameters of the formed tubes.

Figure 6b shows an example illustrating the microchannel network of a microfluidic device used

for the formation of a single layer homogeneous soft material tube. As shown in Figure 6c, the

illustrated microfluidic device consists of three feature layers: an inner streaming layer (top,

green color), an intermediate matrix layer (middle, blue color), and an outer streaming layer

(bottom, red color). Devices used for the formation of heterogeneous tubes may include

additional matrix layers. The device was fabricated using the previously discussed fabrication

sequence. The microchannel network on each layer contains the same key features: the solution

is distributed from a single inlet, circumferentially distributed around the common through a

hierarchical microchannel network with uniform flow resistance. The device architecture makes

it easy to stack additional layers during device fabrication, e.g., for creating heterogeneous

tubular constructs. For example, stacking n matrix layers in between the crosslinking layers

permits the formation of n-layered tubes with a different compatible material for each layer.

Small changes to the matrix layer design allow to introduce heterogeneities within a single layer,

e.g. to define axially aligned stripes. In addition to altering the inner, matrix and outer flow rates,

the range of addressable tube diameters can be further expanded by modifying the size of the

common exit hole and inner diameter of the confinement tube.

Figure 6d shows a photograph of the top view of a fabricated device where the fluids in the

respective channels are labeled with food color. Figure 6e shows a bright field image of the

proximity of the common outlet hole. Figure 6f shows a side view of a microfluidic device with

attached confinement while a soft material tube exits. Figure 6g shows the experimental setup.

Fluid streams may be supplied using three syringe pumps at the well-defined flow rates QI, QM,

and QO. Alternatively, the matrix solution may be supplied from on-chip wells, by controlling

the well head pressure. The produced soft material tube is collected within a liquid-filled

reservoir. The image in Figure 6d is taken from camera 1 and the image in Figure 6e is taken

from camera 2.

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Figure 6. Device design and experimental setup. (a) Schematic illustration of soft material tube formation using a multilayer microfluidic device at the extrusion hole during biopolymer through a confinement tube of length Lc. The biopolymer solution (blue), inner (green) and outer (red) streaming solution meet at the extrusion outlet at the center of the device, and change direction to flow perpendicularly downwards. (b) Top-view of microchannel network of three-layer microfluidic device consisting of inner fluid (top, green color), biopolymer (center, blue

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color) and outer fluid (bottom, red color). Scale bar: 10 mm. Inset: close-up view of region in proximity of common outlet. The bottom layer (red) overlaps with top layer (green) and cannot be seen from top. Scale bar: 2 mm (c) Exploded view of three feature layers and attached confinement. The PDMS layers are bonded to form the extrusion microfluidic device. Scale bar: 15 mm. (d) Photograph of fabricated device with dye-filled channels. Scale bar: 10 mm. (e) Bright-field photograph of resistance channels at extrusion outlet taken with top camera. The resistance channels evenly distribute the solutions around the outlet. Scale bar 1 mm. (f) Photograph taken with side view camera demonstrating continuous tube extrusion through 3.175 mm diameter confining tube, Lc=2.8 cm. The extruded tube is moving downwards into the reservoir. Scale bar: 2 mm. (g) Experimental set-up including fluidic control, the microfluidic device, confinement at extrusion outlet, a liquid filled reservoir and cameras for optical access during extrusion. Camera 1 is responsive for the image in Figure 6e. Camera 2 took the photo for Figure 6f. Scale bar 38.1 mm.

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6.2 Predictive control of tube dimensions In my thesis project, alginate was selected to demonstrate many of the utilities of the tube

extrusion system developed. It is because ionic crosslinking of alginate is a quick process which

fits well in extrusion systems. Many groups discussed above have demonstrated high throughput

extrusion of alginate tubes.

The microfluidic device used by Takeuchi’s group, Kawakami’s group and Ozbolat’s group was

designed to perform in plane extrusion such that the fluid flow inside the device was in plane

with the formed microfiber/tube [8, 10, 11]. In general in-plane extrusion of biopolymer tubes

does not provide a predictive control over the structure dimensions produced. It limits the

scalability of the biomaterial tube extrusion systems. Previous models developed for tube

extrusion in coaxial flows did not account for the fact that the tube wall solidifies during the

process [9, 13]. The increase in viscosity during the extrusion process was not described in the

models. As the inner and outer walls of the tubular structure solidify first, they tend to act as

moving walls instead of fluids. Another model is developed in this thesis which account for tube

wall solidification during the extrusion process.

The flow rates QI, QS, and QO were varied in the experiments to produce alginate tubes with

different diameters in this section. Gelation of the biopolymer solution was initiated upon contact

with the streaming fluids, i.e., at both the inner and outer tube surfaces and then further

progresses to the center of the tube wall. We divide the formation of the soft material tube into

two stages. In the initial first stage, the biopolymer solution remained in a liquid state and the

relative volumes occupied by the three fluid streams that passed through the confinement were

governed by their relative flow rates. In the second stage, after gelation was initiated at the inner

and outer tube surface, we assume a moving tube with template fluid (i.e. inner streaming

solution) at the inside and streaming at the outside.

Figure 7a shows the radial variation of the velocity profile for the 3 fluids passing through the

confinement. In the flow profile, the wall of tube is presented as blue rectangle. The center of

tube is represented as 0 on x-axis. The inner wall of confinement is located at 1 on x-axis. The

top profile shows the status where crosslinking just start when solutions meet. The light blue

region represents biopolymer solution in liquid state. The dark blue region represents the gelled

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biopolymer. The bottom profile shows the status where the tube wall is completely crosslinked.

The crosslinker solutions are assumed to be Newtonian fluids, resulting in a parabolic velocity

profile for the outer fluid layer between the hydrogel tube wall and the inner wall surface. In the

model, the moving tube can be treated as a sheath of fluid with infinitely thin thickness since

there is no relative velocity between the two sides of the tube walls. The velocity and

acceleration of the fluids at the two sides of the wall has to be continuous. As the inner streaming

fluid is completely isolated from the outer streaming fluid, the outer streaming fluid shall feel no

shear from the inner streaming. Therefore, the tube wall can be treated as the central line of fluid

moving in a circular pipe, the fluid here refers to outer streaming which should has a parabolic

flow profile. The tube wall’s velocity equals to the maximum velocity in the outer streaming

fluid. As the gradient of velocity change at maximum velocity is 0. The shear at the outside

surface of the tube wall is 0. As mention before that the inside and outside surface of the tube

wall has continuous velocity and acceleration, the gradient of velocity at the inside surface of the

tube wall is 0, so is the shear. As there is no shear around the inner streaming fluid, the inner

streaming fluid has a flat velocity profile.

The above assumptions can be validated from a view of force balance on the tube wall, while the

wall is not supported, the net force on tube wall has to be 0 (otherwise it will

accelerate/decelerate till its velocity is equal to Vm), where Vm is the radially constant fluid

velocity of the matrix stream. For the net force to be zero, either the two shears on the two side

of wall is opposite (means inner flow is flowing backwards, which is not true because it is driven

by syringe pump with positive infusion rate), or both shears are 0. So dV/dr at wall is 0 (no shear

on both sides of tube wall). V is the velocity of the tube wall.

