tio2 based nanosurface modification of stainless steel (ss...

Amrita Centre for Nanosciences and Molecular Medicine 21

Chapter 4

TiO2 Based Nanosurface Modification of Stainless Steel (SS) Substrates and Coronary Stents

C.C. Mohan, Sujish Kurup, John Joseph, Manitha B Nair, M. Vijayakumar, K.P.

Chennazhi, Shantikumar V Nair and Deepthy Menon. Feasibility of titania nanotexturing

approach to modify stainless steel coronary stents: In-vitro and in-vivo biocompatibility

evaluation (To be communicated to Biomaterials)

Amrita Centre for Nanosciences and Molecular Medicine Page 122

4.1 Introduction

Coronary artery stent implantation has become a more popular treatment

modality for coronary heart diseases such as atherosclerosis (commonly known as

heart blocks) due to its minimal invasiveness for patients undergoing percutaneous

coronary intervention. Current limitations in coronary stenting viz., in-stent

restenosis, and late thrombosis, are all mediated by endothelial damage or

dysfunction. A rapid restoration of a native functional endothelium on the stent

surface (i.e., re-endothelialization) is a possible solution to all the unresolved issues

pertaining to coronary stents, thereby achieving a high patency rate. Incorporation of

nanotopographical cues on metallic substrates would be a plausible approach to

enhance cellular proliferation, selective cell growth and thereby aid in re-

endothelialization. This part of the thesis encompasses the prospect of generating a

stable bioactive nanostructured coating on the currently used bare metal stents to

promote endothelialization. Moreover, stable passive nanocoatings can mask the

stent material surface from corrosive in-vivo environment and prevent the leaching

of toxic ions such as Ni, which is one of the main reasons for allergic reactions in

patients after stent implantation.

In the earlier chapters, the possibilities of nanotechnology combined with

the excellent surface properties of TiO2 were adopted as a potential strategy to

address the question of improved vascular cell response. Hemocompatible TiO2

nanostructures of distinct morphologies such as nanoleaves developed through a

simple hydrothermal route demonstrated better endothelial response in-vitro, coupled

with reduced smooth muscle cell proliferation. Incorporation of such uniform titania

nanotopographies with a specific nanomorphology (of submicron dimension) on

clinically used metallic stents, capable of inducing a differential effect on endothelial

growth and SMC proliferation (critical factor for restenosis), could significantly

improve its bioactivity. In the previous chapters, it could be concluded that amongst

various nanotopographies investigated, the nanoleafy topography promoted an

optimal vascular response, endothelial functionality and hemocompatibility. Hence,


in this part of the thesis, the focus is on assessing the feasibility of employing this

titania based nanotexturing approach for surface modifying the most widely used

stent material, viz., stainless steel. Titania nanostructured SS substrates with

nanoleafy topography were tested in-vitro to evaluate its endothelialization potential,

functionality and smooth muscle cell response. In addition, the ability of the stable

oxide layer to limit the underlying SS substrate from corrosion by preventing the

leaching of toxic ions such as Ni into the surrounding was assessed. The stability and

corrosion behavior of coatings on SS was also evaluated by various mechanical and

electrochemical tests. A detailed analysis of the coated stents was also carried out

following the ISO guidelines as a proof of concept for its clinical application.

4.2 Major research questions and hypotheses

RQ1. What is the effectiveness of the hydrothermal technique in successfully

translating TiO2 nanostructures on to SS as obtained on metallic Ti?

Hypothesis: Establishment of a thin Ti film over the SS metallic surface and its

subsequent hydrothermal treatment can yield nanostructured morphology as obtained

on metallic Ti, for desirable cell response.

RQ2. How can the presence of a Ti precursor and a Ti seed layer influence titania

nanostructuring on SS?

Hypothesis: The presence of Ti precursors in the alkaline solution during

hydrothermal processing can aid in crystal growth onto the Ti seed layer producing a

homogeneous nanotextured and crack-free crystalline coating that covers the entire

SS substrate.

RQ3. How can the above surface modification technique be extended to clinically

used stents?

Hypothesis: Generating a seed layer of Ti on SS stents followed by its hydrothermal

processing in the presence of suitable Ti precursors can yield an intact nanopatterned

TiO2 coating which is stable and adherent upon mechanical deformations of the

stent.


4.3 Materials and methods

4.3.1 Material

Medical grade 316L Stainless steel (SS) plates of 21mm diameter were

procured form Jayon Surgicals Pvt. Ltd., India. The plates were ultrasonically

cleaned in acetone, mechanically polished up to #1200 grit, followed by polishing

using 0.05 µm alumina suspension to get a surface finish of < 50nm. The clinically

used bare metal SS stents (Crypton coronary stents) with a diameter of 3 mm and

length 10 mm employed for the study were a kind gift from Meryl Life Sciences,

India.

4.3.2 Hydrothermal processing

Prior to hydrothermal surface modification, SS substrates were

chemically treated with a mixture of H2SO4 and H2O2 mixture (piranha solution) and

washed several times in distilled water to completely remove the piranha residues.

Piranha treated SS substrates were subjected to magnetron sputtering in an inert

argon environment inside a sputter coater(K550X Sputter Coater, Quorum

Technologies) to generate a layer of Ti on SS at a sputtering current of 120mA for

1h using pure Ti target (99.99%, Titanium Sputter Target - 57mm x 0.5mm thick).

