a term paper on surface modification techniques used in the biomedical sector is presented. Various surface engineering techniques are described.

1. Surface modification techniques in
biomedical sector
Munesh Matwa (10MT30012) pages 7-9
Shashank Kumar Singh(10MT30021) pages 10-13
Sumeet Khanna (10MT30030) pages 1-6

2. 1
1. Introduction
A biomaterial is any matter, surface, or construct that interacts with biological systems. Biomaterials
can be derived either from nature or synthesized in the laboratory. There are varieties of biomaterials
such as metallic, ceramics and polymers that have been used as biomedical devices. Metallic
biomaterials can be grouped as steels, cobalt and titanium based alloys. Among non-metallic
polymeric based biomaterials are polyethylene terephthalate, polytetrafluoroethylene, ultrahigh
molecular weight polyethylene (UHMWPE) and lactide-co-glycolide. While titania (TiO2), titanium
carbide (TiC), titanium nitride (TiC), bioglass, hydroxyapatite (HA), silicon carbide (SiC) are typical
examples of ceramic biomaterials.
1.1 Properties of biomaterials
Biomaterials are implanted inside the body of an individual. This places strict specifications regarding
the reactivity, stability, and compatibility of biomaterials. Bio-integration is the ideal outcome
expected of an artificial implant. This implies that the phenomena that occur at the interface between
the implant and host tissues do not induce any deleterious effects such as chronic inflammatory
response or formation of unusual tissues. Hence, it is of paramount importance to design biomaterials
used in implants with the best surface properties. Meanwhile, these biomaterials must possess bulk
properties that meet other requirements, especially mechanical properties in order to function properly
in a bio-environment. As it is quite difficult to design biomaterials fulfilling both needs, a common
approach is to fabricate biomaterials with adequate bulk properties followed by a special treatment to
enhance the surface properties. In this way, it allows one to make ideal biomaterials with surface
attributes that are decoupled from the bulk properties. Furthermore, the surface properties can be
selectively modified to enhance the performance of the biomaterials. Hence, surface modification of
biomaterials is becoming an increasingly popular method to improve device multifunctionality,
tribological and mechanical properties, as well as biocompatibility of artificial devices. It has become
one of the key methods in biomaterials engineering.
1.2 Objectives of surface modification of biomaterials:
The main reasons to carry out various surface modification processes on implant materials for
biomedical applications can be summarized as follows:
1. Clean implant material surface from contaminations prior to implantation
2. Increase bioactivity, cell growth and tissue attachments after implantation
3. Increase hardness of implant to reduce wear rate especially in articulation joint
applications
4. Introduce passive layer to prevent excessive ion release into body environment
5. Promote antibacterial effect
6. Increase fatigue strength of implants
2 Biomaterial applications
Biomaterials have a very wide range of applications and their uses are ever increasing. Some of the
most common examples are described below.

3. 2
2.1 Hard tissue replacements
Hard tissues are often damaged due to accidents, aging, and other causes are widely used as hard
tissue replacements in artificial bones, joints, and dental
implants[1].
Fig.1: hip and knee joint implant replacement. Fig.2: Schematic diagram of the screw-
shaped artificial tooth.
2.1.1 Cardiac and cardiovascular applications
Cardiovascular applications are prosthetic heart valves, protective cases in pacemakers, artificial
hearts, and circulatory devices. Recently, the use of shape memory nickel–titanium(NITINOL) alloy in
intravascular devices, such as stents and occlusion coils has received considerable attention. The
common design of an artificial heart valve is shown in Fig.3. The ring and struts are made of titanium
or titanium alloys while the disk is made of pyrolytic carbon. Around the ring is a sewing ring made
of knitted Teflon cloth where the sutures anchoring the prosthesis to the heart are placed. The
metals in the prosthetic heart valves are often coated with a thin carbon film to enhance blood
compatibility. At present, stents ( Fig.4) are commonly used in the treatment of cardiovascular
disease. They dilate and keep narrowed blood vessels open. Nickel–titanium alloy is one of the most
common materials used in vascular stents due to its special shape memory effects.
Fig.3: Artificial heart valve Fig.4: Artificial vascular stents