By substituting the following boundary conditions in to Navier-Stokes equations, for r = 0, V(r)=

Vm, dVi(r)/dr = 0; for r = ri, V(r) = Vm, dV(r)/dr = 0; for r = ro, V(r) = Vm, dV(r)/dr = 0; for r = rc,

V(r)= 0. r is the radial coordination along the tube radius. V is the velocity of inner streaming

fluid at r. ri is the radius of the inner tube wall. rC is the inner radius of the confinement tube , rO

is the other radius of the formed tube.

The flow rate of the outer fluid can be expressed as

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(1)

(2)

For the flow rate of the inner stream follows

. (3)

The matrix flow rate follows as

. (4)

Figure 7b shows the parameter space available for successful tube formation for the case of ionic

cross linking of an alginate solution at a fixed flow rate of the biopolymer solution of

QM=200µL/min, while QI and QO were varied over 3 orders of magnitude. Successful tube

formation was observed over a wide range of parameters, indicating the robustness of the

platform. The outer streaming fluid flow rate has the range of 0 to 3500µL/min. We were able to

successfully form tubular structures in the absence of the outer streaming fluid. The inner

streaming fluid flow rate has the range of 20 to 3500µL/min.

By varying the flow rates, we controllably altered the tube diameter and wall thickness. We

selected a series of flow rate combinations to compare our experimental results with the

predictions from the analytical model, i.e solving for VM, ro and ri by selecting QI, QM and QO

using equation 2-4. The total flow rate of the biopolymer solution and the cross-linking solutions,

QI+QM+QO, was kept constant at 1500 µL/min while the flow rate of the biopolymer solution

remained at QM=200µL/min. We were able to produce a tube with varying diameters and wall

thicknesses along the length when varying QI and QO (Figure 7c-e). We found DI and DO

increase, wall thickness δ decrease with increasing ratio of QO/QI. The outer diameter DO varied

between 0.7 mm and 2.5 mm and δ varied between 70µm and 420µm. As shown in Figure 7f,

model predictions agreed favorably with the experimental data at most conditions except for the

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upper and lower end of considered QO/QI ratios. At high QO/QI ratios, the rapid onset of gelation

that occurred even before entering the confinement tube defined the length of the inner

circumference of the inner tube wall. The further decrease in cross sectional area occupied by

inner fluid within the confinement tube caused buckling of the already cross-linked inner wall.

Consequently, the measured mean tube wall thickness values exceeded the ones predicted by the

model. At low QO/QI ratios, the rapid gelation occurring at the outer surface of the tube wall

limited the outer diameter of the tube. Measured tube diameters were therefore smaller than the

model predictions. In order to include errors generated from device to device variance, 3 devices

were used for each extrusion condition. 3 measurements of tube dimensions at each extrusion

condition for each device were taken to eliminate anomalous measurements.

In order to support the assumption we made in our model that the inner stream acted as a

template and does not apply a stress on the inner tube wall, an experiment was conducted in

which the inner fluid contained no crosslinker. As shown in Figure 7g we found the resulting

tube dimensions to still agree with our model predictions suggesting that the inner fluid acts as a

template occupying the tube center and travels at the velocity of the tube wall. As indicated by

reconstructed confocal scans we found the tube inner wall to irregularly deform and the

operating envelope to decrease in size, if no cross linker was added to the inner fluid. For

example, at QI > 1000 µL/min successful tube formation was only possible if a cross-linker was

present in the inner fluid stream. Only one device was used for this experiment so the error

generated from device to device variance is not presented here. 3 measurements of each

extrusion conditions were taken.

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Figure 7. Dimension control of homogeneous tubes. (a) Left: Cross-sectional view of example device at outlet hole with flow rates QI, QM and QO forming soft material tube of outer diameter DO and inner diameter DI. Tubular structure flows with velocity V in direction of tube axis, x, through confinement conduit of diameter DC and length LC. Polymerization is completed at distance L. Right: Model prediction of velocity profile at device outlet prior to confinement. (b) Operating window for QI (100mmol CaCl2 solution, applies to all inner streaming solutions in this figure) and QO (100mmol CaCl2 solution, applies to all outer streaming solutions in this figure) at QM=200µL/min (2% alginate solution, applies to all biopolymer solutions in this figure) using device in Figure 2a with conditions of successful ( ) and unsuccessful tube

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formation ( ). The red line indicate the boundary between successful and unsuccessful extrusion conditions, n=1. Confocal images of homogeneous alginate tube cross sections extruded with (c) DO=1612µm and wall thickness δ=132µm, for QI=500µL/min, QM=200µL/min and Qo=800µL/min. with (d) DO=1901µm and δ=148µm for QI=650µL/min, QM=200µL/min and Qo=650µL/min. and with (e) DO=2240µm and δ=70µm, for QI=1200µL/min, QM=200µL/min and Qo=100µL/min. (scale bars (c-e) 500 µm) (f) Measured DO ( ) and corresponding model prediction (upper solid line), DI ( ) and corresponding model prediction (middle solid line) and δ ( ) along with corresponding model prediction (lower solid line) at QM = 200µL/min. The predicted results were produced by solving equation 2-4 using Vm, ri and ro as unknown variables. The experimental data points are averaged from 9 measurements. 3 cross sections were randomly chosen from a meter-long tube extruded from a microfluidic device. 3 devices hence 9 cross sections were used for measurements, n=9. (g) Measured tube dimensions as function of flow rates in absence of cross-linker in inner fluid stream at QI = 200µL/min. Measured DO with ( ) and without cross-linker added to the inner fluid ( ), DI with ( ) and without cross linker added to inner fluid ( ), δ with ( ) and without cross linker added to inner fluid ( ) for QI= 200µL/min, n=3 (3 cross sections randomly picked from extruded tubes from the same device). Data points and error bars are mean values and standard deviation of the n measurements for f and g.

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6.3 Soft material tubes of non-circular cross section The morphology of the tube inner wall may deviate from the cylindrical shape with a uniform

curvature radius, DI/2 that we have considered so far. By lowering the ratio QI/Qo to lower than

1/12, we obtained tubular constructs with a buckling inner wall and a uniform diameter

cylindrical outer wall (Figure 8a). The experimental setup remained unchanged and the device

shown in Figure 1a was used. The flow rates were separately controlled using syringe pumps.

Since locations where the inner streaming and the biopolymer matrix solutions were the least

proximal to the confinement tube, cross-linking at the inner tube already started prior to entering

the confinement. Low values of QI therefore lead to a small DI and induced buckling at the inner

wall. Cross-linking at the outer wall surface occurred while the tubular construct had already

entered the confinement. We directly compared the two cases with identical flow rates and no

cross-linked present in the inner stream. In this case, cross-linking at the inner tube wall was not

initiated prior from entering the confinement but only at a further downstream position at which

the crosslinker had diffused through the tube wall from outside. Figure A1a in appendices shows

the resulting tube that shares comparable inner and outer diameter with the tube in Figure 8a but

the inner tube wall is not folded. The inner diameter may be calculated as the hydraulic diameter

(5)

where A is the inner cross sectional area of the tube. Alternatively the calculation may be based

upon the inner perimeter of the tube inner wall

(6)

where S is the perimeter of the tube inner wall. The ratio Di,2/Di,1 is indicative of the degree of

buckling observed at the inner wall surface. Large values of Di,2/Di,1 imply a larger degree of

buckling and values equal to unity correspond to a perfectly circular inner surface. Figure 8b

suggests that with increasing QI the degree of buckling decreases.