Nanofabrication was carried out by a hydrothermal process as described in the

earlier chapters 1. Briefly, for hydrothermal treatment, Ti coated SS were immersed

in NaOH solution taken in a Teflon chamber which was housed in a stainless steel

autoclave. The entire assembly was placed in a high temperature furnace and

subjected to the prior optimized processing conditions for getting nanoleafy (NL)

morphology on SS (1M NaOH concentration for 4h at 200°C). 0.05% Titanium

isopropoxide (99.9% pure, Sigma Aldrich, USA) as added as a Ti source/ precursor

into the medium before hydrothermal treatment to facilitate the processing. After

hydrothermal treatment, all samples were dried at 60°C for 1 hr and ultrasonically

cleaned in distilled water.

To deposit a uniform Ti layer on SS coronary stents, a rotary apparatus

was designed indigeneously which was mounted inside the sputtering chamber of the


magnetron sputtering unit (Fig 4.1). To ensure uniformity in Ti deposition, the stents

were placed onto a needle mounted to the rotating shaft of a DC motor, whose

rotation speed was set to 30 rpm. To obtain a uniform surface coating of Ti on the

stents, they were balloon expanded to nearly 50% before Ti deposition. Further to Ti

deposition, the Ti coated stents were also surface modified to generate the nanoleafy

morphology via a precursor mediated hydrothermal processing as described above.

Fig 4.1 A rotary apparatus designed to ensure uniform stent coating with titanium inside a

magnetron sputtering chamber in Argon ambience

4.3.3 Surface characterization- SEM, EDS, XPS and XRD

Hydrothermally modified SS substrates and stents were characterized for

their surface morphology and texture using scanning electron microscopy (SEM;

Model: JEOL JSM-6490L) at different magnifications. Crystallinity of the samples

was analyzed using a glancing angle X-ray diffractometer (PANalytical XPert PRO

system) fitted with Cu-Kα radiation (λ =1.5414 Å). The scan was carried out in the

range of 10-70o at a step size of 0.02. Surface compositional analysis of the samples

were further carried out using Energy Dispersive Analysis (EDS) and X-ray

Photoelectron Spectroscopy (XPS) (Model: ESCALAB220I-XL) over a binding

energy range of 0-1200 eV.


4.3.4 Cell proliferation studies on nanomodified samples using HUVECs

and SMCs

Human umbilical cord vein endothelial cells (HUVECs) and vascular

smooth muscle cells (VSMCs) were isolated from umbilical cord vein adopting the

protocol from Jaffe et al. 2. Cells were cultured in complete IMDM (containing 20%

fetal bovine serum, GIBCO, Invitrogen), 100 U ml-1

pen/strep antibiotic solutions

(GIBCO, Invitrogen, USA). 150 µg ml-1

Endothelial growth supplement (ECGS,

Sigma, USA) was added to the complete media for HUVECs and 0.2 µg ml-1 PDGF

for VSMCs.

Cell proliferation on nanomodified SS was studied in comparison with the

bare SS using Alamar blue assay at three different time points 24h, 72h and 120h.

HUVECs and VSMCs were seeded on various samples at a seeding density of 16000

cells/cm2 and incubated in 20% complete IMDM. At the end of each time point, cells

were incubated with 10% Alamar blue (Invitrogen Bioservices Pvt. Ltd, Bangalore)

in complete media for 4 h and the optical density was recorded using a microplate

spectrophotometer at 570 nm, with 600 nm set as the reference wavelength. Cell

number on each sample was obtained as an extrapolation from a standard graph

drawn for different seeding densities of HUVECs and SMCs, and the OD was

evaluated the same way using the Alamar Blue assay.

4.3.5 Fluorescent micrographic studies: F actin staining of ECs & SMCs and

PECAM1 staining on ECs

HUVECs and VSMCs were seeded on various samples at a seeding density

of 16000 cells/ cm2

and cultured for a period of 7 days in complete IMDM. Cell

grown samples were washed and fixed using 4% PFA in PBS at the end of

incubation. Subsequently, the cells were permeabilized using 0.5% Triton X-100 in

PBS for 5 min, and then blocked with 1% FBS in PBS for 15 min to limit the non-

specific binding sites in accordance with the manufacturer‘s protocol. HUVECs were

then incubated with Texas red conjugated Phalloidin (Molecular probes, Invitrogen)

for 60 min and FITC conjugated mouse anti- human PECAM1 antibody for 30 min


at room temperature to view the F- actin filamental assembly and localization of a

platelet endothelial cell adhesion molecule (PECAM1). VSMCs were stained using

Texas red conjugated Phalloidin and counter stained using DAPI to study their cell

morphology and coverage over the sample surfaces. The fixed samples were then

viewed under an Olympus BX-51 fluorescent microscope for analysis.

4.3.6 Nanoindentation and nanoscratch testing of the coating

Nanoindentation testing was carried out on titania surface modified SS

substrates using a Berckovich indenter with a triagonal pyramidal shape (Hysitron

Triboscope Nanoindenter), with nanoleafy titania substrate as a control. A diamond

tip of radius 200nm was pressed into the substrates at a maximum load of 10mN,

allowing a holding time of 10s followed by unloading. About 10-12 indents were

made on each sample and values of hardness and elasticity were calculated from the

unloading curve using Oliver Pharr method 3. For nanoscratch testing of surface

modified SS and Ti samples, an average of 3-5 scratches of length 10µm were made

on the sample surfaces by applying a progressive load range of 0.1 µN to 10mN at a

scan rate of 0.1 µm/s. Adhesion strength of the coating was determined as the

amount of force required for detachment of the coating from its substrate which is

represented in terms of the critical load 3.