4. 3
2.1.2 Osteosynthesis
Early and full restoration can be achieved by osteosynthesis, a method of treating the bone fracture
by surgical means. Typical implants for osteosynthesis include bone screws, bone plates (Fig. 5).
Bone plates are applied to almost all skeletal areas mostly as bridging devices and even as internal
fixators.
Fig.5: Bone screw and bone plate
3 Surface modification processes
Most of the surface modification techniques used in materials for construction, aerospace, and
energy sectors find application in biomaterials. A select few of the methods have been discussed in
this report.
3.1 Electron Beam Deposition of Ti on Co-Cr alloy implant material
Co-Cr alloys, stainless steel, and Ti-based alloys are most widely used for hard tissue replacements.
Co-Cr alloys and stainless steel have the advantage of high mechanical properties compared with Ti-
based alloys [2]. However, their relatively low biocompatibility is one of the major problems limiting
their wider application. On the other hand, Ti-based alloys are widely used because of their superior
biocompatibility and corrosion resistance.
3.1.1 Process
Co-Cr alloy substrates with dimensions of 10 × 10 × 2 mm3
were prepared. As a target material,
commercially pure grade Ti plates are used. A Ti layer was coated on the Co-Cr substrate at a rate of
about 0.1nm/s up to a thickness of 10μm. During the coating process, the substrate holder was
rotated at 20 rpm to achieve a uniform thickness.

5. 4
3.1.2 Microstructure
Fig.6: Co-Cr substrate Fig.7: Ti coating on Co-Cr substrate by electron
beam deposition
3.2 Micro-Arc Oxidation
The biocompatibility of metallic implant materials is strongly dependant on the surface oxide
layer[2]. The excellent biocompatibility of Ti and its alloys is mostly attributed to the thin TiO2 layer
that is formed spontaneously in ambient air. In addition, the biocompatibility of Ti can be further
improved by modifying the structure, morphology, and composition of the oxide layer.
3.2.1 Process
Fig.8: Coating process (A) as-machined
Co-Cr
(B) Ti coated Co-Cr (C) MAO treated after Ti coating
After the Ti coating, the specimens were oxidized electrochemically by the MAO process. The MAO
treatments of the prepared specimens were carried out in an aqueous electrolyte containing Ca and
P sources by applying a pulsed DC field. As an electrolyte, a solution of 0.15M calcium acetate
monohydrate and 0.02M glycerol phosphate calcium salt was used. The frequency and duty of the
pulsed DC power were 660 Hz and 60%, respectively.
3.2.2 Microstructure
Fig.9: TiO2 by MAO on Ti coated on Co-Cr substrate

6. 5
Fig.10: SEM cross-sectional images with (A)
secondary electron
(B) back scattered electron images of MAO
treated specimen.
3.3 Ion implantation process [3]
Ion implantation or ion beam processing is a procedure in which ions of a material are accelerated in
an electric field and bombarded into the solid substrate surface [1]. Various ions such as oxygen,
nitrogen, carbon, etc. can be implanted on any substrate material for a coating purpose to modify
the substrate surface. The energy of the ions before they penetrate the surface is typically between
20 and 200 keV. Up to date, ion implantation has been successfully applied in metal and polymer
biomaterials.
Two common types of ion implantation process are:
(i) Conventional beam line ion implantation
(ii) Plasma immersion ion implantation (PIII) method.
The basic difference between the beam line ion implantation and plasma immersion ion
implantation method is the target function. In beam line ion implantation, the target is totally
isolated from the ion beam generation. In contrast, the target is an active part of the ion generation
through bias voltage in PIII system.
Schematic diagram of (a) beam line ion implantation system and (b) PIII ion implantation system
By means of the pumping system, an atmosphere of a working gas at a reduced pressure is created,
typically in the range of 10-3-10-’ Pa. From this gas, plasma is generated. The work piece is pulse-
biased to a negative high voltage that ranges typically from several kilovolts up to 100 kV.

7. 6
Schematic presentation of an apparatus for plasma
immersion ion implantation
Schematic drawing of the polyfunctional IBAD system
3.4 Ion-beam assisted deposition
Ion-beam assisted deposition (IBAD) is a vacuum deposition process that combines physical vapor
deposition (PVD) with ion-beam bombardment [3]. Above is a schematic drawing of a polyfunctional
IBAD system. The major feature of IBAD is bombardment with a certain energy (ranging from several
hundred to several thousand eV) ion beam during the deposition of coating. The major processing
parameters are coating materials, evaporation rate or sputtering rate, ion species, ion energy and
ion beam current density.
 Biocoating with higher adhesive strength than traditional coating techniques.
 Low substrate temperature.
Applications:
 Hydroxyapatite coating preparation.
 DLC film and C–N film: chemically inert, extreme high hardness and low friction
coefficient .
 Silver coating is that it prevents bacteria attachment to the biomaterial surface.
3.5 Gas Nitriding
3.5.1 Process
Schematic drawing of gas nitriding process.