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Tubular structures with buckled outer and inner wall surfaces (Figure 8c) can be produce by

using a confinement channel with a diameter DC = 1.59mm which is smaller than the device

central outlet (2.0mm diameter) positioned at the exit of device in Figure 1a. The flow rates were

adjusted to adapt the change of the confinement diameter. The narrower confinement resulted in

forming a tubular construct at reduced diameters and increased velocity. The diameter reduction

induced buckling in both the inner and outer wall surfaces, prior to cross-linking was complete.

The buckled tube shape was therefore retained. We used the ratio Do,1/Do,2 to quantify buckling

behavior in Figure 8d. Do,1 is the hydraulic outer diameter and Do,2 is the outer diameter

calculated based on outer tube wall perimeter. While pressurized, the tube wall will unfold

before it is stretched. The tube can be pressurized and expand without inducing strain in the tube

wall for a diameter change of up to 20%, since Do,2 is 20% larger than Do,1.

A corkscrew shaped inner wall (Figure 8e) was obtained by increasing Qi above 2500µL/min

while reducing Qo below 100µL/min, using the device shown in Figure 1a. The wavelength of

the corkscrew shape inner wall was found to decrease with increasing inner stream flow rate as

shown in Figure 8f. We attribute this shape to a similar reason as in the case of the folded inner

tube wall. As the inner tube wall formed within the device and fixed the perimeter and cross

sectional area of the inner surface before the tube entering the confinement, the only way to

accommodate the large volume of inner streaming entering the luminal side of the tube was to

increase the length of the inner tube wall. In other words, volume of inner streaming fluid equals

to the product of inner cross sectional area and length of inner tube wall. As the cross sectional

area is fixed, an increase in volume resulted in an increase in inner tube wall length. With the

limited length of tube formed, a corkscrew shape inner wall was formed to adapt the increase in

length of the inner tube wall. It also explains the effect that increasing the inner stream flow rate

lowers the wavelength of the corkscrew shape inner wall (increase inner wall length). With

higher volume of the inner streaming enters the tube, the wavelength of the corkscrew shape has

to decrease to fit more waves in a given length of tube.

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Figure 8. Homogenous tube morphologies. (a) Confocal reconstruction of cross section of tube with circumferential buckling of inner wall obtained at: QI=100µL/min (100mmol CaCl2 solution, applies to all inner streaming solutions in this figure), QM=200µL/min (2% sodium alginate, applies to all biopolymer solutions in this figure) and QO=1200µL/min (100mmol CaCl2 solution, applies to all outer streaming solutions in this figure). The tube has dimensions of: DO=1361µm, Di=484µm and δ=438µm. The inner diameter of the tube is calculated based on the cross sectional area enclosed by the inner tube wall. (scale bar 500 µm). (b) Ratio Di,1/Di,2 as

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function of inner streaming flow rate for QM=200µL/min and QTotal=1500µL/min, n=9 (3 measurements each condition for 3 devices). (c) Confocal reconstruction of tube with circumferential buckling inner and outer walls, formed at: QI=23µL/min, QM=10µL/min and QO=277µL/min. The tube has dimensions of: DO=873µm, Di=726µm and δ=74µm. The tube was extruded by a device having outlet diameter of 2.0mm. The confinement has diameter Dc=1.59mm and length Lc=2.8cm. (d) Corresponding ratio Di,1/Di,2 of crimped wall tubes for: QI=23µL/min, QM=10µL/min and QO=277µL/min (scale bar 500 µm), n=3 (from same device). (e) Corkscrew shaped inner wall tube for: QI=3500µL/min, QM=200µL/min, QO=5µL/min. (scale bar 3 mm). (f) Wavelength of corkscrew shape for varying QI, while QM=200µL/min and QO=5µL/min, n=5 (from same device). Data points and error bars are mean values and standard deviation of the n measurements for b, c and f.

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6.4 One-step formation of heterogeneous soft material tubes Incorporating spatial composition heterogeneities of materials on the radial, circumferential and

axial directions in a single system have not been demonstrated by previous in-plane extrusions.

Molding techniques using sacrificial templates also has the limitation producing spatial

composition heterogeneity. In this thesis, the off-plane extrusion was demonstrated extrusion of

biomaterial tubes with spatial composition heterogeneities on radial, circumferential and axial

directions.

In addition to the previously demonstrate continuous formation of single-walled tubular

structures, we now extend the approach towards the formation of heterotypic tubular structures.

We demonstrate the formation of different circular cross section tubes with a heterogeneous

composition as shown in Figure 9.

We first discuss soft material tubes that possess a heterogeneous wall composition in the radial

direction and homogeneous compositions in the circumferential and axial directions. Such

multilayered tubes may be obtained from microfluidic devices that possess two or more feature

layers distributing more than one biopolymer streams. Vertically stacking n biopolymer matrix

layers in between the crosslinking layers permits the formation of n-layered tubular structures

with a different compatible material for each layer. For example, using the microfluidic device

design shown in Figure 1d, a three-layered tube was extruded. Figure 9a shows a confocal

reconstruction of the tube cross-section. The 3 layer tubular structure was prepared using 2%

alginate solutions that were optically distinguished because they carried different payloads of

fluorescent microspheres. The approach may be extended to different biopolymers or payloads.

For instance, a slow gelling biopolymer may occupy the center matrix layer, surrounded on both

sides by layers of rapidly gelling biopolymers. The latter would therefore act as templates and

confine the biopolymer in the center layer while gelation takes place until a monolithic tube is

obtained. As shown in Figure 9 (b, c), the inner diameter DI, outer diameter DO and wall

thickness δ measured for three-layer tubes compare favorably with the single layer tubes

produced at similar (total) matrix and streaming flow rates. Several different combinations of

flow rates were tested and the results are shown in Figure A3 in appendices.

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The cross sectional area of each individual layer of the three-layer tube is uniform as shown in

Figure 9d. The standard deviation of the cross sectional area of the layers is 0.042mm2, which is

9.6% of the total cross sectional area. With the same flow rate of all matrix layers of 100µL/min

each, the thicknesses of the three layers decreased from the inside to the outside, compensating

for the increase in the circumference as shown in Figure 9e.

We now discuss the formation of single-layered tubes where the matrix composition is

heterogeneous in the circumferential direction and homogeneous in the radial and axial

directions. Tubular structures with two or more compatible matrix materials along their

circumference were prepared using device shown in Figure 1b. In addition to delivering primary

biopolymer solution to the outlet, microfluidic channels delivering secondary biopolymer

solutions are inserted surrounding the circumference of the center outlet. One or more matrix

solutions can be delivered independently through these microfluidic channels. The configuration

allows forming vertical tessellations within the tube wall that may predictably alter the

mechanical properties in the circumferential direction along the tube wall. Tube walls composed

of axially aligned stripes of alternating composition may be formed by introducing a secondary

biopolymer solution from a well operated at elevated head pressure Ph while delivering the

primary biopolymer solution using a syringe pump. The details of the experimental setup are

shown in Figure 3. As indicated in Figure 9h, increased (time-constant) head pressures produced

wider stripes in cases where the flow rates of the primary biopolymer solution and streaming

fluids were kept constant. In the examples shown in Figure 9(f, g), the stripes correspond to

alginate doped with payloads of different fluorescent microspheres than the primary biopolymer

(alginate). However, secondary and primary biopolymer solutions are not confined to sodium

alginate. Different tubular structures may be obtained using the same setup and microfluidic

device by varying the extrusion conditions, the base and secondary biopolymer solutions, or the

payload.