4.3.7 Linear sweep voltammetry

Corrosion behaviour of surface modified samples in comparison to

unmodified controls was tested in-vitro by a linear sweep voltammetric technique in

PBS at pH 7.4 as an electrolyte. An electrochemical cell which consisted of a

platinum counter electrode, an Ag/AgCl reference electrode and a working electrode

was used for testing samples. The sample was held as the working electrode and

immersed in PBS to expose a surface area of 1 cm2. Testing was carried out using an

electro-chemical work station (Autolab, Metrohm) at a scan rate of 5mV/s across a

potential range of -1.5V to +1.5mV. The corrosion potential (Ecorr) and corrosion

current density (Icorr) were determined using a Tafel extrapolation method 4.


4.2.8 Ion release study

For evaluating the release of metallic Ni, Fe and Cr ions from SS during its

immersion in various physiological media, SS plates were first surface modified as

detailed earlier to generate a homogeneous titania coating. The unmodified surface

was then masked using a resin to prevent ion leaching from that portion. Samples

were immersed in 12ml/cm2

of two physiological solutions, viz., PBS and HBSS at

pH 7.4 and incubated at 37oC to mimic body conditions along with mild agitation at

75 rpm for uniform mixing. 100 UL of antibiotic Pen Strep was added to prevent

bacterial growth on samples. Leaching of ions (Ni, Cr and Fe) from the samples was

analyzed at different time intervals of 1, 7, 14, 28, and 90 days. The percentage of

ion released from the samples with time was analyzed by inductively coupled plasma

atomic emission spectroscopic (ICP-AES) analysis. Untreated bare polished SS as

well as samples coated with the epoxy resin on both sides were used as controls in

this experiment. All the samples were done in triplicates.

4.3.9 Crimping expansion testing of coated stents

The stability and integrity of nanocoatings on stents was tested by a

mechanical expansion and crimping of the stent using a balloon mounted

angioplastic catheter. Stents mounted on the balloon catheter were expanded to its

maximum diameter (3mm) by applying a 12 atm pressure (as per manufacturer‘s

instructions). Further, for the crimping test, nanomodified stents were manually

crimped back to its initial diameter (1mm) over the balloon catheter. Samples were

analysed using SEM to visualize any cracking or flaking of the coatings on stent

surface.

4.3.10 Durability testing of stents

Durability of the nanomodified coatings on SS stents was tested in an in-

vitro flow model by subjecting the stent to the similar strain conditions as

experienced inside coronary arteries and together with an accelerated high shear

stress on stent surface for a period of 3 months. Stents were expanded to a maximum

of 3.3 mm using a balloon catheter to induce an over stretching of about 10-15%,


inside a latex tubing (3mm inner diameter) by applying 18 atm pressure. This lead to

an increase in tube internal diameter by 1- 1.5% which in turn influenced the preload

on stents that helped in emulating the cyclic loading conditions experienced in

coronary arteries. Viscosity of the circulating fluid was adjusted to 9 cP (9 x10-3

NS/m2) by dissolving 5% dextran in distilled water. This viscous fluid was used for

circulation inside the tube at a flow rate of 2ml/s for imparting a high shear condition

on the stent surface. Wall shear stress on stented section was calculated from the

known tube radius and flow rate by using Hagen Poiseullie equation 5.

𝑊𝑆𝑆 =4𝑄𝜇

𝜋𝑅2

Where, Q is the flow rate in ml/s, µthe fluid viscosity in NS/m2

, R is the inner

radiusof the tube (0.15cm). The frequency of circulation was fixed at 1.6 cycles/s for

3 months to complete a total of ~10 million cycles, which was equivalent to 4

months of in-vivo implantation. Stents surfaces were analysed for the stability of

surface morphology using SEM and the circulated fluid was analyzed by ICP for any

wear debris.

4.3.11 Radial stiffness and trackability testing of stents

Radial stiffness of the modified and bare SS stents was tested according to

European standards (prEN 12006-3) by enclosing it in a tube that simulates a blood

vessel that separated the stent from a temperature-controlled, water-filled test

chamber 6. The tube simulating the blood vessel is of polyurethane (PUR) and the

test chamber for measurement was sealed with a pressure-tight cover and connected

with a tubing system to the pressure controller (GDS Pressure-volume controller,

USA). The radial stiffness measurement was performed by determining the outside

diameter and elastic recoil after expansion, without touching the specimen inside the

water bath using a High-speed, High-accuracy Digital Laser Micrometer

(KEYENCE, LS-7030MT). Stent expansion was computer-controlled by a pressure

controller. The radial stiffness values were measured by determining the pressure

at which the stent can no longer resist the pressure the vessel exerts upon it.


Trackability testing of nanomodified stent in comparison with bare SS stent

was carried out inside a water filled chamber with a custom-made design to provide

a tortuous path for the stents. Stents were firmly crimped on a balloon catheter

before being used for the test. Catheter mounted stents were passed through the

winding path inside the chamber as a check for its flexibility. The easiness of the

stent to move from one end of the chamber across the tortuous trail to the other end

is considered as a measure of its trackability.