8. 7
The gas nitriding process (NP) is schematically shown above. A biomedical, β-type titanium alloy, Ti–
29Nb–13Ta–4.6Zr (TNTZ) is taken. First, the specimen is set in a furnace equipped with a chamber.
Then, by using a rotary vacuum pump, the atmosphere in the chamber is exchanged three times
from air to nitrogen; the purity of nitrogen gas is higher than 99.9995%. Subsequently, the furnace is
heated up to 1023, 1073, 1123, or 1223 K at a reduced pressure below 0.001 MPa. When the
temperature in the furnace reached the desired temperature, the nitrogen gas is introduced into the
chamber at a pressure of up to 0.100 MPa and the recording of the nitriding time is started. The
specimen is kept in the nitrogen atmosphere at each temperature for 21.6 ks. After this duration,
the furnace is cooled down to the room temperature while maintaining the nitrogen atmosphere in
the chamber.
A biomedical, β-type titanium alloy, Ti–29Nb–13Ta–4.6Zr (TNTZ), has been newly developed in order
to achieve a lower Young's modulus similar to that of human hard tissues in addition to excellent
mechanical properties and good corrosion resistance for use as structural biomaterial [1], [2], [3],
[4] and [5]. However, the wear resistance of titanium alloys is generally inferior to that of other
metallic biomaterials [6]. When the titanium alloys are utilized as a material for artificial hip joints,
bone plates, etc., one of the possible risks due to wear includes the loosening of these tools. Thus,
the improvement of wear resistance is required for biomedical titanium alloys.
3.5.2 Effect of Al on diffusion rate of O in TiO2 in gas nitriding
In general, ions can diffuse via the point defect in ionic crystals such as oxides. Depending on the
oxygen partial pressure and ambient temperature, the main point defect in TiO2 are probably oxygen
vacancies under the experimental conditions in this study. Therefore, the inward diffusion of O is
dominant in TiO2. A Ti atom is present as Ti4+
in TiO2, while Al3+
is the stable state of an Al atom.
According to the point defect theory, when Ti4+
is substituted with Al3+
, an oxygen vacancy is
generated in order to satisfy the electroneutrality as follows [5]:
equation(3)
Al2O3=2Al′
Ti+VO +3OO, …………………..(3)
Al′
Ti is Al3+
substituted with Ti4+
in TiO2, VO is an oxygen vacancy in TiO2, OO is the O2−
occupying
intrinsic site of O2−
in TiO2 ( is the positive effective charge and ′ is the negative effective charge).
Optical micrographs obtained
from cross-sections of TNTZ
subjected to gas nitriding

9. 8
Eq.(3) indicates that the dissolution of Al3+
into TiO2 may increase the number of oxygen vacancies in
it. As a result, it is expected to increase the diffusion rate of O ions. In fact, the experimental results
of AES (Fig. 8 and Fig. 9) show that O is present at a greater depth in Ti641073NP than in TNTZ1073NP.
This result may be derived from the enhancement in O diffusion due to Al in TiO2.
Fig. 8.
Compositional depth profiles obtained from surfaces of (a) TNTZST and (b) TNTZ1073NP by AES.
Fig. 9.
Compositional depth profiles obtained from surfaces of (a) Ti64ST and (b) Ti641073NP by AES
3.6 Vacuum plasma sprayed
VPS, Fig. 2, spraying techniques were used to manufacture HA coatings with an approximate
thickness of 40 lm. The VPS coatings were sprayed onto Ti—6Al—4V strips (80]20]2 mm) which had
been previously grit blasted with alumina grit and coated with a pure titanium bond layer. The
particles used in the manufacture of the VPS coatings were angular in nature and had a wide size
distribution. There is a linear increase of pushout failure load with increasing surface roughness for
VPS HA coatings. The coatings had a comparably moderate roughness which should encourage
coating dissolution due to the large surface area exposed to the body’s environment and allow good