The head pressures delivering one or several of the secondary biopolymer solutions may be

altered with time, thereby periodically initiating or preventing the local incorporation of the

secondary biopolymer within the tube wall, and producing heterogeneities not only along the

tube circumference but also along its axis. The volume of the secondary biopolymer solution that

was incorporated per spot as well as the spot size was controlled by the pressure and opening

time of the inlet of the spotting channels. Figure A2(a, b) show that smaller spots were generated

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by reducing upper value of the head pressure. Each spotting channel was independently

controlled, so alternating spotting at different locations can be achieved, as illustrated in Figure

A2c. Multiple spots were produced at the same time, as shown in Figure A2d.

Figure 9i shows the cross-section of a “Janus” tube with a tube wall that has a different

composition at each half and was produced using the microfluidic device shown in Figure 1c.

This is another example of heterogeneous tube with circumferential direction heterogeneity. The

matrix flow rates of the respective halves are Qm,1 and Qm,2. The two halves were formed from

2% alginate solutions that carried different fluorescent microspheres as payloads. Since the two

sides were fed at the same flow rate, Qm,1 = Qm,2, the fraction of the circumference occupied by

the base and secondary material was very similar (Figure 9j). Figure 9(k, l) show that DI, DO, and

δ of the half-half tubes were similar to a single layer tube that was produced from the same

biopolymer solution at comparable flow rates of matrix and streaming flow rates. The application

of this type of tube can be found below while half of the tube is formed by stimulus responsive

hydrogel.

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Figure 9. Heterogeneous tubular structures. (a) Confocal image of cross section of three-layer sodium alginate tube formed at QI=650µL/min (100mmol CaCl2 solution, applies to all inner streaming solutions in this figure), Qm,1=67µL/min (2% alginate solution, applies to all biopolymer matrix solutions in this figure), Qm,2=67µL/min, Qm,3=67µL/min and Qo=650µL/min (100mmol CaCl2 solution, applies to all inner streaming solutions in this figure). (Scale bar 500

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µm). Inset shows three-layer alginate tube formed at QI=650µL/min, Qm,1=67µL/min, Qm,2=67µL/min, Qm,3=67µL/min and Qo=650µL/min. The tube has dimensions of: DO=2079µm and δ=131µm. (Scale bar 200 µm) (b) Comparison of inner and outer diameters of single layer tube (grey color) and three-layer tube (white color), n=3 (from same device). (c) Comparison of wall thicknesses between single layer tube and three-layer tube, showing them having similar dimensions while extruded with same conditions. n=3 (from same device). (d) Comparison of cross sectional area occupied by different layers, showing the layers have similar cross sectional area. The tube is formed at QI=650µL/min, Qm,1=100µL/min, Qm,2=100µL/min, Qm,3=100µL/min and Qo=650µL/min, n=3 (from same device). (e) Comparison of wall thickness of different layers, showing that outer layers are thinner than inner ones. The tube is formed at QI=650µL/min, Qm,1=100µL/min, Qm,2=100µL/min, Qm,3=100µL/min and Qo=650µL/min, n=3 (from same device). (f) Confocal reconstruction of secondary biopolymer stripe incorporated within tube wall. QI=650µL/min, Qm=200µL/min and Qo=650µL/min. Secondary biopolymer incorporated using constant head pressure of 6.89kPa. (Scale bar 500 µm) (g) Cross sectional view of second biopolymer stripes on tubular structure wall. QI=650µL/min, Qm=200µL/min and Qo=650µL/min. Head pressure Ph=13.79kPa. (Scale bar 500 µm) (h) Change of stripe width with respect to change in head pressure Ph, n=3 (from same device). The stripe width increases with increased Ph. (i) “Janus” tube formed at QI=650µL/min, Qm,1=100µL/min, Qm,2=100µL/min, and Qo=650µL/min. The tube has dimensions of: DO=2024µm and δ=118µm. (Scale bar 500 µm) (j) Circumferential length ratio evaluated for “Janus” tubes, showing the two halves have similar arc length, n=3 (from same device). (k) Comparison of inner and outer diameters of single layer tube and “Janus” tube, showing that “Janus” tubes have similar dimensions as single layer tube while extruded with same conditions. Single layer tube: grey. “Janus” tube: white, n=3 (from same device). (l) Wall thickness of single layer tube and “Janus” tube are similar when extruded with the same conditions, n=3 (from same device). Data points and error bars are mean values and standard deviation of the n measurements for b, c, d, e, h, j, k and l.

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6.5 Collagen tube formation Other suitable solidification methods include other forms of polymerization, including physical

and chemical crosslinking, are compatible with the described approach for forming tubular soft

materials. To illustrate the applicability of the approach to temperature induced gelation we now

discuss the gelation of collagen I in a case study.

The disadvantage of alginate based approach is that the shell material (alginate) is not ideal in

terms of cell adhesion. And cell migration in the shell is expected to be poor comparing with

materials like collagen. Therefore, only cell-laden fibers can be made instead of cell-laden tubes

using Takeuchi’s approach [8]. The key factor limited this approach was the slow gelation of

collagen. Alginate was essential for the process that it prevents collagen from leaking out from

the extruded fiber. Moreover, the second step of incubating extruded microfiber limited the

throughput of microfiber production. Below, a process of one-step extrusion of collagen tubes

will be demonstrated which solves the problem of slow gelation of collagen.

Homogeneous tubes were made from collagen solution that was derived from rat tail and had

dimensions that were comparable to the previously described sodium alginate tubes. During the

extrusion, collagen solution was kept at 4°C at pH3and the streaming solutions and reservoir

solution is kept at room temperature. The increase in temperature and pH induced the gelation of

collagen. The inner and outer streaming fluids and the reservoir solution contained 10% wt PEG.

Even though the gelation may be accelerated by exposing the collagen solution to PEG solutions

delivered in the inner and outer fluid streams due to molecular crowding[66], the gelation process

is still slower compared with the alginate case. As a result, a longer residence time is required

while passing through the cylindrical confinement and may be achieved in one of two ways:

either by reducing the streaming and matrix flow rates and thereby the velocity of extrusion, or

by extending the length of the confinement, LC . Both methods were tested and found to be

successful. By slowing the extrusion velocity down to 1mm/s, corresponding to

QI = QO=165µL/min and QM =70µL/min. The residence time required to pass through the Tygon

tube confinement (LC=100mm, DC=3.2mm) was 100s. The PEG solution had a density of

1.011g/mL that was greater than the density of the collagen solution, resulting in the produced

collagen tubes to float and collapse upon exiting the confinement, or to collapse even within the

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confinement, thus impact the yield of collagen I tube production. The extrusion setup is the same

as the alginate extrusion describe in Figure 6.