4.4 Results and discussion

4.4.1 Nanosurface modification

Surface nanotexturing of titanium coated SS samples was carried out by a

simple hydrothermal (HT) processing at 200˚C for 4h in 1M NaOH as detailed

earlier, to generate a uniform nanoleafy titania topography. The thickness of the Ti

layer deposited on SS was measured by surface profilometry to be 300±25nm.

However the seed Ti layer on SS alone could not yield a homogeneous surface

texture after HT treatment as evident from the SEM image depicted in Fig 4.2A.

Here, the percentage of Ti deduced from EDS analysis was found to be very low

(1.2%) (Fig 4.2B). Hence, to obtain a uniform surface nanotexture as well as to

enrich the Ti content, an additional Ti precursor was added to the reaction medium.

It was observed that the addition of 0.05% titanium isopropoxide into the medium

served as specific sites for heteronucleation during hydrothermal processing 7

and

aided the subsequent crystalline growth of titania on SS substrate. This is clearly

evident from the SEM image depicted in Fig 4.2C wherein a uniform and defined

nanoleafy pattern was obtained on the piranha treated SS surface (represented as SS

Ti NL), upon HT treatment in presence of a precursor. This was very similar to the

nanotopography obtained on metallic Ti (Ti NL). The EDAX spectra (Fig 4.2D) also

clearly indicated an elevated Ti content on the precursor treated SS substrate.


Fig 4.2 Morphological and compositional analysis of TiO2 nanostucturing on 316L SS by SEM

and EDS (A, B) without precursor addition and (C,D) with the addition of precursor

Table 4.1 XPS analysis showing the atomic percentages of Ti, O and different alloy components

on bare and TiO2 coated nanomodified samples

Structural characterization of the nanomodified surfaces was deduced using

XRD, EDAX and XPS. Glancing angle XRD patterns of nanomodified SS samples

in comparison to bare SS (Fig 4.2A) clearly indicated the presence of TiO2 on the

surface, marked by the appearance of an anatase peak at 25.3° (Fig 4.3B). The two

Elemental

composition

Atomic percentage

SS Ti NL Bare SS

Fe 2p 0.1 21.0

Cr 2p 1.0 5.2

Ni 2p 0.1 0.0

Mo 3d 0.0 0.2

O 1s 54.0 41.2

C 1s 28.7 32.2

Ti 2p 16.0 0.0


high intensity γ-austenitic peaks of bare 316L SS at 2θ values of 44° and 51.3° were

significantly masked on nanomodified SS by the formation of the TiO2 coating on

their surface.

XPS further confirmed an increased percentage of Ti on nanomodified

surfaces compared to bare SS (Table 4.1). This was again reaffirmed by the surface

compositional analysis by XPS. The wide scan XPS spectra depicted in Fig 4.3C

confirmed the complete masking of the alloy components viz., Fe2+

, Cr3+

, Ni2+

with

the appearance of a new Ti 2p peak at 459 eV and also by the higher atomic

concentration of Ti and O on nanomodified SS compared to bare SS. This was

additionally confirmed by the high resolution XPS spectrum of Ti 2p core level

which showed a strong doublet peak at 458.8 eV and 464.4 eV which can be

attributed to the 4+ oxidation state of Ti as in TiO2 (Fig 4.3D).

Fig 4.3 Compositional analysis of nanostructured SS via precursor mediated hydrothermal

processing using XRD and XPS (A) Glancing angle XRD showing the appearance of an anatase

A

22 24 26 28 30

2 theta

Bare SS

SS Ti NL

A

Bare SSSS-Ti-NL

A- Anatase●- 316L: ɣ-austenite ●

●

●●

2Ө2Ө

SS-Ti-NL

A B

C D

Amrita Centre for Nanosciences and Molecular Medicine 3

peak on nanostructured SS compared to bare 316LSS (B) An enlarged view of the anatase peak

(C) Wide spectrum XPS analysis showing the masking of alloy components on modified SS with

the appearance of a Ti peak (D) High resolution XPS spectra of Ti 2p showing the presence of

Ti in its 4+ oxidation state as in TiO2.

By adopting the same methodology as above, coronary SS stents were also

successfully surface modified to generate a uniform nanostructured titania coating

with well defined nanoleafy patterns on their surface, as evident from Fig 4.4B.

Featureless unmodified SS stents are depicted in Fig 4.4A at different

magnifications. EDAX spectra confirmed the presence of TiO2 on the nanomodified

surface compared to bare SS stents (Fig 4.5A, 4.5B).

Fig 4.4 Nanostructuring of SS coronary stents via precursor mediated hydrothermal processing

(A) Bare SS stent at different magnifications (B) TiO2 nanostructured SS stent at different

magnifications (SS Ti NL). Higher magnification image clearly shows the appearance of

nanoleaf-like patterns on their surface.

Ai Aii Aiii

Bi Bii Biii


Fig 4.5 Compositional analysis of SS coronary stents by EDS (A) Bare SS (B) SS Ti NL

4.4.2 Cell material interaction studies

Interaction of vascular cells viz., HUVECs and VSMCs, with nanomodified

SS substrates (plates) was quantitatively analyzed by a cell proliferation assay using

Alamar blue, at three time points 24h, 72h and 120h. Vascular cell proliferation is

graphically represented as cell numbers at different time points in Fig 4.6.