10. 9
mechanical interlocking with bone, without impairing the mechanical strength of the surface. It was
found that there is higher crystallinity and lower residual stress in the VPS coatings, which will result
in a slow rate of dissolution in vitro and in vivo relative to the DGUN coatings.
3.7 Detonation Gun (DGUN)
DGUN, Fig. 3, spraying techniques were used to manufacture HA coatings with an approximate
thickness of 40 lm. The DGUN coatings were sprayed directly onto grit blasted Ti—6Al—4V with no
intervening bond layer. Coatings were applied to one side of the substrate only. The higher
temperature which the powder particles reached during detonation spraying should impose a higher
degree of melting on the starting powder creating a more amorphous coating. This should allow
better adhesion of the hydroxyapatite to the substrate. The higher velocity, higher energy DGUN
process imposed a greater degree of melting on the powder but this process may be
counterbalanced by the extremely short dwell time of the particles in the plasma. These effects
should combine to produce a more amorphous, more dense coating which, despite being better
adhered to the substrate will undergo a more rapid dissolution in vitro than the VPS coatings. The
detonation process resulted, to some degree, in the degradation of pure hydroxyapatite to beta-
tricalcium phosphate in the final coating.
The DGUN spray process is a higher temperature, higher velocity technique which is thought to
impose a higher degree of melting on the ceramic starting powder. The DGUN process producing a
denser coating which had a higher proportion of the amorphous phase with some evidence for the
appearance of beta-tricalcium phosphate.
There was a considerable difference in the crystallinity of the VPS and DGUN coating types.

11. 10
3.8 Laser
Laser carburizing process involves carbon diffusion into the metal substrate using laser
irradiation[6]. The typical source of carbon is graphite powder. This process involved heating of
specimen through continuous or pulse wave laser irradiation, rapid melting, intermixing or diffusion
of carbon particle, and rapid solidification of the pre-deposited alloying elements on substrate to
form an alloyed zone or carburized layer.
Lasers can rapidly and specifically induce surface changes in organic and inorganic materials.
The advantages of using lasers for such modification are the precise control of the frequency of the
light, the wide range of frequencies available, the high energy density, the ability to focus and faster
the light, the possibilities for using both heat and specific excitation to effect change and the ability
to pulse the source and control reaction time. Treatments are pulsed (100 nanoseconds to
picoseconds pulse times) and continuous wave (CW), with interaction times often less than 1
microsecond. Laser-induced surface alterations include annealing, etching, deposition and
polymerisation.
The three requirements generally expected of biomaterials coating are:
1) crystallinity,
2) porosity
3) Adhesion.
Common advantages of laser surfacing compared to alternatives are :
● chemical cleanliness
● controlled thermal penetration and, therefore, distortion
● controlled thermal profile and, therefore, shape and location of the heat
● affected region
● less after-machining, if any, is required
● remote non-contact processing is usually possible
● relatively easy to automate
At present, the lasers are being used in the following surface modifications of the biomaterials:
3.8.1 Laser patterning and microfabrication
Laser patterning is based on the possibility of focusing an intense laser beam at certain spots on a
surface, where the high beam intensity causes evaporation of the material. By this approach, pits
can be produced down to 1 mm, in the size range of interest to match cell sizes.

