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

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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.

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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

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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.

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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

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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.

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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.

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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

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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

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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.

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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.

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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

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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.