Fig 4.6 Vascular cell response on nanomodified and bare SS substrates Cell proliferation studies

by Alamar blue assay using (A) HUVECs and (B) SMCs. (C) Functionality analysis of HUVECs

on SS Ti NL by NO release assay

0.00 1.00 2.00 3.00 4.00 5.00 6.00 7.00 8.00 9.00 10.00

keV

0

50

100

150

200

250

300

350

400

450

500

550

600

Counts

CK

aO

Ka

SK

aS

Kb

CrL

lC

rLa

CrK

a

CrK

b

MnL

lM

nL

a

MnK

a

MnK

b

FeL

lF

eL

a

FeK

esc

FeK

a

FeK

b

MoL

lM

oL

a

0.00 1.00 2.00 3.00 4.00 5.00 6.00 7.00 8.00 9.00 10.00

keV

0

200

400

600

800

1000

1200

1400

1600

1800

2000

Counts

OK

aT

iLl

TiL

a

TiK

a

TiK

b

CrL

lC

rLa

CrK

a

CrK

b

FeL

lF

eL

a

FeK

esc

FeK

a

FeK

b

NiL

lN

iLa

NiK

a

NiK

bMoL

lM

oL

a

Ti 4.7%A B

0

10000

20000

30000

40000

50000

Bare SS Ti NL Cover slip

Ce

ll n

um

be

r

Day 1

Day 3

Day 5

*

**A

Co

nce

ntr

atio

n o

f N

O (

µM

)

00.5

11.5

22.5

33.5

44.5

SS Ti NL BARE SS NC

Day 1

Day 3

*

**

Cover slip

B

** **

0

5000

10000

15000

20000

25000

Bare Ti SS Cover slip

Ce

ll n

um

be

r

Day 1

Day 3

Day 5

SS Ti NLBare SSC

Bare SS


HUVEC cell proliferation was found to be significantly enhanced on SS Ti

NL after 5th

day of in-vitro culture compared to bare SS (Fig 4.6A). However it was

observed that, the SMC viability was strikingly reduced on the nanomodified SS

compared to bare SS after 5 days (Fig 4.6B). Additionally, to assess the functionality

of HUVECs grown on surface modified vs unmodified SS substrates, NO release

from the endothelium formed was measured using a Griess assay. It was found that

endothelial cells grown on nanomodified SS showed a statistically significant

elevation in the NO levels compared to bare SS (Fig 4.6C). This is a clear indication

of the influence of specific TiO2 nanotopographical features viz., nanoleaves, in

promoting functional endothelialization of stent materials.

Furthermore, to characterize the intactness of the endothelium formed on

the test samples, the expression of PECAM1/CD31, which plays a key role in

endothelial cell– cell adherence and migration, was evaluated 8. Fig 4.7 depicts the

PECAM1 expression at cell junctions and cytoskeletal staining of HUVECs on SS Ti

NL in comparison with the bare SS after 7 days of in-vitro culture. Immunostaining

revealed that PECAM1 expression was notably improved on HUVEC monolayers

grown over SS Ti NL, with clear localization primarily towards their periphery (Fig

4.7A), similar to that observed on a gelatin coated cover slip which served as a

positive control (Fig 4.7B). However the higher magnification images of bare SS

substrates clearly showed a reduced PECAM1 expression at their cell junctions,

indicative of the lack of stable cell contact (Fig 4.7C). Thus a rapid coverage of an

intact endothelial monolayer was achieved on nanomodified SS after 7 days of in-

vitro culture, while the bare surfaces still remained devoid of a complete endothelial

coverage.


Fig 4.7 Fluoroscent micrographs showing the PECAM1 expression at cell junctions and

cytoskeletal staining of HUVEC monolayer on (A) SS Ti NL (B) gelatin coated coverslips and

(C) bare SS after 7 days of in-vitro culture at low (left panel) and high (right panel)

magnifications.


Fig 4.8 Fluorescent micrographs showing the cytoskeletal staining of SMCs (Red) on (A) bare

SS substrates in comparison to (B) nanomodified SS substrates after 5 days of in-vitro culture at

different magnifications [(i) - low and (ii) – high]. Cell nucleus is counterstained with DAPI

which appears blue.

In sharp contrast to this observation, the titania nanoleafy topography on

SS hampered the SMC proliferation, unlike on bare SS, where cells achieved a good

coverage with well spread morphologies for the same culture conditions in-vitro (Fig

4.8). Fluorescent micrographs of F-actin staining shown in Fig 4.8 depict cells with a

shrunken morphology on nanosurfaces (Fig 4.8B) unlike those on unmodified SS

(Fig 4.8A), which is indicative of its reduced viability. These results showed a very

good correlation with the quantification results of SMC cell proliferation (Fig 4.6B)

wherein the cells had minimum viability on the nanosurface. The results are in good

Ai Aii

Bi Bii

100µm100µm

100µm100µm


agreement with the previous data on similar hydrothermally created TiO2

nanostructures on Ti (Fig 2.14 & 2.16 in Chapter 2) that showed a preferential

endothelial growth with reduced SMC proliferation 1.

Analogous differential behavior of vascular cells on surface features with

submicron dimensions are reported on other substrates such as SS and TiO2

nanotubes in literature 9,10

. Such behaviors can be attributed to the changes in

mechanotransduction cascades that get initiated as a part of their stress relieving

mechanisms such as membrane stretching and cytoskeletal reorganization that occur

during cell spreading on nanosurfaces via the formation of focal adhesions 11,12

.