Collagen gelation rate was found to increase when applying shear to the gel. Therefore, it gels

faster with higher flow rates hence higher extrusion velocity. In section 6.2, an assumption of no

shear at tube wall was made. This assumption does not hold in the case of collagen I extrusion

because collagen solution remain as liquid for most of the time while in the confinement. The

tube needs to gel to the extent that it does not collapse after coming out from the confinement.

While the tube comes out too early in the crosslinking stage, the tube will collapse even when it

is submerged. To achieve this extent of gelation, as mentioned before, 100s is required for 1mm/s

extrusion velocity. On the other hand, only 45s is required for 4mm/s extrusion velocity. For an

extrusion velocity of 4mm/s that corresponded to a QI=QO=500µL/min and QM= 200µL/min. A

confinement with LC=180mm and DC=3.2mm was used, resulting in a residence time of 45s. As

the gelation process was faster at higher extrusion velocity (higher shear) and the tube exiting the

confinement was stiff enough to not collapse easily. Figure 10(a, b) show images of as produced

collagen tubes. As illustrated in Figure 10(c, d), the collagen tubes obtained at these flow rates

had smaller diameters and exhibited higher elastic moduli as compared to the alginate tube

produced with the same flow rates of fluids. The elastic modulus (evaluated in the axial

direction) of the collagen tube was 63kPa as compared with 38kPa for the alginate case (Figure

10e). The elasticity of tube segments was measured using a custom tensile tester (840LE2, Test

Resources, Shakopee, USA). Tube segments of 2cm length were pulled with speed ramp of

0.1mm/s until failure. The load cell used was 1000g.

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Figure 10. Collagen tubes (a) Collagen tubes collected in 90mm diameter petri dish. Conditions for tube formation: QI=500µL/min (10%wt PEG, 2%wt NaCl solution), Qm=200µL/min (3.7mg/mL acid-solubilized collagen I solution at pH3) and Qo=500µL/min (10%wt PEG, 2%wt NaCl solution). (b) Confocal image of collagen tube cross section shown in Figure 10a. The tube has dimensions of: DO=1738µm and δ=141µm. (c) Comparison of inner and outer diameters of single layer alginate tubes and collagen tubes. Alginate tubes were extruded at QI=500µL/min (100mmol CaCl2 solution), Qm=200µL/min (2% alginate solution) and Qo=500µL/min (100mmol CaCl2 solution). Collagen tubes were extruded at QI=500µL/min (10%wt PEG, 2%wt NaCl solution), Qm=200µL/min (3.7mg/mL acid-solubilized collagen I solution at pH3) and Qo=500µL/min (10%wt PEG, 2%wt NaCl solution). Collagen tube: white. Alginate tube: grey, n=3 (from same device). (d) Comparison of wall thicknesses between collagen and alginate tubes. The tubes were extruded at the same conditions as in Figure 10c, n=3 (from same device). (e) The elastic moduli of collagen tube and alginate tubes are in the same order. The tubes were extruded at the same conditions as in Figure 10c, n=5 (from same device). Data points and error bars are mean values and standard deviation of the n measurements for c-e.

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6.6 Stimulus responsive tubular structures There is a great interest in the applying spatially organized soft materials to undergo

programmed shape transformations. I am exploring such shape transformations in the contexts of

homogeneous and heterogeneous tubular/planar soft materials with incorporated stimulus

responsive nanoparticle payloads.

Poly(N-isopropylacrylamide-co-vinylacetic acid) (PNIPAM-VAA) particles with large swelling

ratios in response to stimuli (i.e., pH and temperature changes) are dispersed as a payload in

alginate solution before gelation. PNIPAM and PNIPAM-VAA particles were synthesized by

precipitation polymerization using the recipe in Table 3, which is a widely used synthesis

method for polymers[58]. PNIPAM is only responsive to temperature changes while PNIPAM-

VAA is responsive to both temperature and pH changes. As described in the experimental

section 5.11, the recipe can be used to synthesis 160mL of stimulus responsive nanoparticle

suspension solution. PNIPAM-VAA is used for the experiments in this section unless stated

otherwise. The stimuli-responsive microgel of PNIPAM-VAA has a NIPAM core and a

vinylacetic acid shell as indicated in Figure 11a. NIPAM formed the core because it has greater

reaction ratio than VAA. Therefore NIPAM tends to polymerize before VAA and formed a block

polymer which possess the core shell structure. The core is responsive to temperature changes as

temperature passes its LCST (Figure 11b). The shell is responsive to pH changes due to the

similar charge repulsion between carboxylic groups (Figure 11c)[67]. As the particles are doped

with a high density of 0.089g/mL, the microscopic swelling behavior of the particles will lead to

the macroscopic volume change of the hydrogel sheets. In the figure, base hydrogel refers to

alginate hydrogel. The blue and red arrows only indicate the direction of reaction. They do not

mean temperature change along the direction.

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Table 3 Recipe of stimuli-responsive particles synthesis

Poly (PNIPAM) (mol) Poly (PNIPAM-VAA) (mol)

N-Isopropylacrylamide (NIPAM) 3.5×10-3 3.5×10-3

Vinylacetic Acid (VAA) 0 6.5×10-4

Sodium Dodecyl Sulfate (SDS) 1.7×10-4 1.7×10-4

Ammonium Persulfate (APS) 4.4×10-4 4.4×10-4

N,N-Methylene Bisacrylamide (MBA) 1.24×10-2 1.24×10-2

Figure 11. Stimuli-responsive PNIPAM-VAA nanoparticles. (a) Structure and composition of PNIPAM-VAA particles and its temperature responsiveness. The particle has a VAA shell and a NIPAM core. The core swells as temperature become lower than LCST of the material and the

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reaction is reversible. (b) pH responsiveness can be induced on the shell of the particle. The shell swells as environmental pH become basic and the reaction is reversible. (c) Shape change of hydrogel sheet can be imposed by triggering volumetric change of the doped stimulus responsive particles. A

PNIPAM-VAA particles swell in basic environment and shrink in acidic environment. The

hydrodynamic diameter of PNIPAM-VAA particles are measured using dynamic light scattering

(DLS). Based on DLS measurements, the polydispersity index (PDI) of the particles is 0.012

(Figure 12a). The hydrodynamic diameter distribution of each batch is narrow even the absolute

values are different. A representative distribution is selected for illustration. Particles with PDI

less than 0.1 are considered monodisperse. The hydrodynamic diameter dh at 40 ̊C is half of dn at

room temperature (Figure 12b). The hydrodynamic diameter at pH10 db is 2 times of

hydrodynamic diameter dn at neutral at room temperature (Figure 12c). At different pH, the

LCST of the particles are different. LCST tends to shift higher at higher pH and shift lower at

lower pH (Figure 12d). As for the measurements of the normalized or absolute hydrodynamic

diameters of the particles, since the results are already obtained from average of all the particles

in the solution, there is no error bar for the plot. For each data point, the result of a single

measurement is represented because multiple DLS measurements do not change the result. The

measurements were taken from a single batch of polymer synthesis to remove human errors

(weighing ingredients). Batch to batch difference of the polymer synthesis is not studied here.