12. 11
3.8.2 Pulsed laser deposition (PLD) of biocompatible ceramics
Pulsed laser deposition (PLD) is a new technique for the deposition of thin films of biocompatible
ceramics . Pulsed laser deposition is especially well suited to the deposition of bone-like ceramics
(e.g. hydroxyapatite (HA) and calcium phosphates) on to metal, ceramic, semiconductor or polymer
substrates for potential application in medical implants, prosthetic devices and biocompatible
probes or sensors.
3.8.3 Laser surface treatment for improving corrosion
Laser treatment improvement resistance is explained by a combination of the homogenisation of the
surface by melting, the hardening due to N incorporation and the thickening of the oxide layer.
Moreover, excimer laser surface treatment in air showed a remarkable improvement in pitting
corrosion resistance for 316LS biograde stainless steel. The results show that excimer laser surface
melting can effectively eliminate carbides and second phases alike, while also serving the function of
homogenising the microstructure. N2 induced into the laser treated surface could promote new
precipitates and as a result lowered the corrosion resistance of 316LS stainless steel and Ti–6Al–4V
alloy.
3.8.4 Laser grafting
For the purpose of improved surface hydrophilicity and biocompatibility of ethylene–propylene
rubber, 2-hydroxyethyl methacrylate (HEMA) and N-vinyl pyrrolidone (NVP) have been grafted on to
the surface of this polymer using a CO2 pulsed laser at different fluence (output power J/cm2)
as the excitation source.
3.9 High Velocity Oxygen Fuel (HVOF) coating
High Velocity Oxygen Fuel (HVOF) coating is a thermal spray coating process used to improve or
restore a component’s surface properties or dimensions, thus extending equipment life by
significantly increasing erosion and wear resistance, and corrosion protection.
As with all thermal spray coating processes, HVOF coating material is heated and accelerated by a
gas stream to a component’s surface, to attain better properties. With the HVOF coating process,
the gas stream is produced by mixing and igniting oxygen and fuel (gas or liquid) in a combustion
chamber and allowing the high pressure gas to accelerate through a nozzle. Powder is introduced
into this stream where it is heated and accelerated towards a component’s surface. The resulting
thermal spray coating consists of thin overlapping platelets
The application of hydroxyapatite (HA) coatings on Ti-6Al–4V based prosthetics has been widely
used due to the unique biocompatibility of HA. For a long term usage, an HA coating must exhibit a
high biocompatibility and adequate mechanical properties, such as a high bond strength and an

13. 12
elastic modulus value close to that of the bone. The biocompatibility and the mechanical properties
will depend on the coating microstructure, crystallinity and phase composition.
Cross-section of the HA coating after a 7-day incubation in the SBF solution.
3.10 Sputtering
A technique used to deposit thin films of a material onto a surface (substrate)[7]. By first creating a
gaseous plasma and then accelerating the ions from this plasma into some source material (target),
the source material is eroded by the arriving ions via energy transfer and is ejected in the form of
neutral particles - either individual atoms, clusters of atoms or molecules. As these neutral particles
are ejected they will travel in a straight line unless they come into contact with something - other
particles or a nearby surface. If a "substrate" such as a Si wafer is placed in the path of these ejected
particles it will be coated by a thin film of the source material
Schematic presentation of an apparatus for Sputtering
Calcium ion implantation where calcium ions are implanted into biomedical titanium alloys, calcium
ion mixing method where Ca is sputtered on the surface of biomedical titanium alloys followed by
Ar ion implantation, etc. CaP precipitation is enhanced on the surface of biomedical titanium alloys
conducted with these treatments when they are implanted into living body.

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4 Conclusion
● We have described the current applications of surface modification techniques in the field of
biomaterials and bioengineering.
● It is observed that the overall trends of surface modification methods has shifted from the
use of conventional source (chemical, induction heater and gas) to the application of
advanced technology (electrolyte based, laser, plasma and ion).
● The works on surface modifications has expanded from focusing on tribological issues such
as wear resistance, corrosion resistance and hardness of modified layer to clinical issues
such as cell growth, cell attachment and antibacterial effects.
5 References
[1] I. S., "Surface Modification Techniques for Biomedical Grade of Titanium Alloys: Oxidation,
Carburization and Ion Implantation Processes," [Online]. Available: www.intechopen.com.
[2] M. H. Cheol, "Enhanced Biocompatibility of Co-Cr Implant Material by Ti Coating and Micro-Arc
Oxidation," Wiley InterScience.
[3] F. Z. Cui , "Biomaterials modification by ion-beam processing," Surface and Coatings Technology
112 (1999), vol. 112, p. 278–285, 1999.
[4] E. W, "Modification of mechanical and chemical surface properties of metals by plasma immersion
ion implantation," Surface and Coatings Technology, Vols. 100-101, pp. 341-352, 1998.
[5] M. Nakai, "Surface hardening of biomedical Ti–29Nb–13Ta–4.6Zr and Ti–6Al–4V ELI by gas
nitriding," Materials Science and Engineering: A, vol. 486, no. 1–2, p. 193–201, 15 July 2008.
6 [6] R.S. Lima, K.A. Khor, H. Li, P. Cheang B.R. Marple, HVOF spraying of nanostructured
hydroxyapatite for biomedical applications, Materials Science and Engineering: A Volume 396,
Issues 1–2, 15 April 2005, Pages 181–187.
7 [7] Mitsuo Niinomi “Recent research and development in titanium alloys for biomedical
applications and healthcare goods”, Science and Technology of Advanced Materials Volume 4,
Issue 5, September 2003, Pages 445–454.