Mechanical stimulations experienced by cells can directly activate nucleus though

cytoskeletal assembly or indirectly through various downstream signalling, thereby

influencing events such as proliferation, migration, gene expression, protein

expression, cell cycle progression etc 10

.

4.4.3 Nanoindentation and nanoscratch testing

It is important that the coatings developed on the metallic surfaces withstand

the shear forces experienced during its contact with blood. To evaluate the adhesion

strength as well as other mechanical properties of titania films created on SS

substrates, nanoindentation and nanoscratch testings were carried out. In a typical

nanoindentation experiment, a controlled load was applied to the sample using an

indenter tip and the values depicting the mechanical characteristics of the sample

were obtained from a load‐displacement curve, as depicted in Fig 4.9A for the

nanomodifed samples. Here, the mechanical properties of nanostructured coating

developed on SS have been compared with the titania structures developed on

metallic Ti (discussed in Chapter 2). The loading and unloading curves for both the

nanocoatings exhibited a relatively smooth profile (without any discontinuity)

indicating that the coatings resisted crack propagation upon surface indentation. The

SPM images shown in Fig 4.9B provided visual evidence to the crack resistant

nanomodified surface upon nanoindentation.


Fig 4.9 Mechanical studies of TiO2 film on SS substrate by nanoindentation (A) Load

displacement graph showing the loading unloading curves of different samples (B)

Representative SPM image showing the tip imprint upon indentation (C) Table showing the

values of elastic modulus and hardness

Although literature reports ceramic coatings to be brittle in nature 13

, the

titania coating generated on SS through hydrothermal technique revealed excellent

flexibility and elasticity, perhaps due to its reduced thickness (320±25nm).

Nanoindentation results demonstrated a higher overall hardness and elastic modulus

for nanostructured titania coating on SS (E= 210.7±5 GPa, H= 4.3±0.1 GPa) and

nanomodified metallic titanium (E= 187.9±7 GPa, H= 2.8±0.4 GPa) as depicted by

the tabulated values given in Fig 4.9C. The nanomodified samples developed via the

aqueous hydrothermal route showed increased values of hardness as well as elastic

modulus than reported in literature 13,14

for TiO2 coatings. Sample processing through

hydrothermal treatment offers an annealing effect to the coatings which could impart

improved crystallinity and mechanical properties 15

.

Further, adhesion of coatings to its substrates was studied by nanoscratch

testing. Three scratches were made on each of the sample substrates by applying a

ramping load from 0.1µN- 10mN over a distance of 10 µm. Fig 4.10A and 4.10B

SamplesElastic modulus

Er (GPa)Hardness H (GPa)

SS Ti NL 214.3 3.6

Ti NL 202.6 3.1

0 100 200 300 400 500 600 7000

2000

4000

6000

8000

10000

12000

14000

Loa

d (

P)

Displacement (nm)

Bare Ti

Bare SS

Ti NL

SS Ti NL

A B

C


shows the representative SPM images of scratches made on SS Ti NL and Ti NL

respectively.

Fig 4.10 Adhesiveness of TiO2 film on SS by nanoscratch testing. Representative SPM image

showing the scratches made on (A) SS Ti NL (B) Ti NL. Graph showing the variation in force

along the scratch length on (C) SS Ti NL (D) Ti NL.

The critical load (Lc) in the scratch adhesion test is a measure of adhesive

strength in the absence of complete delamination of the film from the substrate 16

.

Low adhesion strength results in failure of the coating during scratching which is

shown as cracking on the trail region; and the load at which this is observed is

recorded as the critical load, Lc. The Lc value for SS Ti NL was deduced to be 4.3

mN at a distance of 5 µm, while for Ti NL the initial delamination of coating

occurred soon at an Lc value of 1.8 mN and before 1µm (Fig 4.10C, 10D) which is

A B

-4 -2 0 2 4 6 8 100

1000

2000

3000

4000

5000

SS Ti NL

La

tera

l fo

rce

(N

)

Lateral displacement (m)

Lc

C

-4 -2 0 2 4 6 8 100

1000

2000

3000

4000

5000

No

rma

l fo

rce

(N

)

Lateral displacement (m)

Ti NL

Lc

D


also evident from the SPM images. A higher Lc of SS Ti NL is suggestive of better

scratch resistance or adhesion strength of the titania coatings on SS substrates when

compared to that on metallic Ti. Magnetron sputter coated Ti underlayer on SS as

well as the piranha treatment of SS before Ti deposition might have provided better

adhesion strength to the TiO2 nanocoating as reported earlier in literature 17

.

Moreover, the nanoleafy topography (Ra= 81nm, Rq= 171) could offer good

resistance to the propagation of crack through the layer probably due to its high

elastic modulus, crystalinity and fracture toughness compared to the conventionally

used TiO2/ Al2O2 coatings 18

. SPM images also revealed no evidence of cracking or

flaking of the nanostructured coatings on both substrates, SS and pure Ti, while

scratching is indicative of their better mechanical stability and elasticity upon

hydrothermal treatment, all of which are crucial factors for its translation on to

coronary stents 19

.