In order to crosslink alginate hydrogel, Ca2+ is required. However, it is known that Ca2+ will

cancel pH response by interconnect carboxylic groups on the particle surface [57]. In basic

environment, the particles do not swell significantly in presence of Ca2+ (Figure 12e). The

method overcomes this problem is to use pH as trigger for stimulus response of the temperature

responsive cores of the particles. In order to trigger the stimuli-responsiveness by pH, the VAA

shell is added to PNIPAM so that a change in pH will change LCST of the particles. Therefore,

by moving LCST, the state of the particles can be changed at a fixed temperature by changing

pH (Figure 12f). Though PNIPAM-VAA was synthesized and studied before, this method of

triggering its pH response have never been used. It is observed that PNIPAM-VAA particles tend

to aggregate at temperature higher than its LCST. In presence of Ca2+, the aggregation becomes

more prominent at low pH. Though this aggregation process is reversible by lowering the

temperature below LCST, no meaningful measurements of the hydrodynamic diameter of the

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particles can be obtained through DLS in presence of Ca2+ and at temperature higher than their

LCST.

Figure 12. PNIAPAM-VAA particle size and its stimuli-responsive volumetric changes. (a) DLS polydispersity measurement of PNIPAM-VAA particles. The particles are monodispersed with a PDI of 0.012. (b) Hydrodynamic diameter change of PNIPAM-VAA particles with respect to pH at room temperature. The particle size is largest at pH=10 and is smallest at pH=3. (c) Hydrodynamic diameter change of PNIPAM-VAA particles with respect to temperature at

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pH7. The LCST of the particle is at 32°C at pH 7. (d) Normalized hydrodynamic diameter change of PNIPAM-VAA particles in water for, pH3 ( ), pH7 ( ) and pH10 ( ). The particle is more sensitive (low LCST) to temperature changes at low pH and is less sensitive (high LCST) at high pH. (e) Hydrodynamic diameter change of PNIPAM-VAA particle at room temperature in water ( ) and in calcium ( ) at different pH at room temperature. The hydrodynamic diameter of the particle does not as much when Ca2+

is present. It implies that Ca2+ impedes swelling behavior of PNIPAM-VAA particle during pH change. (f) Normalized hydrodynamic diameter change of PNIPAM-VAA particles in 100mmol Ca2+ solution for pH3 ( ), pH7 ( ) and pH10 ( ). In presence of Ca2+ the particles are more sensitive to temperature at lower pH. n=1 for all measurements.

Temperature and pH can both be used as triggers of stimuli responses of the doped hydrogels.

Homogeneous stimulus responsive hydrogel tubes were successfully extruded using the same

setup for homogeneous tubes discussed before, shown in Figure 14(a, b). The outer diameter of

stimulus responsive tube extruded at QI=650µL/min, QM=200µL/min and Qo=650µL/min is

2640±230µm, while alginate tubes extruded at the same condition has outer diameters of

1920±180µm. The PNIPAM-VAA nanoparticles were concentrated in the stimulus responsive

hydrogel solution such that the nanoparticles were not fully swelled. Therefore, the nanoparticles

absorb water from environment after extrusion and result in the increase of extruded tube size.

Hydrogel sheets were extruded in order to study the stimulus responsive properties of the

material. In-plane extrusion with a PDMS device was used to obtain hydrogel sheet of 200μm

thickness and 5mm width. The outlet of the device in immersed in 100mmol Ca2+ solution made

of calcium chloride and DI water. Stimulus responsive polymer solution was pumped into the

microfluidic device and was distributed as a planar flow. Gelation happened at the outlet of the

device. Formed hydrogel sheet was collected by a rotating drum with linear velocity of 4mm/s.

The length change of stimuli-responsive hydrogel induced by temperature change at low pH can

be as high as 105%. The particle diameter change measurements agree with the hydrogel sheet

length change measurements. 3 sheets extruded at different times were used for the length

change measurements. The stimuli-response of the hydrogel is completely reversible (Figure

14c). After switching water bathes of 25 ̊C and 40 ̊C for 9 times, there is no observable change in

swelling and shrinking magnitude of the homogeneous stimuli-responsive hydrogel sheet. In

terms of responsive time, for a hydrogel sheet which is 200μm thick and 5mm wide, shrinking is

completed in 60 seconds while swelling requires 10 minutes to finish, shown in Figure 14(d, e).

It is because water is actively pumped out from the hydrogel while the hydrogel becomes

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insoluble and collapses. On the other hand, swelling requires water to penetrate the hydrogel

matrix via diffusion, which takes a much longer time

Stimuli-responsive particles are doped into the alginate solution at a high intensity (0.089g/mL

dry mass) such that the particles are closely packed in the hydrogel matrix after gelation. Stimuli-

responsive particles filled up the void spaces between alginate fibers. Figure 14f shows a

Scanning electron microscope (SEM) image of the alginate matrix produced by extrusion. My

hypothesis is that the particles and the alginate fibers are connected by interpenetration. If the

particles are trapped in alginate matrix rather than interpenetrated, shrinking of particles will

release water into the matrix but will not induce an overall volumetric change of the hydrogel.

SEM images confirmed this hypothesis (Figure 14g). The alginate fibers penetrate the particles

and form a “branches and leaves” structure, whereas alginate fibers are the branches and the

particles are the leaves. In terms of the type of connection between alginate and PNIPAM-VAA,

they are physically winded instead of ionic bonded. Experiment evidence shows that doping

PNIPAM particles will produce the same temperature responsiveness as doping PNIPAM-VAA

particles. This means Ca2+ did not establish connection between the carboxylic groups in VAA

and the carboxylic groups in alginate. In a case such connection is established, the temperature

responsiveness produced by doping PNIPAM-VAA will be stronger than doping PNIPAM.

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Figure 13. Schematic of homogeneous stimulus responsive sheet extrusion. The stimulus responsive hydrogel solution is driven by a syringe pump into a microfluidic device which evenly distributes the solution into a planar flow. The solution gels after exiting from the device and entering a bath of crosslinker solution.

Figure 14. Homogeneous stimuli-responsive hydrogel and its response to temperature and pH.. (a) Homogeneous stimuli-responsive hydrogel tube responses to temperature changes. Lc is the length of tube at room temperature. Lh is the length of tube at 40°C. Homogeneous tube was extruded at 4mm/s in 100mM Ca2+ solution. The hydrogel tube is 150μm thick and 2.4mm in outer diameter in swell state using the homogeneous tube extrusion device shown in Figure 1a at QI=650µL/min, QM=200µL/min and Qo=650µL/min. Particle density in the hydrogel solution is 0.089g/ml (dry mass) Scale bar: 2mm. (b) Length change of stimuli-responsive hydrogel material in response to pH and temperature, with the comparison with hydrodynamic diameter changes of stimuli-responsive particles. pH3 Particle Diameter, pH7 Particle Diameter, pH10

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Particle Diameter, pH3 Material Length pH7 Material Length, pH10 Material Length. The measurements in (b-g) are based on extruded sheets using the setup shown in Figure 13 using 100mmol CaCl2 crosslinking solution. The sheet was extruded at a velocity of 4mm/s. n=3 (3 hydrogel sheets extruded at different times). (c) Length change of stimuli-responsive hydrogel material in cycles of temperature change, showing the reversibility of the stimulus response. Temperature is switch from 25ºC to 40ºC back and forth, starting at 25 ºC, n=1. (d) Length changes over time while hydrogel material is swelling, n=3 (3 hydrogel sheets extruded at different times). (e) Length changes over time while hydrogel material is shrinking, n=3 (3 hydrogel sheets extruded at different times). (f) SEM image of alginate matrix. Network of alginate fibers is visible in the image. Scale bar: 2µm. (g) SEM image of alginate matrix doped with particles. Individual particles can be seen in this image. Scale bar: 2µm. Data points and error bars are mean values and standard deviation of the n measurements for b-e.