4.4.4 In-vitro corrosion studies

Stainless steel biomaterials are prone to localized corrosion inside human

body over time which is a leading cause for inflammatory responses due to Ni ion

leaching 20

. Allergic reactions to the leached Ni ions could trigger the mechanisms

responsible for in-stent restenosis inside stainless steel coronary stents 21

. Improved

corrosion resistance of SS substrates could guarantee a better biomedical implant

response 22

. The present study employed an accelerated corrosion testing of various

TiO2 modified substrates such as Ti SS NL and Ti NL in comparison with bare

substrates of 316L SS and pure Ti in PBS. The corrosion behavior of test samples

were interpreted from the Tafel plot which depicts the variations in current density

across an applied potential range of -1.5 to +1.5V represented in Fig 4.11A. The

values of Icorr and Ecorr were obtained as an extrapolation from the Tafel plots and are

tabulated in Fig 4.11B. Bare SS was found to corrode faster with a lower Ecorr value

amongst different samples compared. Metallic Ti is well known for its good

corrosion resistance due to the presence of a native TiO2 layer indicated by a clear

shift towards the more anodic region and a substantially lesser current density 13

.

Similar trend was observed for titania modified samples of SS and Ti i.e., SS Ti NL


and Ti NL depicting higher corrosion potential along with reduced current densities,

indicative of an improved corrosion resistance of these surfaces. This was further

confirmed by the corrosion rate values, which were significantly lesser compared to

that of the bare SS (Fig 4.11B).

Fig 4.11 Corrosion studies on TiO2 nanomodified samples by linear sweep voltammetry (A)

Tafel plot showing the variation in current density plotted against a constant potential range (B)

Table showing the values of corrosion potential (Ecorr), corrosion current (Icorr) and corrosion

rate among various test samples

Additionally, the corrosion behavior was analyzed over a period of time upto

3 months by measuring the amount of ion released from the sample surfaces in

physiological conditions (pH 7.4). The ion release profile showed a gradual increase

in the amount of free Ni2+and Cr3+ ions in the release medium (PBS and HBSS)

from bare SS samples with increasing incubation time. Ni and Cr release was

substantially increased for bare SS after 14 days while their release concentration

from SS Ti NL was still retained at a low level (Fig 4.12 A, 12B). The higher

amount of ion release from bare SS can be correlated well with their relatively

higher corrosion rate in comparison to TiO2 nanomodified SS under physiological

pH. The resin coated negative control did not show any ion release from their surface

(data not included), confirming the masking of SS surfaces preventing the discharge

of ions.

-2.0 -1.5 -1.0 -0.5 0.0 0.5 1.0 1.5

-8

-7

-6

-5

-4

-3

-2

E(V)

log

i(m

A/c

m2)

Ti Bare

Ti NL

SS Ti NL

Bare SS

Samples E corr I corr Corrosion rate(mpy)

Bare SS -1.24 2.3X10-6 2.1X10-4

Bare Ti -0.58 0.2X10-5 4.9X10-5

SS Ti NL -0.47 0.4X10-5 9.8X10-5

Ti NL -0.32 0.2X10-5 4.9X10-5

BA


Fig 4.12 Graph showing the ion leaching from SS TI NL compared to bare SS as a part of the

corrosion behavior analysis for a period of 3 months (A) Nickel (Ni) ion leaching (B) Chromium

(Cr) ion leaching. Solid lines indicate ion release in PBS and dotted line indicates that in HBSS

release medium. **p<0.01w.r. to control SS.

4.4.5 Coronary stent testing

Bio and hemocompatible, inorganic titania (TiO2) is an ideal stent material

owing to its multifaceted characteristics which include its capability to provide an

inert coating on other metals like SS, as well as the desirable electrochemical

properties for reduced thrombosis, platelet adhesion, and immune response 23

.

However, stability and durability are most essential requisites for any implant

surface for its long-term existence in the living body 24

. Hence, coating durability

and stability are crucial for vascular implants as well 17

.

Inorganic stents coatings should resist cracking upon its mechanical

expansion and crimping. Conventionally used inorganic coatings are thicker of the

order of several microns, which are prone to cracking due to lack of elasticity and

stability 17, 25-27

. The present nanofabrication technique yielded thin titania coatings

(typically 320 nm) on SS which would impart stability to the coatings.

Nanostructuring of SS coronary stents via the precursor mediated hydrothermal

processing as done for the SS plate yielded a nanoleafy pattern which was uniform

and homogeneously distributed over the entire stent surface. The uniformity of the

nanotextured titania coating on the luminal and abluminal areas of the SS stent was

confirmed by SEM analysis at multiple sites on the inner and outer stent surfaces


(Fig 4.13). As seen in the image, a uniform coverage of the stent by a nanoleafy

texture is clearly evident at higher magnifications.

Fig 4.13 SEM micrographs depicting the uniformity of nanotexturing on luminal and abluminal

areas of the stent. Images were taken at multiple regions by specifically focusing inside and

outer stent surface.

Metallic coronary stents have to be expanded and/or crimped at various

stages of deployment. This necessitates the testing of the coatings upon stent

deformation. Hence, the nanotexured TiO2 coatings on SS stents were mechanically

tested for its ductility, flexibility and stability upon expansion and crimping of stents

to check its applicability for stenting. SEM images were taken at various stages of

expansion and crimping and are shown from different regions of the stent in Fig

4.14. Nanotitania coatings showed good stability on SS stent surface during

expansion and also crimping, without any evidence of cracking or flaking (Fig 4.14).

This may be attributed to the good elasticity and adhesion strength of the coating on

SS substrate.


Fig 4.14 SEM micrographs depicting the stability of TiO2 coating on stent surface upon

crimping and expansion. (A) Unmodified stent (B, C) Modified stents after 50%, 100%

expansion and (D) after hand crimping

Furthermore, it is imperative that the titania coating developed on SS

coronary stents is durable and stable under circulation over a period of time. To

check the durability of stent coating, stents were kept in a viscous medium to

emulate a high shear stress under flow as detailed in the methods section. An

increased WSS value of about 7 N/m2

was employed for subjecting the stents to an

accelerated durability testing compared to that experienced in normal coronary

arteries (1-2 N/m2)

5. This elevated WSS was achieved by increasing the viscosity of

the circulating fluid to three times that of blood (3.4cP). For a period of 10 years

further to implantation inside coronary arteries, stents will be subjected to nearly 400

million cycles under circulation. In our study, the coated stents were subjected to a

maximum of 10 million cycles over a period of 3 months under high shear

Aft

er

Cri

mp

ing

(1m

m, h

and

cr

imp

ing)

50

% e

xpan

sio

n

(1.5

mm

; 6 a

tm)

At

full

exp

ansi

on

(3m

m; 1

2 a

tm)

Di Dii Diii Div

Ai Aii Aiii Aiv

Bi Bii Biii Biv

Ci Cii Ciii Civ

Un

mo

dif

ied

5

0%

exp

ansi

on

(1

.5m

m; 6

atm

)


accelerated conditions, as a proof of concept for its durability. Titanium being

physically deposited and modified into a thin titania layer on SS, it was important to

check for any leaching of Ti from the stent surface under flow conditions. This was

assessed here by ICP analysis of the media at different time points, by aliquoting

samples and checking for the Ti content. The results of the 3 month study indicated

that, the concentration of Ti ions in the circulating viscous fluid increased from the

basal values over the study period of 90 days (Fig 4.15A). This might be due to the

leaching out of Ti ions from the SS stent surface under the constant high shear

applied in this experiment. However, the leaching profile of other ions such as Ni

and Cr was not found to increase significantly over time. Considering the fact that

this durability testing was only a one-time study and that it was not subjected to a

pulsatile flow condition mimicking those in human coronary arteries, the results

cannot be considered as strictly conclusive. A proper durability testing of modified

stents should be checked under an accelerated condition to complete 400 million

cycles which is equivalent to 10 years of implantation inside human body as per ISO

standards prior to implantation.

4.4.6 Radial stiffness and trackability testing

The measurement of radial stiffness was performed by measuring the

maximum pressure a stent can withstand before getting deformed (collapse

pressure), indicated by a change in its diameter which is measured accurately using a

laser micrometer. Stiffness measurements done on the polyurethane tube alone

without stents served as the control. Measurement performed on commercial

coronary SS stent showed a collapse pressure value of 280 kPa, while the

nanomodified stent yielded a value of 299 kPa, which was marginally higher than

the control stent, perhaps due to the annealing effect of the stent upon hydrothermal

treatment. This in turn implies that the radial stiffness, which is a significant

parameter for coronary stent in order to resist the inward recoiling pressure of the

arteries after stent deployment, is not significantly altered upon nanomodification.

High stiffness values are advantageous for coronary stents for its use as endovascular

supporting devices. Additionally, it is extremely important to determine how the


stent stiffness affects the longitudinal flexibility of the stent. This was qualitatively

assessed by a trackability experiment wherein the stent was made to move along a

tortuous path inside a custom-made water filled chamber. The trackability was

determined by the ability of the stent to move from one end of the chamber to the

other along the winding path. The high stiffness value also did not seem to affect the

flexibility of the nanomodified stent which was more or less comparable to that of

the unmodified SS stent. This in turn can reflect the easiness with which the surface

modified SS stent can be deployed inside the coronary artery, from the site of its

insertion, i.e., femoral/carotid artery.

Fig 4.15. Durability testing of nanomodified stents under high shear flow condition over a

period of 3 months. (A) The graph depicts the variation Ti ion concentration in the circulating

fluid from the stent coating at different time points. (B) Radial stiffness measurements of

nanomodified SS stents in comparison to unmodified SS stents

4.5 Conclusion

Simple, translatable nanoscale topographical features of distinct nanoleafy

morphology were developed on 316L stainless steel substrates/stents through a

precursor mediated hydrothermal processing. TiO2 nanoleafy structures on SS

showed improved biocompatibility in-vitro by promoting a rapid endothelialization

at 7 days, with elevated nitric oxide release and substantially lesser SMC viability in

comparison to bare SS. Hydrothermally nanostructured TiO2 layer on 316L SS also

exhibited better surface mechanical properties such as elasticity and improved

corrosion resistance, thereby minimizing the leaching of Ni and Cr ions. These


properties were critical for the development of a ceramic TiO2 coating on SS stents

which was sufficiently thin, and imparted better mechanical stability to resist any

delamination/ cracking during stent expansion and crimping. A successful translation

of generating a nanostructured titania coating on bare metal coronary SS stents was

demonstrated. This surface modification approach elicited good uniformity, crack

resistance and radial stiffness, all of which are crucial properties for coronary

stenting applications. However, a preliminary durability testing carried out in an

accelerated flow system under constant high shear stress resulted in slight leaching

of Ti ions with time. Further repeated trials under pulsatile flow conditions are

necessary to confirm the durability of the coatings on SS. Nonetheless, such

biocompatible titania nanostructures on SS could be a potential surface modification

strategy that does not utilize any polymers or drugs, to combat the problems of in-

stent restenosis and thrombosis, which are commonly associated with clinical bare

metal coronary stents.

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