Tubular structure is one of the most commonly seen 3D structures in natural tissues and

organisms, namely vascular structure. It is a perfusable structure which allows effective deliver

of liquids. Nature uses this structure to distribute nutrition across large tissues and organisms. By

creating a stimuli-responsive tube, the flow rate of liquid can be automatically adjusted. For

example, by using a pH responsive tube which shrinks in low pH, flow rate of ions can be

automatically regulated, as the flow rate decreases at low pH and increases at high pH due to

stimuli-responsive diameter changes of the tube. Heterogeneous tubular structure made of

stimuli-responsive soft material will possess the ability of change flow direction according to the

stimuli it receives. By designing the spatial composition along the peripheral of the tube, the

bending direction of the tube can be designed for selective stimuli. Moreover, stimulus

responsive tubes can be assembled easily by shrinking one end of a tube and insert into another

tube as discussed in Figure 15.

To demonstrate the ability of assembly stimulus responsive tubes, by taking advantage of

stimulus responsiveness, the following experiment is conducted. A stimulus responsive tube is

cut into two segments. One segment is immersed in hot water bath of 50°C for 5mins. The heat

segment is shrank in diameter and then transfer to the room temperature water bath where the

other segment is residing. The swelling of the tube needs 15mins which is relatively slower than

shrinking. Therefore, there is enough time to insert the shrink tube segment into the tube segment

of the original size. After 15mins, the shrink tube is fully swelled and went back to its original

size. As the result, the inserted tube segment is sealed inside the other tube segment. Figure 15

shows a schematic diagram of the tube assembly process.

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In the case that a Janus tube is made that only half of the tube is stimulus responsive, a curvature

change can be induced by induce stimulus response using temperature change. As one side of the

tube is swelled/shrank, internal stresses are induced in the tube and the tube will bend. In the

example shown in Figure 16, the curvature of the tube is changed between 0.060mm-1 and

0.138mm-1, which is a 231% change of curvature. The concept of extruding patterned stimulus

responsive hydrogel was first validated using in-plane extrusion of heterogeneous hydrogel

sheets. Heterogeneous hydrogel sheets can be patterned such that the shape deformation is

predictable. Details of heterogeneous stimulus responsive hydrogel sheets are described in

appendices.

Figure 15. Schematic of stimulus responsive tube assembly. A homogeneous stimulus responsive tube is cut into two segments. One segment is heated and shrank so it can be inserted into the other segment which is unchanged. The tube was extruded at QI=650µL/min(100mmol CaCl2 solution,), Qm,=67µL/min(1.5% alginate, 0.087g/mL PNIPAM-VAA particle solution) and Qo=650µL/min (100mmol CaCl2 solution,)

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Figure 16. Temperature induced response of half-half stimulus responsive tube. The “Janus” tube with half of the tube stimulus responsive was extruded using the device shown in Figure 1c. Upon heating, the tube bends because of the shrinking of the stimulus responsive half. The process is reversible. QI=650µL/min(100mmol CaCl2 solution,), Qm,1=67µL/min (2% alginate solution), Qm,2=67µL/min(1.5% alginate, 0.087g/mL PNIPAM-VAA particle solution) and Qo=650µL/min (100mmol CaCl2 solution,). (Scale bars 2mm)

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7 SUMMARY Tubular constructs of engineered biomaterials increasingly gain attention as they are potential

therapeutic replacements of human tissues. In vitro, they are extensively studied as platform of

cell-cell and cell-matrix interactions. The methods described in this thesis demonstrated

continuous extrusion of hydrogel tubes using a microfluidic platform. Tubular structures were

formed through off-plane extrusion of coaxial flow of biopolymer solutions and crosslinker

streaming solutions. The developed system is robust that it operates at a wide range of flow rates

(0-3500µL/min for streaming solutions). The throughput of tube extrusion is up to 20mm/s. By

controlling the flow rates of streaming solutions, the diameters and wall thicknesses of the

extruded tubes can be tuned predictively. The produced tube outer diameter ranged from 0.7mm

to 2.5mm. Alginate and collagen I hydrogel tubes can be consistently formed by coupling with

suitable crosslinkers in the streaming solutions. The flexible and automated nature of this tube

extrusion approach should enable formation tissue constructs of many cell types at different

scales.

Tube morphologies such as crimped and corkscrew shape inner tube walls were produced by

bringing the flow rates ratio of inner and outer streaming solutions to extremes. Spatial

composition heterogeneities along radial (multilayer tube), axial (tube with stripe patterns) and

circumferential (“Janus” tube) directions were incorporated in the extruded alginate tubes using

the same system. These morphologies and heterogeneities may lead to develop tubular tissue

constructs mimic natural tissues in terms of morphologies and spatial organization of different

cells.

Homogeneous and heterogeneous stimulus responsive tubes were extruded using the same

system, by doping synthesized PNIPAM-VAA nanoparticles in alginate solution. The

nanoparticle doped alginate based hydrogel was characterized in terms of its volumetric change

in changing temperature and pH, responsive time and reversibility of the reaction. Assembly of

homogeneous hydrogel tubes using their volumetric stimulus responsive change and bending of

“Janus” stimulus responsive hydrogel tube was demonstrated. This work is an important step

towards applications such as high order hydrogel construct assembly, responsive actuators and

flow regulators.

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8 REMAINING WORK AND FUTURE DIRECTIONS As formation of homogeneous and heterogeneous tubular structures was extensively discussed in

the context of alginate, other materials need to be tested to validate the results obtained using

alginate. For example, the experiments on collagen tube formation can be explored to evaluate

operational parameters of collagen using the same system. Introducing heterogeneity in collagen

tubes was not validated and may be tested in future experiments.

Methods of fixation and perfusion of produced hydrogel tubes should be developed in order to

utilize the different types of tubular structure produced. Hosting of tubes can be achieved by

using reversible sealed microfluidic devices or cannulation. Quantitative measurements such as

elasticity of the tube wall, diffusivity of different biomaterials and magnitude of stimulus

responsiveness can be evaluated if the tubes can be hosted and perfused in a controlled manner.

Biocompatible materials were used to form these tubular structures. Culturing different cell types

in the tubular structures can be tested, in order to study properties such as viability of cells,

migration of cells and biodegradation of the tubular structures. For example, endothelial cells can

be cultured to form vascular networks. Induced pluripotent stem cell can be cultured to and

developed in to various cell types on the tubular structure. Multiple cell types can be introduced

into the heterogeneous tubular structures at the same time to study cell-cell interactions.

Moreover, methods of assembling the tubular structures into complex 3D structures as building

blocks can be developed. Perfusable scaffolds of biomaterials can be made based on the

produced tubes.

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Appendices

A.1 Tube extrusion without crosslinker in the inner stream

Alginate tubes are extruded using the setup in Figure 2. The inner streaming fluid contains no

calcium ions. The crosslinking of alginate completely relies on diffusion of calcium ions from

the outside of the tube. The gelation of the inner wall is relative slow than the case where

calcium ions are present in the inner streaming fluid. Alginate is free to move in the lumen of the

tube and hence formed the structure shown in Figure A1b.

Figure A1. Tube extruded with no crosslinker in inner streaming fluid. (a) Confocal image of tube cross section extruded at QI=100µL/min (DI water), QM=200µL/min (2% alginate soltuion) and Qo=1200µL/min (100mmol CaCl2 solution). (Scale bar 500µm) (b) Confocal image of tube cross section extruded at QI=100µL/min (DI water), QM=200µL/min (2% alginate soltuion) and Qo=1200µL/min (100mmol CaCl2 solution). (Scale bar 500µm)

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A.2 Spotted tubes

Figure A2. Spotted tubes. (a) shows a fluorescence image of large spots produced by subjecting the spotting inlet channel to a pressure of 1 Psi during a solenoid valve opening time of 100ms. QI=500µL/min (100mmol CaCl2 solution), QM=200µL/min (2% alginate soltuion) and Qo=500µL/min (100mmol CaCl2 solution). (Scale bar 500 µm) (b) shows a fluorescence image of small spots produced by 200ms opening time of spotting channel with 0.6Psi. (Scale bar 500 µm) (c) shows a fluorescence image of alternating spots produced by 100ms opening time of spotting channel with 1Psi. (Scale bar 1 mm) (d) shows a fluorescence image of 2 rows of parallel alternating spots produced by 100ms opening time of spotting channel with 1Psi. (Scale bar 1 mm).

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A.3 Comparison of dimensions of single layer/three-layer tubes

Figure A3. Comparison of dimensions of single layer/Three-layer tubes. The format of the x-axis label is QI/QM(Qm,1/ Qm,2/ Qm,3)/QO. n=3 (3 measurements from same device). The tubes are made of alginate where the streaming fluids are 100mmol CaCl2 solution and the biopolymer solution is 2% alginate solution.

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A.4 One-step formation of heterogeneous stimuli responsive hydrogels

(PNIPAM-VAA) particles with large swelling ratios in response to stimuli (i.e., pH and

temperature changes) are dispersed as a payload in alginate solution before gelation. This hybrid

hydrogel can be controllably incorporated within a base hydrogel by a microfluidic device. Both

hydrogel materials undergo gelation after leaving the microfluidic device and exposing to a

solution containing crosslinkers which are Calcium ions at room temperature. The thickness of

the hydrogel sheet can be controlled by tuning the rotation speed of the drum which guides the

sheet at the device exit. Mosaic hydrogels with different tessellations can be produced by

controlling the pressure of each inlet of the device. The schematic diagram illustrates the

extrusion process of mosaic hydrogel sheet (Figure A4a).

There are 3 devices made for simple heterogeneous stimuli-responsive hydrogel sheets. For all

the devices, the channels are designed such that the flow resistances for all channels are the same

in each device. 1) Bilayer extrusion device (Figure A4b): this device has two layers on top of

each other. Each layer has two inlets which join at a T junction in the device. Different solutions

can be flowed in each layer by controlling the pressure of each inlet. This device is aimed to

extrude a bilayer hydrogel sheet where one layer is stimuli-responsive and the other is not. 2)

Side stripes extrusion device (Figure A4d): this device is designed to extrude hydrogel sheets

which has different materials at the center and at the two sides. The width of the central region

and the side stripes can be adjusted by tuning the relative flow rates of hydrogel solutions. For

example, while the flow rate in central channels is increased, the central region of the hydrogel

sheet will become wider. The purpose of this device is to extrude hydrogel sheets that the central

region is inert while the two sides of the hydrogel sheets are stimuli-responsive. 3) Side by side

extrusion device (Figure A4f): This device has two inlets and each inlet is responsible for

extruding one side of the hydrogel sheet. Therefore, the materials on each side of the hydrogel

sheet can be different. The width of each side of the hydrogel sheet is again controlled by flow

rate in each inlet. This device was made to extrude hydrogel sheets which one side is stimuli-

responsive and the other side is not.

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Bilayer heterogeneous hydrogel sheet can bend in response to external change of temperature or

pH (Figure A4c and Figure A5a). While one layer of the sheet is contracting and the other layer

remains unchanged, stress is created between two layers and causes the hydrogel sheet to bend.

The higher degree of contraction in the stimuli-responsive layer, the higher the curvature of the

hydrogel sheet produced by stresses. The curvature of the hydrogel sheet is a function of

temperature (Figure A5b). The curvature measurements were taken from 3 separately extruded

hydrogel sheets.

Stimuli-responsive stripes at sides hydrogel sheet bends globally and form saddle-like structure

locally due to the internal stresses caused by shrinking of the stripes at the two sides (Figure A4e)

[68].

Side by side heterogeneous hydrogel sheet buckles and form a wavy structure on the side of

hydrogel which has lower elastic modulus, while one side of the hydrogel is shrinking and the

other side remains unchanged (Figure A4g and Figure A5c) [69]. This morphology is also caused

by the stress at the interface of different materials, which is a result of swelling ratio difference

between the two sides. The wavelength of the wave structure decreases as the internal stress

increases. This implies that the wavelength is a function of temperature Figure A5d).

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Figure A4. Heterogeneous hydrogel extrusion. (a) Overview of extrusion of heterogeneous hydrogel with designed stimulus responses using a microfluidic printing device. Stimulus responsive and non-responsive solutions are driven into a microfluidic device with spatially align the solutions before gelation which happens at the exit of device. The extruded sheet is collected by a rotating drum. (b) Microfluidic device for double layer hydrogel sheet formation. The red layer is on top of the blue layer. There are two inlets for each layer which allow switch of stimulus responsive and non-responsive hydrogel solution within a layer. (c) Dual layer with one

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side stimulus responsive and other side inert roll up upon heating. (d) Microfluidic device for stripes at sides hydrogel sheet formation. (e) Hydrogel sheet forms wave and saddle shape and overall bending upon triggered response. (f) Microfluidic device for side by side stripes hydrogel sheet formation. (g) Wave and spiral shape formed on one side of the sheet while response of the material is triggered.

Figure A5. Heterogeneous stimuli-responsive hydrogel sheets and their response to temperature and pH. (a) Heterogeneous stimulus responsive hydrogel sheet rolls while temperature is changed from at 25ºC to 40ºC at pH 3. The sheets used in this figure was extruded at 4mm/s. The stimulus responsive hydrogel solution consist 1.5% alginate and 0.087g/mL PNIPAM-VAA. The non-responsive hydrogel is 1.5% alginate. The crosslinker used is 100mmol CaCl2 solution. (b) Curvature change of double layer heterogeneous stimulus responsive sheets. The curvature of the sheet increases upon heating and decreases while cooled. The environment is pH 3. n=3(3 hydrogel sheets extruded at different times). (c) Side by side hydrogel sheet forms spiral shape on one side while temperature is changed from at 25ºC to 40ºC at pH 3. (d) Side by side hydrogel sheet (spiral) wavelength vs. temperature at pH 3. The wavelength decreases as the temperature increases. n=9 (3 waves was measured from each hydrogel sheet for 3 sheets extruded at different times). Data points and error bars are mean values and standard deviation of the n measurements for b and d.